Nanoparticulate
Drug Delivery Systems
DRUGS AND THE PHARMACEUTICAL SCIENCES
A Series of Textbooks and Monographs
Executive Editor
James Swarbrick
PharmaceuTech, Inc.
Pinehurst, North Carolina
Advisory Board
Larry L. Augsburger
Harry G. Brittain
University of Maryland
Baltimore, Maryland
Center for Pharmaceutical Physics
Milford, New Jersey
Jennifer B. Dressman
Johann Wolfgang Goethe University
Frankfurt, Germany
Anthony J. Hickey
University of North Carolina School of
Pharmacy
Chapel Hill, North Carolina
Jeffrey A. Hughes
University of Florida College of
Pharmacy
Gainesville, Florida
Trevor M. Jones
The Association of the
British Pharmaceutical Industry
London, United Kingdom
Vincent H. L. Lee
Ajaz Hussain
Sandoz
Princeton, New Jersey
Hans E. Junginger
Leiden/Amsterdam Center
for Drug Research
Leiden, The Netherlands
Stephen G. Schulman
University of Southern California
Los Angeles, California
University of Florida
Gainesville, Florida
Jerome P. Skelly
Elizabeth M. Topp
Alexandria, Virginia
University of Kansas School of
Pharmacy
Lawrence, Kansas
Geoffrey T. Tucker
University of Sheffield
Royal Hallamshire Hospital
Sheffield, United Kingdom
Peter York
University of Bradford School of
Pharmacy
Bradford, United Kingdom
1. Pharmacokinetics, Milo Gibaldi and Donald Perrier
2. Good Manufacturing Practices for Pharmaceuticals: A Plan for Total
Quality Control, Sidney H. Willig, Murray M. Tuckerman,
and William S. Hitchings IV
3. Microencapsulation, edited by J. R. Nixon
4. Drug Metabolism: Chemical and Biochemical Aspects, Bernard Testa
and Peter Jenner
5. New Drugs: Discovery and Development, edited by Alan A. Rubin
6. Sustained and Controlled Release Drug Delivery Systems, edited by
Joseph R. Robinson
7. Modern Pharmaceutics, edited by Gilbert S. Banker
and Christopher T. Rhodes
8. Prescription Drugs in Short Supply: Case Histories, Michael A. Schwartz
9. Activated Charcoal: Antidotal and Other Medical Uses, David O. Cooney
10. Concepts in Drug Metabolism (in two parts), edited by Peter Jenner
and Bernard Testa
11. Pharmaceutical Analysis: Modern Methods (in two parts), edited by
James W. Munson
12. Techniques of Solubilization of Drugs, edited by Samuel H. Yalkowsky
13. Orphan Drugs, edited by Fred E. Karch
14. Novel Drug Delivery Systems: Fundamentals, Developmental Concepts,
Biomedical Assessments, Yie W. Chien
15. Pharmacokinetics: Second Edition, Revised and Expanded, Milo Gibaldi
and Donald Perrier
16. Good Manufacturing Practices for Pharmaceuticals: A Plan for Total
Quality Control, Second Edition, Revised and Expanded, Sidney H. Willig,
Murray M. Tuckerman, and William S. Hitchings IV
17. Formulation of Veterinary Dosage Forms, edited by Jack Blodinger
18. Dermatological Formulations: Percutaneous Absorption, Brian W. Barry
19. The Clinical Research Process in the Pharmaceutical Industry, edited by
Gary M. Matoren
20. Microencapsulation and Related Drug Processes, Patrick B. Deasy
21. Drugs and Nutrients: The Interactive Effects, edited by Daphne A. Roe
and T. Colin Campbell
22. Biotechnology of Industrial Antibiotics, Erick J. Vandamme
23. Pharmaceutical Process Validation, edited by Bernard T. Loftus
and Robert A. Nash
24. Anticancer and Interferon Agents: Synthesis and Properties, edited by
Raphael M. Ottenbrite and George B. Butler
25. Pharmaceutical Statistics: Practical and Clinical Applications,
Sanford Bolton
26. Drug Dynamics for Analytical, Clinical, and Biological Chemists,
Benjamin J. Gudzinowicz, Burrows T. Younkin, Jr.,
and Michael J. Gudzinowicz
27. Modern Analysis of Antibiotics, edited by Adjoran Aszalos
28. Solubility and Related Properties, Kenneth C. James
29. Controlled Drug Delivery: Fundamentals and Applications, Second
Edition,
Revised and Expanded, edited by Joseph R. Robinson and Vincent H.
Lee
30. New Drug Approval Process: Clinical and Regulatory Management,
edited by Richard A. Guarino
31. Transdermal Controlled Systemic Medications, edited by Yie W. Chien
32. Drug Delivery Devices: Fundamentals and Applications, edited by
Praveen Tyle
33. Pharmacokinetics: Regulatory • Industrial • Academic Perspectives,
edited by Peter G. Welling and Francis L. S. Tse
34. Clinical Drug Trials and Tribulations, edited by Allen E. Cato
35. Transdermal Drug Delivery: Developmental Issues and Research
Initiatives, edited by Jonathan Hadgraft and Richard H. Guy
36. Aqueous Polymeric Coatings for Pharmaceutical Dosage Forms,
edited by James W. McGinity
37. Pharmaceutical Pelletization Technology, edited by Isaac GhebreSellassie
38. Good Laboratory Practice Regulations, edited by Allen F. Hirsch
39. Nasal Systemic Drug Delivery, Yie W. Chien, Kenneth S. E. Su,
and Shyi-Feu Chang
40. Modern Pharmaceutics: Second Edition, Revised and Expanded,
edited by Gilbert S. Banker and Christopher T. Rhodes
41. Specialized Drug Delivery Systems: Manufacturing and Production
Technology, edited by Praveen Tyle
42. Topical Drug Delivery Formulations, edited by David W. Osborne
and Anton H. Amann
43. Drug Stability: Principles and Practices, Jens T. Carstensen
44. Pharmaceutical Statistics: Practical and Clinical Applications,
Second Edition, Revised and Expanded, Sanford Bolton
45. Biodegradable Polymers as Drug Delivery Systems, edited by
Mark Chasin and Robert Langer
46. Preclinical Drug Disposition: A Laboratory Handbook, Francis L. S. Tse
and James J. Jaffe
47. HPLC in the Pharmaceutical Industry, edited by Godwin W. Fong
and Stanley K. Lam
48. Pharmaceutical Bioequivalence, edited by Peter G. Welling,
Francis L. S. Tse, and Shrikant V. Dinghe
49. Pharmaceutical Dissolution Testing, Umesh V. Banakar
50. Novel Drug Delivery Systems: Second Edition, Revised and Expanded,
Yie W. Chien
51. Managing the Clinical Drug Development Process, David M. Cocchetto
and Ronald V. Nardi
52. Good Manufacturing Practices for Pharmaceuticals: A Plan for Total
Quality Control, Third Edition, edited by Sidney H. Willig
and James R. Stoker
53. Prodrugs: Topical and Ocular Drug Delivery, edited by Kenneth B. Sloan
54. Pharmaceutical Inhalation Aerosol Technology, edited by
Anthony J. Hickey
55. Radiopharmaceuticals: Chemistry and Pharmacology, edited by
Adrian D. Nunn
56. New Drug Approval Process: Second Edition, Revised and Expanded,
edited by Richard A. Guarino
57. Pharmaceutical Process Validation: Second Edition, Revised
and Expanded, edited by Ira R. Berry and Robert A. Nash
58. Ophthalmic Drug Delivery Systems, edited by Ashim K. Mitra
59. Pharmaceutical Skin Penetration Enhancement, edited by
Kenneth A. Walters and Jonathan Hadgraft
60. Colonic Drug Absorption and Metabolism, edited by Peter R. Bieck
61. Pharmaceutical Particulate Carriers: Therapeutic Applications, edited by
Alain Rolland
62. Drug Permeation Enhancement: Theory and Applications, edited by
Dean S. Hsieh
63. Glycopeptide Antibiotics, edited by Ramakrishnan Nagarajan
64. Achieving Sterility in Medical and Pharmaceutical Products, Nigel A. Halls
65. Multiparticulate Oral Drug Delivery, edited by Isaac Ghebre-Sellassie
66. Colloidal Drug Delivery Systems, edited by Jörg Kreuter
67. Pharmacokinetics: Regulatory • Industrial • Academic Perspectives,
Second Edition, edited by Peter G. Welling and Francis L. S. Tse
68. Drug Stability: Principles and Practices, Second Edition, Revised
and Expanded, Jens T. Carstensen
69. Good Laboratory Practice Regulations: Second Edition, Revised
and Expanded, edited by Sandy Weinberg
70. Physical Characterization of Pharmaceutical Solids, edited by
Harry G. Brittain
71. Pharmaceutical Powder Compaction Technology, edited by
Göran Alderborn and Christer Nyström
72. Modern Pharmaceutics: Third Edition, Revised and Expanded, edited by
Gilbert S. Banker and Christopher T. Rhodes
73. Microencapsulation: Methods and Industrial Applications, edited by
Simon Benita
74. Oral Mucosal Drug Delivery, edited by Michael J. Rathbone
75. Clinical Research in Pharmaceutical Development, edited by Barry Bleidt
and Michael Montagne
76. The Drug Development Process: Increasing Efficiency and Cost
Effectiveness, edited by Peter G. Welling, Louis Lasagna,
and Umesh V. Banakar
77. Microparticulate Systems for the Delivery of Proteins and Vaccines,
edited by Smadar Cohen and Howard Bernstein
78. Good Manufacturing Practices for Pharmaceuticals: A Plan for Total
Quality Control, Fourth Edition, Revised and Expanded, Sidney H. Willig
and James R. Stoker
79. Aqueous Polymeric Coatings for Pharmaceutical Dosage Forms:
Second Edition, Revised and Expanded, edited by James W. McGinity
80. Pharmaceutical Statistics: Practical and Clinical Applications,
Third Edition, Sanford Bolton
81. Handbook of Pharmaceutical Granulation Technology, edited by
Dilip M. Parikh
82. Biotechnology of Antibiotics: Second Edition, Revised and Expanded,
edited by William R. Strohl
83. Mechanisms of Transdermal Drug Delivery, edited by Russell O. Potts
and Richard H. Guy
84. Pharmaceutical Enzymes, edited by Albert Lauwers and Simon Scharpé
85. Development of Biopharmaceutical Parenteral Dosage Forms, edited by
John A. Bontempo
86. Pharmaceutical Project Management, edited by Tony Kennedy
87. Drug Products for Clinical Trials: An International Guide to Formulation •
Production • Quality Control, edited by Donald C. Monkhouse
and Christopher T. Rhodes
88. Development and Formulation of Veterinary Dosage Forms:
Second Edition, Revised and Expanded, edited by Gregory E. Hardee
and J. Desmond Baggot
89. Receptor-Based Drug Design, edited by Paul Leff
90. Automation and Validation of Information in Pharmaceutical Processing,
edited by Joseph F. deSpautz
91. Dermal Absorption and Toxicity Assessment, edited by Michael S.
Roberts
and Kenneth A. Walters
92. Pharmaceutical Experimental Design, Gareth A. Lewis, Didier Mathieu,
and Roger Phan-Tan-Luu
93. Preparing for FDA Pre-Approval Inspections, edited by Martin D. Hynes III
94. Pharmaceutical Excipients: Characterization by IR, Raman, and NMR
Spectroscopy, David E. Bugay and W. Paul Findlay
95. Polymorphism in Pharmaceutical Solids, edited by Harry G. Brittain
96. Freeze-Drying/Lyophilization of Pharmaceutical and Biological Products,
edited by Louis Rey and Joan C. May
97. Percutaneous Absorption: Drugs–Cosmetics–Mechanisms–Methodology,
Third Edition, Revised and Expanded, edited by Robert L. Bronaugh
and Howard I. Maibach
98. Bioadhesive Drug Delivery Systems: Fundamentals, Novel Approaches,
and Development, edited by Edith Mathiowitz, Donald E. Chickering III,
and Claus-Michael Lehr
99. Protein Formulation and Delivery, edited by Eugene J. McNally
100. New Drug Approval Process: Third Edition, The Global Challenge,
edited by Richard A. Guarino
101. Peptide and Protein Drug Analysis, edited by Ronald E. Reid
102. Transport Processes in Pharmaceutical Systems, edited by
Gordon L. Amidon, Ping I. Lee, and Elizabeth M. Topp
103. Excipient Toxicity and Safety, edited by Myra L. Weiner
and Lois A. Kotkoskie
104. The Clinical Audit in Pharmaceutical Development, edited by
Michael R. Hamrell
105. Pharmaceutical Emulsions and Suspensions, edited by Francoise
Nielloud
and Gilberte Marti-Mestres
106. Oral Drug Absorption: Prediction and Assessment, edited by
Jennifer B. Dressman and Hans Lennernäs
107. Drug Stability: Principles and Practices, Third Edition, Revised
and Expanded, edited by Jens T. Carstensen and C. T. Rhodes
108. Containment in the Pharmaceutical Industry, edited by James P. Wood
109. Good Manufacturing Practices for Pharmaceuticals: A Plan for Total
Quality Control from Manufacturer to Consumer, Fifth Edition, Revised
and Expanded, Sidney H. Willig
110. Advanced Pharmaceutical Solids, Jens T. Carstensen
111. Endotoxins: Pyrogens, LAL Testing, and Depyrogenation, Second Edition,
Revised and Expanded, Kevin L. Williams
112. Pharmaceutical Process Engineering, Anthony J. Hickey
and David Ganderton
113. Pharmacogenomics, edited by Werner Kalow, Urs A. Meyer
and Rachel F. Tyndale
114. Handbook of Drug Screening, edited by Ramakrishna Seethala
and Prabhavathi B. Fernandes
115. Drug Targeting Technology: Physical • Chemical • Biological Methods,
edited by Hans Schreier
116. Drug–Drug Interactions, edited by A. David Rodrigues
117. Handbook of Pharmaceutical Analysis, edited by Lena Ohannesian
and Anthony J. Streeter
118. Pharmaceutical Process Scale-Up, edited by Michael Levin
119. Dermatological and Transdermal Formulations, edited by
Kenneth A. Walters
120. Clinical Drug Trials and Tribulations: Second Edition, Revised
and Expanded, edited by Allen Cato, Lynda Sutton, and Allen Cato III
121. Modern Pharmaceutics: Fourth Edition, Revised and Expanded, edited by
Gilbert S. Banker and Christopher T. Rhodes
122. Surfactants and Polymers in Drug Delivery, Martin Malmsten
123. Transdermal Drug Delivery: Second Edition, Revised and Expanded,
edited by Richard H. Guy and Jonathan Hadgraft
124. Good Laboratory Practice Regulations: Second Edition, Revised
and Expanded, edited by Sandy Weinberg
125. Parenteral Quality Control: Sterility, Pyrogen, Particulate, and Package
Integrity Testing: Third Edition, Revised and Expanded, Michael J. Akers,
Daniel S. Larrimore, and Dana Morton Guazzo
126. Modified-Release Drug Delivery Technology, edited by
Michael J. Rathbone, Jonathan Hadgraft, and Michael S. Roberts
127. Simulation for Designing Clinical Trials: A PharmacokineticPharmacodynamic Modeling Perspective, edited by Hui C. Kimko
and Stephen B. Duffull
128. Affinity Capillary Electrophoresis in Pharmaceutics and Biopharmaceutics,
edited by Reinhard H. H. Neubert and Hans-Hermann Rüttinger
129. Pharmaceutical Process Validation: An International Third Edition,
Revised and Expanded, edited by Robert A. Nash and Alfred H. Wachter
130. Ophthalmic Drug Delivery Systems: Second Edition, Revised
and Expanded, edited by Ashim K. Mitra
131. Pharmaceutical Gene Delivery Systems, edited by Alain Rolland
and Sean M. Sullivan
132. Biomarkers in Clinical Drug Development, edited by John C. Bloom
and Robert A. Dean
133. Pharmaceutical Extrusion Technology, edited by Isaac Ghebre-Sellassie
and Charles Martin
134. Pharmaceutical Inhalation Aerosol Technology: Second Edition,
Revised and Expanded, edited by Anthony J. Hickey
135. Pharmaceutical Statistics: Practical and Clinical Applications,
Fourth Edition, Sanford Bolton and Charles Bon
136. Compliance Handbook for Pharmaceuticals, Medical Devices,
and Biologics, edited by Carmen Medina
137. Freeze-Drying/Lyophilization of Pharmaceutical and Biological Products:
Second Edition, Revised and Expanded, edited by Louis Rey
and Joan C. May
138. Supercritical Fluid Technology for Drug Product Development, edited by
Peter York, Uday B. Kompella, and Boris Y. Shekunov
139. New Drug Approval Process: Fourth Edition, Accelerating Global
Registrations, edited by Richard A. Guarino
140. Microbial Contamination Control in Parenteral Manufacturing, edited by
Kevin L. Williams
141. New Drug Development: Regulatory Paradigms for Clinical Pharmacology
and Biopharmaceutics, edited by Chandrahas G. Sahajwalla
142. Microbial Contamination Control in the Pharmaceutical Industry, edited
by Luis Jimenez
143. Generic Drug Product Development: Solid Oral Dosage Forms, edited by
Leon Shargel and Izzy Kanfer
144. Introduction to the Pharmaceutical Regulatory Process, edited by
Ira R. Berry
145. Drug Delivery to the Oral Cavity: Molecules to Market, edited by
Tapash K. Ghosh and William R. Pfister
146. Good Design Practices for GMP Pharmaceutical Facilities, edited by
Andrew Signore and Terry Jacobs
147. Drug Products for Clinical Trials, Second Edition, edited by Donald
Monkhouse, Charles Carney, and Jim Clark
148. Polymeric Drug Delivery Systems, edited by Glen S. Kwon
149. Injectable Dispersed Systems: Formulation, Processing, and
Performance,
edited by Diane J. Burgess
150. Laboratory Auditing for Quality and Regulatory Compliance,
Donald Singer, Raluca-Ioana Stefan, and Jacobus van Staden
151. Active Pharmaceutical Ingredients: Development, Manufacturing,
and Regulation, edited by Stanley Nusim
152. Preclinical Drug Development, edited by Mark C. Rogge and David R. Taft
153. Pharmaceutical Stress Testing: Predicting Drug Degradation, edited by
Steven W. Baertschi
154. Handbook of Pharmaceutical Granulation Technology: Second Edition,
edited by Dilip M. Parikh
155. Percutaneous Absorption: Drugs–Cosmetics–Mechanisms–Methodology,
Fourth Edition, edited by Robert L. Bronaugh and Howard I. Maibach
156. Pharmacogenomics: Second Edition, edited by Werner Kalow,
Urs A. Meyer and Rachel F. Tyndale
157. Pharmaceutical Process Scale-Up, Second Edition, edited by
Michael Levin
158. Microencapsulation: Methods and Industrial Applications, Second
Edition,
edited by Simon Benita
159. Nanoparticle Technology for Drug Delivery, edited by Ram B. Gupta
and Uday B. Kompella
160. Spectroscopy of Pharmaceutical Solids, edited by Harry G. Brittain
161. Dose Optimization in Drug Development, edited by Rajesh Krishna
162. Herbal Supplements-Drug Interactions: Scientific and Regulatory
Perspectives, edited by Y. W. Francis Lam, Shiew-Mei Huang,
and Stephen D. Hall
163. Pharmaceutical Photostability and Stabilization Technology, edited by
Joseph T. Piechocki and Karl Thoma
164. Environmental Monitoring for Cleanrooms and Controlled Environments,
edited by Anne Marie Dixon
165. Pharmaceutical Product Development: In Vitro-In Vivo Correlation, edited
by Dakshina Murthy Chilukuri, Gangadhar Sunkara, and David Young
166. Nanoparticulate Drug Delivery Systems, edited by Deepak Thassu,
Michel Deleers, and Yashwant Pathak
167. Endotoxins: Pyrogens, LAL Testing and Depyrogenation, Third Edition,
edited by Kevin L. Williams
168. Good Laboratory Practice Regulations, Fourth Edition, edited by
Sandy Weinberg
169. Good Manufacturing Practices for Pharmaceuticals, Sixth Edition,
edited by Joseph D. Nally
Nanoparticulate
Drug Delivery Systems
edited by
Deepak Thassu
UCB Pharma, Inc.
Rochester, New York, U.S.A.
Michel Deleers
UCB Pharma, Chemin du Foriest
Braine l'Alleud, Belgium
Yashwant Pathak
UCB Manufacturing, Inc.
Rochester, New York, U.S.A.
New York London
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Library of Congress Cataloging-in-Publication Data
Nanoparticulate drug delivery systems / edited by Deepak Thassu, Michel Deleers,
Yashwant Pathak.
p. ; cm. -- (Drugs and the pharmaceutical sciences ; v. 166)
Includes bibliographical references and index.
ISBN-13: 978-0-8493-9073-9 (alk. paper)
ISBN-10: 0-8493-9073-7 (alk. paper)
1. Drug delivery systems. 2. Nanoparticles. I. Thassu, Deepak. II. Deleers, Michel.
III. Pathak, Yashwant. IV. Series.
[DNLM: 1. Drug Delivery Systems--methods. 2. Nanostructures. 3. Drug Carriers.
W1 DR893B v.166 2007 / WB 340 N184 2007]
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2006051461
Foreword
The use of molecular or macromolecular entities and superstructures derived
thereof for the delivery of drugs has a long history. Antibodies, for instance, were
suggested early last century as a means to direct anticancer drugs to tumor cells in
the body expressing the corresponding antigen. Their use in the form of monoclonals
is now at the forefront of targeted therapy. Following advances in the discovery of
cell receptors, receptor-binding macromolecules were added to the armamentarium
of systems for the targeting of drugs. Parallel to these developments has been, since
the early 1970s, the exploitation of liposomes as a delivery system for drugs and
vaccines. These superstructures, formed spontaneously from amphipathic lipid
molecules, together with a diverse collection of other promising superstructures
derived from a huge variety of natural and synthetic monomeric or polymeric units,
have evolved to sophisticated versions through the incorporation onto their surface
of macromolecules that contribute to optimal pharmacokinetics of actives and their
delivery to where they are needed. An ever increasing number of drug- and vaccinedelivery systems are being tested clinically, with many already marketed.
Recently, drug-delivery systems have been rediscovered as the biological dimension of nanotechnology. A leading article in a prestigious scientific journal tells us
that “biologists are embracing nanotechnology—the engineering and manipulation
of entities in the 1 to 100 nm range—and are exploiting its potential to develop new
therapeutics and diagnostics.” What else is new?, you might say! Nonetheless, the
prefix nano (from the Greek word for dwarf ) is a useful one because it helps define
drug-delivery systems of a certain size range. Reflecting this trend of size definition,
Nanoparticulate Drug Delivery Systems is a worthy attempt to bring together a wide
range of drug-delivery systems for the delivery (targeted or otherwise), through a
variety of routes of administration, of drugs, diagnostics, and vaccines in the
treatment or prevention of disease, now encapsulated in the term “nanomedicine.”
Importantly, the book includes a wealth of the latest advances in the technology of
nanoparticulates, including electrospinning, formation of microcrystals, production
of liquid crystalline phases, and, last but not least, the technology of metallic
nanoparticles. The editors, Deepak Thassu, Michel Deleers, and Yashwant Pathak,
are to be complimented for both their judicial selection of nanosystems and choice
of the international panel of contributors.
Gregory Gregoriadis
The School of Pharmacy
University of London
London, U.K.
iii
Preface
For many decades, the interest in modifying drug-delivery systems has been a
prominent thrust of pharmaceutical research. In recent years, due to tremendous
expansion in the different scientific domains and skill sets, the scope has been widened to incorporate many faculties in the drug-delivery research covering physics,
polymer sciences, electrical engineering, bioelectronics, genetics, biotechnology,
and molecular pharmaceutics.
Pharmaceutical industry research culture is facing an uncertain future. Higher
clinical development cost coupled with declining drug-discovery process and lower
clinical success rates is decreasing the flow of new chemical entities in the research
and development pipeline.
Due to the advent of analytical techniques and capabilities to measure the
particle sizes in nanometer ranges, particulate drug-delivery systems research and
development has been moving from the micro- to the nanosize scale. Significant
research interests are geared towards utilizing the techniques where the particles
can be reduced almost to nanometer ranges, thus reducing the dose and reactive
nature of the molecule. This can deliver the drug at the targeted sites.
The book presented herewith is an attempt to describe the research efforts
being done in this direction by the global scientific community. Nanoparticulate
drug-delivery systems are a challenging area, and there are pulsating changes happening almost every day. This is an attempt to cover the recent trends and emerging
technologies in the area of nanoparticulate drug-delivery systems.
The first chapter covers a complete overview of the nanoparticulate drugdelivery system, covering wide applications and evaluation of the nanoparticulate
drug-delivery system in various fields. Chapter 2 encompasses formulations of
nanosuspensions for parenteral delivery. The third chapter covers the polymerbased nanoparticulate drug-delivery systems. Chapters 4 to 6 focus on nanofibers,
nanocrystals, and lipid-based nanoparticulate drug-delivery systems, respectively.
Chapters 7 to 10 discuss the engineering aspects and different techniques used
for nanoparticulate drug-delivery systems, including nanoengineering, aerosol flow
reactor, supercooled smectic nanoparticles, and metallic nanoparticles, respectively.
Chapters 11 and 12 focus on biological requirements and the role of nanobiotechnology in the development of nanomedicines. Chapters 13 to 21 extensively cover the
applications of nanoparticulate drug-delivery systems, including lipid nanoparticles for dermal applications; gene carriers for restenosis; ocular, central nervous
system, gastrointestinal applications; adjuvant for vaccine development; and transdermal systems.
It is our hope that this multiauthored book on nanoparticulate drug-delivery
systems will assist and enrich the readers in understanding the diverse types of
nanoparticulate drug-delivery systems available or under development, as well as
highlight their applications in the future development of nanomedicines. This book
is equally relevant to academic, industrial, as well as scientists working in pharmaceutical drug delivery worldwide. The text is planned in such a way that each
v
vi
Preface
chapter represents an independent area of research and can be easily followed without referring to other chapters.
We would like to express our sincere thanks to Tony Benfonte for the figures in
Chapters 1 and 13 and to Linda Glather for reading the manuscript and suggesting
corrections and punctuation. Special thanks to our editors, Stevan Zolo, Yvonne
Honigsberg, and Sherri Niziolek, who helped us to get through the project
successfully.
Last, but not least, we would like to express our sincere gratitude to all the
authors who have taken time from their busy schedules to be part of this project and
written wonderful chapters that added both the depth and value to this book.
Deepak Thassu
Michel Deleers
Yashwant Pathak
Contents
Foreword
Gregory Gregoriadis . . . . iii
Preface . . . . v
Contributors . . . . ix
1. Nanoparticulate Drug-Delivery Systems: An Overview 1
Deepak Thassu, Yashwant Pathak, and Michel Deleers
2. Nanosuspensions for Parenteral Delivery 33
Barrett E. Rabinow
3. Nanoparticles Prepared Using Natural and Synthetic Polymers 51
Sudhir S. Chakravarthi, Dennis H. Robinson, and Sinjan De
4. Nanofiber-Based Drug Delivery 61
Matthew D. Burke and Dmitry Luzhansky
5. Drug Nanocrystals—The Universal Formulation Approach for Poorly Soluble
Drugs 71
Jan Möschwitzer and Rainer H. Müller
6. Lipid-Based Nanoparticulate Drug Delivery Systems 89
Jun Wu, Xiaobin Zhao, and Robert J. Lee
7. Nanoengineering of Drug Delivery Systems 99
Ashwath Jayagopal and V. Prasad Shastri
8. Aerosol Flow Reactor Method for the Synthesis of Multicomponent Drug
Nano- and Microparticles 111
Janne Raula, Hannele Eerikäinen, Anna Lähde, and Esko I. Kauppinen
9. Supercooled Smectic Nanoparticles 129
Heike Bunjes and Judith Kuntsche
10. Biological and Engineering Considerations for Developing Tumor-Targeting
Metallic Nanoparticle Drug-Delivery Systems 141
Giulio F. Paciotti and Lawrence Tamarkin
11. Biological Requirements for Nanotherapeutic Applications
Joseph F. Chiang
159
12. Role of Nanobiotechnology in the Development of Nanomedicine
K. K. Jain
vii
173
viii
Contents
13. Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
Yashwant Pathak, Deepak Thassu, and Michel Deleers
185
14. Lipid Nanoparticles (Solid Lipid Nanoparticles and Nanostructured Lipid
Carriers) for Cosmetic, Dermal, and Transdermal Applications 213
Eliana B. Souto and Rainer H. Müller
15. Nano-Carriers of Drugs and Genes for the Treatment of Restenosis 235
Einat Cohen-Sela, Victoria Elazar, Hila Epstein-Barash, and Gershon Golomb
16. Ocular Applications of Nanoparticulate Drug-Delivery Systems
Annick Ludwig
271
17. Nanoparticulate Systems for Central Nervous System Drug Delivery
Jean-Christophe Olivier and Manuela Pereira de Oliveira
18. Nanoparticles for Gene Delivery: Formulation Characteristics
Jaspreet K. Vasir and Vinod Labhasetwar
281
291
19. Gastrointestinal Applications of Nanoparticulate Drug-Delivery Systems
Maria Rosa Gasco
20. Nanoparticles as Adjuvant-Vectors for Vaccination 317
Socorro Espuelas, Carlos Gamazo, María José Blanco-Prieto, and Juan M. Irache
21. Transdermal Applications of Nanoparticulates
Jongwon Shim
Index . . . . 339
327
305
Contributors
María José Blanco-Prieto Department of Pharmacy and Pharmaceutical Technology,
University of Navarra, Pamplona, Spain
Heike Bunjes Department of Pharmaceutical Technology, Institute of Pharmacy,
Friedrich Schiller University Jena, Jena, Germany
Matthew D. Burke Department of Pharmaceutical Development, GlaxoSmithKline,
Research Triangle Park, North Carolina, U.S.A.
Sudhir S. Chakravarthi Department of Pharmaceutical Sciences, University of
Nebraska Medical Center, Omaha, Nebraska, U.S.A.
Joseph F. Chiang Department of Chemistry and Biochemistry, State University of New
York at Oneonta, Oneonta, New York, U.S.A., and Department of Chemistry, Tsinghua
University, Beijing, China
Einat Cohen-Sela Department of Pharmaceutics, School of Pharmacy, The Hebrew
University of Jerusalem, Jerusalem, Israel
Sinjan De Research and Development, Perrigo Company, Allegan, Michigan, U.S.A.
Michel Deleers Global Pharmaceutical Technology and Analytical Development
(GPTAD), UCB, Braine l’Alleud, Belgium
Hannele Eerikäinen Pharmaceutical Product Development, Orion Corporation
Orion Pharma, Espoo, Finland
Victoria Elazar Department of Pharmaceutics, School of Pharmacy, The Hebrew
University of Jerusalem, Jerusalem, Israel
Hila Epstein-Barash Department of Pharmaceutics, School of Pharmacy, The Hebrew
University of Jerusalem, Jerusalem, Israel
Socorro Espuelas Department of Pharmacy and Pharmaceutical Technology,
University of Navarra, Pamplona, Spain
Carlos Gamazo Department of Microbiology, University of Navarra, Pamplona, Spain
Maria Rosa Gasco
Nanovector srl, Torino, Italy
Gershon Golomb Department of Pharmaceutics, School of Pharmacy, The Hebrew
University of Jerusalem, Jerusalem, Israel
Juan M. Irache Department of Pharmacy and Pharmaceutical Technology,
University of Navarra, Pamplona, Spain
K. K. Jain
Jain PharmaBiotech, Basel, Switzerland
Ashwath Jayagopal Biomaterials, Drug Delivery, and Tissue Engineering Laboratory,
Department of Biomedical Engineering, Vanderbilt University, Nashville,
Tennessee, U.S.A.
Esko I. Kauppinen NanoMaterials Group, Laboratory of Physics and Center for New
Materials, Helsinki University of Technology, and VTT Biotechnology, Helsinki, Finland
ix
x
Contributors
Judith Kuntsche Department of Pharmaceutical Technology, Institute of Pharmacy,
Friedrich Schiller University Jena, Jena, Germany
Vinod Labhasetwar Department of Pharmaceutical Sciences, University of Nebraska
Medical Center, Omaha, Nebraska, U.S.A.
Anna Lähde NanoMaterials Group, Laboratory of Physics and Center for
New Materials, Helsinki University of Technology, Helsinki, Finland
Robert J. Lee Division of Pharmaceutics, College of Pharmacy, NCI Comprehensive
Cancer Center, NSF Nanoscale Science and Engineering Center for Affordable
Nanoengineering of Polymeric Biomedical Devices, The Ohio State University,
Columbus, Ohio, U.S.A.
Annick Ludwig Department of Pharmaceutical Sciences, University of Antwerp,
Antwerp, Belgium
Dmitry Luzhansky Department of Corporate Technology, Donaldson Company, Inc.,
Minneapolis, Minnesota, U.S.A.
Rainer H. Müller Department of Pharmaceutical Technology, Biotechnology, and
Quality Management, Freie Universität Berlin, Berlin, Germany
Jan Möschwitzer Department of Pharmaceutical Technology, Biotechnology, and
Quality Management, Freie Universität Berlin, Berlin, Germany
Jean-Christophe Olivier Pharmacologie des Médicaments Anti-Infectieux, Faculty of
Medicine and Pharmacy, and INSERM, ERI 023, Poitiers, France
Giulio F. Paciotti
CytImmune Sciences, Inc., Rockville, Maryland, U.S.A.
Yashwant Pathak UCB Manufacturing, Inc., Rochester, New York, U.S.A.
Manuela Pereira de Oliveira Pharmacologie des Médicaments Anti-Infectieux,
Faculty of Medicine and Pharmacy, and INSERM, ERI 023, Poitiers, France
Janne Raula NanoMaterials Group, Laboratory of Physics and Center for
New Materials, Helsinki University of Technology, Helsinki, Finland
Barrett E. Rabinow
Baxter Healthcare Corporation, Round Lake, Illinois, U.S.A.
Dennis H. Robinson Department of Pharmaceutical Sciences, University of Nebraska
Medical Center, Omaha, Nebraska, U.S.A.
V. Prasad Shastri Biomaterials, Drug Delivery, and Tissue Engineering Laboratory,
Department of Biomedical Engineering, Vanderbilt University, Nashville, Tennessee, U.S.A.
Jongwon Shim Nanotechnology Research Team, Skin Research Institute, R&D
Center, Amorpacific Corporation, Kyounggi, South Korea
Eliana B. Souto Department of Pharmaceutical Technology, Biotechnology, and
Quality Management, Freie Universität Berlin, Berlin, Germany
Lawrence Tamarkin CytImmune Sciences, Inc., Rockville, Maryland, U.S.A.
Deepak Thassu UCB Pharma, Inc., Rochester, New York, U.S.A.
Jaspreet K. Vasir Department of Pharmaceutical Sciences, University of Nebraska
Medical Center, Omaha, Nebraska, U.S.A.
Jun Wu Division of Pharmaceutics, College of Pharmacy, The Ohio State University,
Columbus, Ohio, U.S.A.
Xiaobin Zhao Division of Pharmaceutics, College of Pharmacy, The Ohio State
University, Columbus, Ohio, U.S.A.
1
Nanoparticulate Drug-Delivery
Systems: An Overview
Deepak Thassu
UCB Pharma, Inc., Rochester, New York, U.S.A.
Yashwant Pathak
UCB Manufacturing, Inc., Rochester, New York, U.S.A.
Michel Deleers
Global Pharmaceutical Technology and Analytical Development (GPTAD),
UCB, Braine l’Alleud, Belgium
INTRODUCTION
Nanotechnology and nanoscience are widely seen as having a great potential to bring
benefits to many areas of research and applications. It is attracting increasing investments from governments and private sector businesses in many parts of the world.
Concurrently, the application of nanoscience is raising new challenges in the safety,
regulatory, and ethical domains that will require extensive debates on all levels.
The prefix nano is derived from the Greek word dwarf. One nanometer (nm) is
equal to one-billionth of a meter, that is, 10−9 m. The term “nanotechnology” was first
used in 1974, when Norio Taniguchi, a scientist at the University of Tokyo, Japan,
referred to materials in nanometers. The size range that holds so much interest is typically from 100 nm down to the atomic level approximately 0.2 nm, because in this
range materials can have different and enhanced properties compared with the same
material at a larger size. Figure 1 shows the nanometer in context (1). Nanotechnologies
have been used to create tiny features on computer chips for the last 25 years. The
natural world also contains many examples of nanoscale structures, from milk (a
nanoscale colloid) to the sophisticated nanosized and nanostructured proteins that
control a range of biological activities, such as flexing muscles, releasing energy, and
repairing cells. Nanoparticles (NPs) occur naturally and have been in existence for
thousands of years as products of combustion and cooking of food.
Nanomaterials differ significantly from other materials due to the following
two major principal factors: the increased surface area and quantum effects. These
factors can enhance properties such as reactivity, strength, electrical characteristics,
and in vivo behavior. As the particle size decreases, a greater proportion of atoms are
found at the surface compared to inside. For example, a particle size of 30 nm has 5%
of its atoms on the surface, at 10 nm 20%, and at 3 nm 50% of the atoms are on surface
(1). Thus, an NP has a much greater surface area per unit mass compared with larger
particles, leading to greater reactivity. In tandem with surface area effects, quantum
effects can begin to dominate the properties of matter as size is reduced to the nanoscale. These can affect the optical, electrical, and magnetic behavior of materials. Their
in vivo behavior can be from increased absorption to high toxicity of nanomaterials.
METHODS OF MEASUREMENTS AND CHARACTERIZATION
OF NANOMATERIALS
Nanometrology is the science of measurements at the nanoscale, and its application
underlies all the nanoscience and nanotechnology. The ability to measure and
1
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FIGURE 1 (See color insert.) Length scale showing the nanometer in context. The length scale of interest for nanoscience and nanotechnologies is from 100 nm down to the atomic scale approximately 0.2 nm. Source: From Ref. 1.
Nanoparticulate Drug-Delivery Systems
3
characterize materials, as well as determine their shape, size, and physical properties at the nanoscale is vital for nanomaterials and devices. These need to be
produced to a high degree of accuracy and reliability, to realize the applications
of nanotechnologies. Nanometrology includes length and/or size (where dimensions are typically in nanometers) as well as measurement of force, mass, electrical,
and other properties. Four commonly used techniques are: transmission
electron microscopy (TEM), scanning electron microscopy (SEM), scanning probe
techniques [scanning probe microscopy (SPM)], and optical tweezers (single-beam
gradient trap).
Transmission Electron Microscopy
TEM is used to investigate the internal structure of micro- and nanostructures. It works
by passing electrons through the samples and then using magnetic lenses to focus the
image of the structure. TEM can reveal the finest details of the internal structure, in
some cases the individual atoms. TEM with high-resolution transmission electron
microscopy is the important tool for the study of NP.
Scanning Electron Microscopy
SEM uses the basic technology developed for TEM, but the beam of electrons is
focused to a diameter spot of approximately 1 nm on the surface of the specimen and
scanned repetitively across the surface. It reveals that the surface topography of the
sample with the best spatial resolution currently achieved is on the order of 1 nm.
Scanning Probe Techniques (Scanning Probe Microscopy)
SPM uses the interaction between a sharp tip and a surface to obtain the image. The
sharp tip is held very close to the surface to be examined and is scanned back and
forth. As the tip is scanned across the sample, the displacement of the end of the
cantilever is measured, using a laser beam. This can image insulating materials
simply because the signal corresponds to the force between the tip and the sample,
which reflects the topography being scanned. The scanning tunneling microscope
brought a Noble prize for physics to Gerd Binnig and Heinrich in 1986. The atomic
force microscope uses this SPM technique, which reflects the surface topography of
the samples.
Optical Tweezers (Single-Beam Gradient Trap)
Optical tweezers use a single laser beam (focused by a high-quality microscope
objective) to a spot on the specimen plane. The radiation pressure and gradient
forces from the spot create an optical trap, which holds a particle at its center. Small
interatomic forces and displacements can be measured by this technique. Samples
that can be analyzed range from single atoms to micrometer-sized spheres to strands
of DNA and living cells. Numerous traps can be used simultaneously with other
optical techniques, such as scalpels, which can cut the particle being studied. Various
analytical techniques utilized in nanometrology are enumerated in Table 1.
MANUFACTURE OF NANOMATERIALS
There are a wide variety of techniques that are capable of creating nanostructures
with various degrees of quality, speed, and cost. These manufacturing approaches
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TABLE 1 Analytical Techniques Used for Characterization of Nanoparticles
Name of the technique
Laser diffraction
Photon correlation spectroscopy
Wide-angle X-ray scattering
Differential scanning colorimetry
Proton nuclear magnetic resonance spectroscopy
Electron spin resonance
Electron transmission microscopy
Sedimentation velocity analysis and EM
DLS and cryo-TEM
DLS and TEM
Flow cytometry and ELISA method
Fluorometry
Fluorescence and TEM
Reference
(2)
(3)
(4)
(5)
(6)
(7)
(8)
(9)
Abbreviations: DLS, dynamic light scattering; ELISA, enzyme-linked immunosorbent assay; EM,
electron microscopy; NP, nanoparticle; TEM, transmission electron microscopy.
fall under two categories: bottom-up and top-down. Figure 2 illustrates the types of
materials and products which can be manufactured using these two approaches (1).
Bottom-Up Manufacturing
Bottom-up manufacturing involves the building of nanostructures atom by atom or
molecule by molecule. This can be done in three ways: chemical synthesis, selfassembly, and positional assembly.
Chemical synthesis is a method of producing raw materials, such as molecules or
particles, which can then be used either directly in products in their bulk-disordered
form or as the building blocks of more advanced ordered materials. Figure 3 represents the generic processes that are involved in the production of NPs (1):
1. Self-assembly is a production technique in which atoms or molecules arrange
themselves into ordered nanoscale structures by physical or chemical interactions
within the smaller units. The formation of salt crystals and snowflakes with their
intricate structure are examples of the self-assembly process. Although selfassembly occurs in nature, in industry it is relatively new and not a wellestablished process (1).
FIGURE 2 The use of bottom-up and top-down techniques in manufacturing nanoparticles.
Abbreviation: MEMS, microelectromechanical system. Source: From Ref. 1.
Nanoparticulate Drug-Delivery Systems
5
FIGURE 3 The generic processes that are involved in the production of nanoparticles. Source:
From Ref. 1.
2. In positional assembly, atoms, molecules, or clusters are deliberately manipulated
and positioned one by one. Techniques such as SPM for work on surfaces or
optical tweezers in free space are used for this. Positional assembly is extremely
laborious and rarely used as an industrial process.
Top-Down Manufacturing
Top-down manufacturing involves starting with a larger piece of material, and
etching, milling, or machining a nanostructure from it by removing material. Topdown methods offer reliability and device complexity. These are higher in energy
usage and produce more waste than the bottom-up methods.
Although the nanotechnologies have been used by industries for many
decades (semiconductor and chemical industry), it is still very much at infancy
stage. In recent years, the tools used to characterize materials (Table 1) have led to
better understanding of the behavior and properties of matter on a very small scale.
Increased knowledge of the relationship between the structure and properties of
nanomaterials has enabled the production of materials and devices with higher
performance and increased functionality. At the same time, there are uncertainties
which need to be addressed about the direction that nanotechnology will take, and
about the hazards to humans and the environment that are presented by certain
aspects of this technology (10).
There are several good reports and reviews which cover the production and
characterization of NPs and nanoparticulate drug-delivery systems (NPDDSs).
Venkateswarlu and Manjunath (11) have discussed the preparation and characterization of clozapine NPs. They used hot homogenization and later ultrasonication
method to formulate solid–lipid nanoparticles (SLNs) incorporating clozapine.
Dingler and Gohla (12) have discussed the production of SLN and scaling up studies
and Gasco (13) has patented a method for producing SLN. Mehnert and Mader (14)
have written an excellent review about the SLN production and characterization.
Many reviews are reported by Muller et al. (15) and others (2–5,16–18). Rigaldie
et al. (19) have shown the high-hydrostatic-pressure technique to preserve and
sterilize the spherulites, an NPDDS. Several papers and patents are reported by our
group (20–26) and Rodriguez et al. described a high-pressure emulsification and
homogenization process for NPDDS preparation (17).
Microfluidics is being explored for applications in NPDDS. It is based on
instruments that are capable of transferring small volumes of liquids ranging from
microliters to nanoliters. Microfluidic “lab-on-the-chip” technology requires an
understanding of the forces that control fluid movement and reaction conditions
and brings the potential benefits of miniaturization, integration, and automation.
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Manufacturing such chips combines methods from microchip industry with
expertise in fluid mechanics, biochemistry, and hardware engineering to create
miniature integrated biochemical-processing systems. A microfluidics platform
provides better quality data and allows shorter assay development times. Owing to
the direct measurement at nanoscale and the high-quality data generated by
microfluidics, this technology platform is finding a place in drug discovery as well
as NPDDS (27–31).
DRUG-DELIVERY SYSTEMS
An ideal drug-delivery system possesses two elements: the ability to target and to
control the drug release. Targeting will ensure high efficiency of the drug and reduce
the side effects, especially when dealing with drugs that are presumed to kill cancer
cells but can also kill healthy cells when delivered to them. The reduction or prevention of side effects can also be achieved by controlled release. NPDDS provide a
better penetration of the particles inside the body as their size allows delivery via
intravenous injection or other routes. The nanoscale size of these particulate systems also minimizes the irritant reactions at the injection site. Early attempts to
direct treatment to a specific set of cells involved attaching radioactive substances to
antibodies specific to markers displayed on the surface of cancer cells. Antibodies
are the body’s means of detecting and flagging the presence of foreign substances.
Antibodies specific to certain proteins can be mass produced in laboratories, ironically using the cancer cells. These approaches have yielded some good results, and
NPDDSs are demonstrating lot of potential in this area.
Lipid-Based Colloidal Nanodrug-Delivery Systems
Lipid nanocapsules are submicron particles made of an oily liquid core surrounded
by a solid or semisolid shell. NPDDSs were primarily developed to combine the
colloidal stability of solid particle suspensions in biological fluids and the solubilizing properties of liquids (32,33). SLNs were invented at the beginning of 1990s
and are produced either by high-pressure homogenization or by microemulsion
technique (34). SLNs consist of solid matrix and can be described as parenteral
emulsions in which the liquid–lipid oil is replaced by a solid–lipid. Owing to their
solid particle matrix, they can protect incorporated ingredients against chemical
degradation (35) and allow modification of release of the active compounds (36).
Homogenization followed by ultrasonication was used for the production of
clozapine-loaded SLNs (11).
Colloidal drug carriers offer a number of potential advantages as delivery
systems, such as better bioavailability for poorly soluble drugs. Other advantages of
these SLNs include: use of physiological lipids, the avoidance of organic solvents in
the preparation process, a wide potential application spectrum (oral, dermal, and
intravenous), high-pressure homogenization as an established production method
(which allows large-scale production), improved bioavailability, protection of sensitive drug molecules from the environment (water and light), and a controlled release
characteristic (14). Common disadvantages of SLNs include: particle growth, unpredictable gelation tendency, unexpected dynamics of polymorphic transitions, and
inherently low incorporation capabilities due to crystalline structure of the SLN
(14). The key parameters in characterizing the SLN include: concentration, nanocapsule size and shape, thickness, and shell composition, defining the freeze-drying
Nanoparticulate Drug-Delivery Systems
7
FIGURE 4 Proposed topology for lipid nanocapsules
freeze-dried in the presence of trehalose. Source: From
Ref. 37.
conditions such as cryoprotectant, pressure, and temperature cycle. Some of the factors for the formulation of the lipid NP are: the drug payload depends on the oil
content, the evolution of the hydrophilic–lipophilic balance of solutol HS15 is the
driving force of SLN formation (Fig. 4), and the SLN diameter depends on both the
Foil/F solutol and the solutol HS15/Lipoid S 100 ratios (37). Besides nanoemulsions,
nanosuspensions (38), mixed micelles and liposomes, melt-emulsified NP-based
lipids, and solids at room temperature have been developed (15). The low incorporation capabilities were overcome by using liquid–lipid nanostructured carriers (39).
Several excellent reviews and papers on the SLN are reported (14,38,40–44).
Recent Trends in Solid–Lipid Nanoparticle Research
Recently, a lipid-based solvent-free formulation process has been developed to
prepare lipid nanocapsules in the nanometer range (32,45). This process takes
advantage of the variation of the hydrophilic–lipophilic balance of an ethoxylated
hydrophilic surfactant as a function of the temperature, leading to an inversion
phase. In the first step, several temperature cycles applied around the inversionphase temperature lead to droplet size decrease and homogenization. In a second
step, fast cooling leads to the crystallization of the lecithin (introduced in the
formulation both as lipophilic cosurfactant and constituting material of the rigid
shell), which leads to the formation of a stable lipid nanocapsule suspension. This
suspension can be freeze-dried and resuspended in an aqueous medium extemporaneously. The freeze-drying can alter the topology of the NPs; hence while doing
so, the structure of the NPs needs to be preserved (37).
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Jores et al. (2) have studied physicochemical investigations of SLN and oilloaded SLN using nuclear magnetic resonance and electron spin resonance. They
have investigated various techniques to evaluate and characterize them using
photon correlation spectroscopy. Laser diffraction was used for particle size determination, and field flow fractionation with multiangle light scattering detection was
used to separate the particles due to their Stokes radius. It helped in understanding
the topography of the particles. Cryo-TEM was used to study the ultrastructure
of the NPs.
SLNs have been shown to condense DNA into nanometric colloidal particles
capable of transfecting mammalian cells in vitro (46). Compared with standard
DNA carriers such as cationic lipids or cationic polymers, SLN offers several technological advantages such as a relative ease of production without organic solvents,
the possibility of large-scale production with qualified production lines, good storage stabilities, possibility of steam sterilization, and lyophilization (47). In a study
by Rudolph et al. (47), a diametric tyrosine aminotransferase (TAT) peptide derived
from the arginine-rich motif of the HIV-1 TAT protein that functions as nuclear
localization sequence and as a protein transduction domain could be used to
substantially enhance gene transfer efficiency of SLN-based vectors, leading to gene
expression levels even higher than observed for polyethylenimine (PEI) gene vectors.
This might allow aerosol application of fragile gene delivery systems to lungs in the
future studies.
The common ground of cationic liposome nanoemulsions (48) and SLNs for
transfection is the need for cationic lipids to facilitate deoxyribonucleic acid (DNA)
binding. In liposome formulations, these lipids are arranged as bilayers around an
aqueous core. Interaction with DNA initiates structural rearrangements into different
structures depending on the kind of lipid, lipid/DNA ratio, incubation media, and
time (38). Tabatt et al. (49) have shown equipotency of SLNs and liposome formulated from the cationic lipids in in vitro DNA transfection efficiency.
A study by Kogure et al. (50) demonstrated the development of a multifunctional envelope-type nanodevice for a gene-delivery system. This contained
membrane-permeable peptide R8 with less cytotoxicity. This system can incorporate
various functional devices such as a specific ligand to a specific cell, intracellular
sorting devices that permit endosomal escape, and nuclear localization. This lipidbased device can be a useful tool for gene delivery for gene therapy and biochemical
research (Fig. 5: schematic steps for nanodevices). Lee et al. (51) reported an increased
stability and controlled release of lovastatin by microencapsulating the drug-loaded
lipid NPs. Several studies have shown the application of SLN formulation for the
delivery of paclitaxel and its pro-drug for cancer treatment (52).
Hou et al. (53) have described the modified high shear homogenization and
ultrasound techniques to produce SLNs. Model drug mifepristone was incorporated
in SLNs, and the mean particle size was found to be 106 nm. The drug entrapment
efficiency was more than 87% and showed relatively long stability, as the leakages
were small. Olbrich et al. (54) described the potential delivery of hydrophilic antitrypanosomal drug diminazine diaceturate to the brain of infected mice formulating
the lipid drug conjugate NP by combination of stearic and oleic acids. An excellent
work on an in vivo evaluation of tobramycin SLNs and their duodenal administration is described by Cavalli et al. (55), and is further discussed in the following
chapters in detail.
Williams et al. (56) have studied lipid-based NP formulation of SN38, a camptothecin analog used as antineoplastic drug. They showed improved drug loading
Nanoparticulate Drug-Delivery Systems
9
FIGURE 5 Three steps involved in constructing the
multifunctional envelope-type nanodevice. Source: From
Ref. 50.
and good lactone stability in the presence of human serum albumin (HSA). The
NPDDS showed prolonged circulation in murine blood and better efficacy against a
resistant model of human colon carcinoma in nude mice. It was also demonstrated
that the blood half-life of SN38 was greatly prolonged by incorporation in NPs.
Nanoparticulate Polymeric Micelles as Drug Carriers
Polymeric micelles have attracted much attention in drug delivery, partly because
of their ability to solubilize hydrophobic molecules, their small particle size, good
thermodynamic solution stability, extended release of various drugs, and prevention of rapid clearance by the reticuloendothelial system (RES) (57). Critical micelle
concentration (CMC), similar to low-molecular-weight surfactants, is the key characterization parameter for polymeric micelles. CMC is the concentration at which
the amphiphilic polymers in aqueous solution begin to form micelles while coexisting in the equilibrium with the individual polymer chains or unimers. At CMC or
slightly above the CMC, the micelles form loose aggregates and contain some water
in the core (58). With further increases in amphiphilic polymer concentration, the
unimer to micelle equilibrium shifts towards micelle formation. The micellar structure then becomes more compressed and stable, whereas residual solvent is excluded
from the core, and the micelle size is reduced. The lower CMC values correlate to
more stable micelles. This concept is especially important from the pharmacological
point of view, as upon dilution with a large volume of the blood, micelles with high
CMC values may dissociate into unimers and their content may precipitate out,
whereas the micelles with low CMC are more likely to remain the same. Thus, to
develop improved drug-delivery systems, amphiphilic molecules that are able to form
more stable micelles with lower CMC values are appropriate candidates. A fascinating study reported by Djordjevic et al. (59) utilized scorpion-like amphiphilic
macromolecules. They used indomethacin as a model drug for the study and
reported this method as convenient for drug delivery while minimizing drug toxicity
and maximizing the drug effectiveness.
Generally, the amphiphilic core/shell structure of polymeric micelles is formed
from block copolymers, which are hydrophobic polymer chains linked to hydrophilic
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Thassu et al.
polymer chains (60). Association of the hydrophobic portions of the block copolymers creates the inner micelle core due to their cohesive interactions with each other
in aqueous media (i.e., hydrophobic interactions), whereas the outer hydrophilic
portions surround the inner hydrophobic core as a hydrated shell (61).
Polymeric micelles are self-assemblies of block copolymers in aqueous media.
Many advantages have been demonstrated with their unique core shell architecture.
Hydrophilic shells from the aqueous exterior segregate the hydrophobic cores.
Hydrophobic drugs can be solubilized into the hydrophobic core structures of polymeric micelles at concentrations much higher than their intrinsic water solubility.
Polymeric micelles are known to have high drug-loading capacity, high water solubility, and appropriate size for long circulation in the blood (62). The hydrophilic
shell surrounding the micellar core can protect undesirable phenomena such as
intermicellar aggregation, or precipitation, protein adsorption, and cell adhesion.
The chemical composition of the polymeric micelles can be tailor-made to have
desirable physicochemical properties for drug solubilization (63). The hydrophobic
drug is incorporated into the hydrophobic core by interactions such as metal–ligand
coordination bonding and electrostatic interaction. The extent of drug solubility
depends on the compatibility between the drug and the micelle core (64). One of the
limitations of drug-loaded polymeric micelles is low stability in aqueous solution,
and the stability becomes even lower as the drug-loading content increases (65).
Various types of drugs can be loaded into the hydrophobic core of polymeric
micelles by chemical conjugation or physical entrapolymeric micelles sent utilizing
various interactions such as hydrophobic interactions, or ionic interactions, or hydrogen bonding. Furthermore, the hydrophobic core serves as a reservoir from which
the drug is released slowly over an extended period of time. The hydrophobic inner
core is solubilized by the hydrophilic shell, which prevents the inactivation of the
core-encapsulated drug molecules by decreasing the contact with the inactivating
species in the aqueous (blood) phase. As the outer hydrophilic part of the polymeric
micelles interacts with biocomponents such as cells and proteins, it affects their
pharmacokinetics and disposition, as well as their surface properties (66).
Polymeric Micelles and Solubilization of Drugs
Solubilization of drugs is a complex mechanism that involves different parameters,
for example, hydrophobicity, molecular volume, crystallinity, flexibility, charge, and
interfacial tension against water. The lack of water solubility hampers the use of
many potent pharmaceuticals. Polymeric micelles are self-assembled nanocarriers
with versatile properties that can be engineered to solubilize, target, and release
hydrophobic drugs in a controlled release fashion. Unfortunately, their large-scale
use is limited by the incorporation methods available. This poses a problem when
sterile dosage forms are formulated. Polymeric micelles present a core shell architecture that results from the self-assembly of the amphiphilic block polymers in a
selective solvent above a threshold concentration referred to as critical association
concentration (67). Their structure is such that the core provides an isolated cargo
space where hydrophobic drugs can partition. This is of great significance as many
potent pharmaceuticals are highly hydrophobic by nature. The nanometric size of
polymeric micelles varies from 10 to 100 nm and the flexible highly hydrated corona
minimizes nonselective scavenging and rapid clearance by the monolayer phagocyte system. These drug carriers can extravasate and accumulate passively in regions
presenting leaky vasculatures such as tumors, inflamed and infracted tissues (68).
Nanoparticulate Drug-Delivery Systems
11
FIGURE 6 Scheme of the freeze-drying
procedure for water-soluble amphiphilic
nanocarriers. Source: From Ref. 44.
Recently, polymeric micelles have also been shown to distribute to defined cytoplasmic organelles (69) and increasing efforts are now directed at targeting the
subcellular components (44). A simple method to have higher drug loading in the
amphiphilic nanocarriers polymeric micelles was developed by Fournier et al. (44).
Figure 6 shows the schematic production of polymeric micelles and its freeze-drying
procedure.
Polymeric Micelles and Reticuloendothelial System
Polymeric micelles provide an attractive characteristic in that they can avoid uptake
of the drugs by RES in vivo and hence these can circulate in the blood for a longer
time. This advantage comes from the structure of a micelle, the hydrophilic portions
of the amphiphilic block copolymer form the outer shell and are exposed to body
fluid, and hence the micelles can be protected from phagocytic cells and plasma
proteins in blood. Another important biological advantage of polymeric micelles is
the EPR9-enhanced permeability and retention effect or passive targeting. As a
result, polymeric micelles can slowly accumulate in malignant or inflamed tissues
due to the elevated levels of vascular permeability factors in such cells (70). Polymeric
micelles seem to be ideal carriers for poorly water-soluble drugs because of their
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distinct advantages such as high solubility, long circulation of drug in blood, permeation of an anticancer drug by the EPR effect (71), and simple sterilization. They
have two major disadvantages: physical instability upon dilution limits their
pharmaceutical application and water-soluble drugs cannot be in the micelles.
Recent Trends in Polymeric Micelles Research
Francis et al. (72) have studied the polysaccharide-based polymeric micelles for the
delivery of cyclosporine A. They demonstrated that coupling of hydrophobic groups
to water-soluble polysaccharides significantly promotes the solubilizing power of
either dextran or hydroxypropyl cellulose (HPC) toward cyclosporine A. The bioadhesive properties of HPC enhance the association of polymeric micelles toward
caco-2 cell monolayers and facilitate internalization of the polymer and the transport
of the drug. The polysaccharide-based polymeric micelles offer unique opportunities for the oral delivery of lipophilic drugs. Similar studies for cyclosporine have
been reported using SLN (73), polycaprolactone NPDDS (74), poly-lactic acid polyethylene glycol NPs (75), and chitosan derivatives (76).
A new modality of drug targeting tumors is based on drug encapsulation in
polymeric micelles followed by a localized triggering of the drug intracellular uptake
induced by ultrasound, which is focused into the tumor volume (77). A rationale
behind this approach is that drug encapsulation in polymeric micelles decreases a
systemic concentration of free drug, diminishes intracellular drug uptake by normal
cells, and provides passive drug targeting of tumors via enhanced penetration and
retention effect as a result of abnormal permeability of tumor blood vessels (78).
Drugs targeting tumors reduce unwanted drug interactions with healthy tissues
(79). With micelle accumulation in the tumor interstitium, an effective intracellular
drug uptake by the tumor cells should be ensured, making it possible for ultrasonic
irradiation to be used (77). The in vitro and in vivo experiments have suggested that
polymeric micelles can be degraded into unimers under the action of ultrasound,
which may provide an additional advantage of in vivo sensitization of multidrugresistant cells (80). It is suggested that this technique can be useful in treatment
of ovarian carcinoma tumors of small size; hence early detection is necessary for
tumor treatment.
POLYMER-BASED NANOPARTICULATE DRUG-DELIVERY SYSTEMS
Several polymers and nonlipid materials have been evaluated as carriers for drugs
in the nanoparticulate forms. These materials have shown different properties and
advantages when formulated as drug-delivery systems. A brief description of each
of the polymeric systems follows.
Hydrogel-Based Nanoparticulate Drug-Delivery Systems
A progressively increasing interest has been paid to self-assembled hydrogel NPs
from hydrophobized water-soluble polymers due to their potential biomedical and
pharmaceutical applications (81). The NPs have shown various structural and
rheological features in aqueous solutions depending on the structure of the parent
water-soluble polymer, conjugated hydrophobic moiety or groups, and the degree
of substitution. The formation of self-assembled NPs is theorized by a free-energyminimized structure, sharing a common feature of assembly of polymeric micelles.
However, there exists a difference in the interior structure between NPs and
Nanoparticulate Drug-Delivery Systems
13
polymeric micelles formed from amphiphilic block copolymers. The interior of
polymeric NPs consists of dispersed multiple hydrophobic island domains in a
hydrophilic sea domain due to the random association of hydrophobic moieties
conjugated to soluble macromolecules. Polymeric micelles provide one inner core of
hydrophobic segments with a hydrophilic shell (82,83). The NPs formed from polymers containing moiety switching its hydrophilicity by external stimuli is expected
to exhibit stimuli responsive surface property plus macroscopic hydrogel bulk
property. These properties might lead to the accumulation of the NPs at a disease
site and the change of drug-release behavior from slow-to-fast drug release. An
interesting study using pullulan acetate/sulfonamide conjugates in self-assembled
NPs responsive to pH change was reported by Na et al. (81).
Amphiphilic block copolymers are widely studied as potential NPDDSs as
they are capable of forming aggregates in aqueous solutions (84,85). These aggregates are comprised of a hydrophilic shell and hydrophobic core. They are good
vehicles for delivering hydrophobic drugs because the drugs are protected from
possible degradation by enzymes. Changing the composition of hydrophobic and
hydrophilic blocks on the polymer chains can vary the morphology of NPs produced from amphiphilic block copolymers. Various forms of morphologies such as
sphere, vesicles, rods, lamellas, tubes, large compound micelles, and large compound vesicles have been reported. Some of these structures are good candidates
for drug-delivery applications (86). Compared with normal shell micelles, vesicles
with a hydrophilic core and hydrophobic layers are better for drug delivery. In the
clinical studies, it has been shown that vesicles improve the treatment efficacy of
anticancer drugs such as doxorubicin due to enhanced permeability and retention
properties (87). The block copolymers comprised of commercial pluronic systems
and biodegradable poly(lactic acid) are very good carriers for drug delivery and
controlled release applications (88).
Zhang et al. (89) synthesized triblock copolymers of poly(caprolactoneco-lactide)–b-poly(caprolactone-co-lactide) (PCLLA–PEG–PCLLA) by ring-opening
copolymerization of caprolactone and lactide in the presence of polyethylene glycol.
They entrapped an anticancer drug, a camptothecin derivative by nanoprecipitation
technique. The in vitro and in vivo evaluation of this NPDDS showed a potential for
use with poorly soluble anticancer drugs. They demonstrated that the drug release
from these systems can be controlled by controlling the particle size, as they found
the larger the NPs size, the lower was the drug release. The body distribution of
these NPs showed that the blood concentration can be maintained for a longer time,
and the tissue body distribution was affected by the particle size (90). Several other
groups have shown the application of the triblock copolymers for NPDDSs (91–93).
Yoo and Park (94) have shown folate receptor-targeted PLGA–PEG micelles entrapping a high loading amount of doxorubicin, showing better uptake of the drug. The
in vitro and in vivo studies have shown the accumulation of the drug in the tumor
cells in a site-specific manner.
An excellent review on block copolymer micelles for drug delivery, design,
characterization, and biological significance is written by Kataoka et al. (60). Another
review on applications of poly(ethylene oxide) block copolymer–poly(amino acids)
micelles is published by Lavasanifar et al. (95). Vriezema et al. (96–98) have reported
some interesting methods to produce block copolymers.
NPs based on hydrogels are being developed for the delivery of macromolecules, and some of the candidates of hydrogel utilized for this purpose are
enumerated in Table 2. Many polymeric carriers were reported useful in the
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TABLE 2 Hydrogel Matrices
Based on natural
materials
Collagen
Gelatin
Starch
Alginates
Chitosans
Dextrans
Synthetic polymers
Responsive polymers
Poly(n-vinyl pyrrolidine)
Poly(vinyl alcohol)
Poly(phosphazenes)
Poly[ethylene oxide-b-poly(propylene
oxide)] copolymers
PEO–PPO–PAA graft copolymer
PL(G)A/PEO/Pl(GA) copolymers
PVA-g-PLGA graft polymers
PEGT–PBT copolymers (polyactive)
MA–oligolactide–PEO–oligolactide–MA
Methacrylates
Poly(N-isopropylacrylamide)LCST
PEO–PPO–PEO Pluronics
PLGA–PEO–PLGA (LCST)
Source: From Ref. 104.
formulation of NPDDSs, especially in the treatment of cancer, for example, poly(2
ethyl-2-oxazoline) block-poly-caprolactone (99), polyalkyl cyanoacrylate polymers
(100), PLGA NPs (101), polysaccharide decorated polyisobutyl cyanoacrylate NPs
(102), and serum albumin NPs (103).
Dendrimer-Based Drug-Delivery Systems
Three-dimensional tree-like branched macromolecules possess some fascinating
characteristics: a well-defined structure, a very narrow molecular weight distribution,
a three-dimensional structure tuned by dendrimer generation and dendron structure,
and flexibility for tailored functional groups with high density on the periphery (66).
Studies of biomedical application of dendrimers are becoming more and more
attractive especially in the field of nonviral gene vector and NPDDS (105,106).
Photodynamic therapy of cancer involves the systemic administration of
photosensitizers to solid tumor tissues and local illumination with light of a specific
wavelength, leading to photochemical destruction of cancer cells via generation of
singlet oxygen or superoxide from molecular oxygen. Suitable carriers and delivery
of photosensitizers should have a simple but effective strategy to realize high
selectivity, high photodynamic efficacy, and have less side effects. It is a challenge to
formulate the photosensitizers. Zhang et al. (107) reported the use of dendrimer
polymeric micelles for the delivery of photosensitizers successfully.
Calcium Carbonate Nanoparticles
Ueno et al. (108) have reported the incorporation of hydrophilic drugs and bioactive
proteins into solid calcium carbonate NPs. The size of the NPs was controlled by
mixing speed and was around 105 to 128 nm. These CaCO3 NPs were stable and
sustained the release of the drug betamethasone phosphate.
Proticles: Protamine-Based Nanoparticulate Drug Carriers
Protamine is a nonantigenic and virtually nontoxic peptide from the sperm, the
compound derived from salmon, the most widely used source, and has a molecular
mass around 5000 g/mol. It can be used as a carrier system for delivery of DNA or
oligonucleotides and it is being used as the cationic component. Several groups have
described the applications of proticles as drug-carrier systems (109–112). In most
studies, the peptide was employed together with relatively large double-stranded
Nanoparticulate Drug-Delivery Systems
15
DNA in a two-step procedure. In the first step, it is condensed with DNA into a compact particle and subsequently the complex was incorporated into protamine or suitable cationic liposomes. In some cases, transferring was also used. The term “proticles”
was used to represent oligonucleotides/protamine NPs by Dinaure et al. (95).
Vogel et al. (109) showed inclusion of HSA in the proticles led to dramatic
stability of the particles. They reported many advantages of this system such as: the
proticle production by self-assembly is simple and rather rapid. The excipients used
are nonantigenic and have very low toxicity and are well accepted in pharmaceutics. The particles are relatively stable in water and cell culture medium. They show
an increased uptake by a variety of cells, as compared to naked oligonucleotides,
and after cellular uptake, the oligodeoxyribonucleotide (ODN)/protamine NPs
readily release the active agent. They reported two distinct disadvantages: first,
they immediately show massive aggregation and precipitation when produced or
transferred into solutions containing salts at physiological ionic strength or even at
concentrations in the range of mmol/l, and secondly, those containing the more
stable (phosphorothioates PTOs) instead of ODNs (diesters) do not release their
nucleic acid after particle uptake by cells (109).
Chitosan-Based Nanoparticulate Drug-Delivery System
Chitosan, a polycationic polymer, comprising d-glucosamine and N-acetyl-dglucosamine linked by b-(1,4)-glycosidic bonds, has been extensively researched for
NPDDSs for delivering anticancer drugs, genes, and vaccines (113–116). In these
applications, it is important to assess the effectiveness of uptake of the carrier and
associated drug cargo into the target cells. Chitosan, being a natural polymer, is biocompatible. Chitosan NPs were also evaluated for ocular applications. The cationic
polysaccharide chitosan showed excellent properties such as biodegradability, nontoxicity, biocompatibility, and mucoadhesiveness, which are desirable for the ocular
delivery systems. An interesting study by Campos et al. (117) demonstrated that
Chitosan NPs were able to interact and remain associated to the ocular mucosa for
an extended period of time, thus promising carriers for enhancing and controlling
the release of drugs to the ocular surface. A review published by Hejazi et al. (103)
discusses various aspects of chitosan-delivery systems covering the availability,
physicochemical, and biological properties of chitosan. The review covered various
applications of chitosan-delivery systems for colon-specific delivery, as absorption
enhancers, and for GI tract delivery systems including the NPDDS. Park et al. (119) have
assessed the application of self-aggregates formed by modified glycol chitosan as a
carrier for peptide drugs. They exhibited comparable biological activity to parenteral
peptides. Fluorescein isothiocyanate (FITC)-labeled peptides were released from
the self-aggregates in a sustained manner for approximately a day. A report using
chitosan alginate combination nanospheres showed the utility of these nanospheres
for drug-delivery systems formulation (120). Self-assembled NPs containing hydrophobically modified chitosan for gene delivery was reported by Yoo et al. (121).
They demonstrated that modified glycol chitosan NPs composed of hydrophobized
DNA enhanced the transfection efficiencies in vitro as well as in vivo. Kumar et al.
(122) showed the application of chitosan-based NPs in treating allergic asthma.
Silicone Nanopore-Membrane-Based Drug-Delivery System
Top-down microfabrication techniques have been used to create nanopore membranes
consisting of arrays of parallel rectangular channels, which range from 7 to 50 nm.
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The original method was pioneered by Chu et al. (123) and consisting of two basic
steps, surface micromachining of nanochannels into a thin film on the top of the
silicon wafer and forming the nanopore membrane by etching away the bulk of the
silicone wafer underneath the thin-film structure. The experimental and mathematical results have shown that the devices outfitted with silicone nanopore membranes
can regulate the drug-delivery kinetics of a wide range of drugs. Moreover, the
mechanism of release is attributable to a novel constrained diffusion mechanism
provided by the precise geometry of the nanopore membrane itself, and no moving
parts such as pistons are required. The drugs can likely be loaded into the device
reservoir in a range of physical states, including solutions, crystalline, or micronized suspensions. Flexibility with respect to the physical form of encapsulated
drugs provides options to substantially increase the loaded dose and duration of the
therapy, as well as promoting approaches to increase stability of proteins, which are
intrinsically unstable in an aqueous solution at body temperature (124–126).
Polyester Polysaccharide Nanoparticles
NPs can be prepared from preformed copolymers by methods such as emulsificationsolvent evaporation, nanoprecipitation, or salting out, all of which require dissolution of polymers in organic solvents. Lemarchand et al. (127) reported a study using
interfacial migration-solvent evaporation method leading to NP formation using a
novel family of amphiphilic copolymers based on Dextran grafted with polycaprolactone side chains. They reported that these materials were found to be able to
self-organize and precipitate in the presence of mixtures of water and ethyl acetate.
Ethyl acetate in a water emulsion was stable and produced the best NPs.
Albumin and Gelatin Nanospheres
Since the first reports on the preparation of uniformly sized albumin microspheres
in the early 1970s, these biodegradable, biocompatible particles have found various
applications. Initially conceived as a diagnostic tool, albumin particles have been
utilized as drug-carrier systems (128). More than 100 therapeutic and diagnostic
agents were incorporated into albumin particles and have been investigated for
intravenous, intramuscular, intra-arterial, and intra-articular administration. Albumin
particles are well suited for drug targeting and drug delivery because of their lack
of toxicity and antigenicity. Compared with other colloidal carrier systems such as
liposomes, albumin nanospheres have better stability, shelf life, controllable drugrelease properties, and higher loading properties for hydrophilic molecules due to
drug-binding properties of native albumin. Albumin particles can be obtained by
many methods. In a study by Muller et al. (128), they optimized the manufacturing
techniques of albumin nanospheres with average diameter of 200 nm. They studied
the effect of five different process variables on particle size, polydispersity, and
yield, to optimize the preparation technique to reach sub-200-nm particles.
An interesting novel drug-delivery system for improved all-trans retinoic acid
(atRA) therapy for external treatments of photo-damaged skin was developed. The
research team prepared inorganic-coated atRA NPs using boundary-organized reaction droplets. The interfacial properties of organic architecture in atRA micelles were
used to template the nucleation of inorganic materials. When administered, they found
a boost in the production of hyaluronan among the intercellular spaces of the basal
and spinous cell layers of the epidermis. Nano-atRA technology for atRA therapy
could not only efficiently regulate keratinocyte cell proliferation and differentiation,
Nanoparticulate Drug-Delivery Systems
17
but also markedly produce the additional benefit. Human skin severely injured by
chronic ultraviolet irradiation may be completely repaired due to the accelerated
turnover of skin tissue, which is induced by nano-atRA (129–131).
Antibody-modified gelatin NPs have been reported to be a carrier system for
targeting the specific T-lymphocytes by Balthasar et al. (132). Gelatin NPs were
formed by two-step desolvation process. They showed the utility of this system for
targeting the lymphocytes.
Polymeric Nanocapsules as Drug Carriers
These were first prepared by solubilization of the outer shell material in an organic
solvent (133). Interesting biopharmaceutical performances of drugs encapsulated in
polymeric NPs have been reported for the oral (134), the parenteral (135,136), and
the ocular routes (137). However, the industrial constraints of solvent handling,
limited scale, and particular efforts needed to decrease residual solvent down to few
parts per million induced high manufacturing costs.
A clear aqueous nanodispersion of porpofol, a lipophilic anesthetic agent, was
developed by Chen et al. (138), which possessed physical and chemical stability. It had
better red blood cell compatibility and improved microbial resistance compared with
the marketed product diprivan that is an oil-based emulsion. They showed the new
nanodispersion using a combination of poloxamer, PEG 400, polysorbate 80, propylene glycol, and citric acid known as TPI 213 F (138). An interesting study of the in vitro
degradation of polymer poly-dl-lactic acid (PDLLA), poly-dl-lactic-co-glycolic acid
(PLGA), and polyethylene-oxide-based NPs showed that it took two years to degrade
the PDLLA-based NPs, whereas it took 10 weeks to degrade PLGA NPs (139).
Poly(methyl vinyl ether-co-maleic anhydride) (PVA/MA) is a biodegradable
poly-anhydride widely used for developing polymeric micelles as NPDDSs which
possess bioadhesive as well as mucoadhesive properties. Arbos et al. (140) reported
the application of these in the formulation of NPDDSs and showed that the bioadhesive properties of these NPs appear to modulate gastrointestinal transit profiles.
An interesting study was reported by Luu et al. (141) on developing polymeric
micelles of nanostructured DNA delivery scaffold by electrospinning of PLGA
(polylactide-co-glycolide) and PLA–PEG (poly(dl-lactide)-poly(ethylene glycol)).
They showed that the release of plasmid DNA was sustained over 20 days, and the
DNA released was structurally intact and capable of cell transfection and bioactivity. It was the first successful demonstration of plasmid DNA incorporation into a
polymer scaffold using electrospinning. Other groups have also used the nanosized
scaffold for drug delivery successfully for vascular endothelial growth factor (142),
for osteotropic factors (143), and for plasmid DNA (144).
A hydrotropic polymer system using N,N-diethyl nicotinamide was reported
to be useful for the NPDDS of the poorly water-soluble drug paclitaxel. The micelles
ranged between 30 and 50 nm and could be easily redissolved in an aqueous system.
They demonstrated higher loading capacity and physical stability compared with
other polymeric micelles (66).
Son et al. (130) have shown the accumulation of doxorubicin-loaded glycol
chitosan nanoaggregates in tumor cells by enhanced and permeation effects. Several
other studies also reported the accumulation of drugs in tumor cells in vivo using
NPDDSs (146,147). Aneed (148) has written an excellent overview of the current
drug-delivery systems used for cancer gene therapy, covering both the viral and
nonviral vectors for carrying the therapeutic genes.
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Polystyrene Nanospheres
These monodisperse polystyrene microparticles, also called latex microspheres, are
used in a wide range of immunodiagnostic assays, as size standards for calibration
of equipolymeric micellesent, in cell biology applications, and so on. The physical
adsorption of the polystyrene particles is used to bind ligands to the surface of the
particles. Sakuma et al. (149) studied the mucoadhesion of polystyrene NPs having
surface hydrophilic polymeric chains in the GI tract in rats. They reported that the
muco-adhesion of poly(N-isopropylacrylamide) NPs strongly enhanced the absorption of salmon calcitonin. Another interesting study showing applications of polystyrene NPs was reported by Hayakawa et al. (150). To establish an effective tool for the
prevention of HIV-1 transmission, lectin-immobilized polystyrene NPs were synthesized and examined for their HIV-1 capture activity. It showed that when concanavalin A was immobilized on the surface of polystyrene NPs (mean diameters of 400 nm)
with poly(methacrylic acid) branches and incubated with HIV-1 suspension at room
temperature for 60 minutes, the NPs achieved >3.3log and a 2.2log reduction of viral
infectivity in HIV-1 suspension at a concentration of 2 and 0.5 mg/ml, respectively,
demonstrating the potential of this technique for prevention of viral transmission
(150). Similar studies were also reported by Akashi et al. (151). Ogawara et al. (152)
wrote a review on hepatic disposition of polystyrene NPs and the implications for
rational design of particulate drug carriers. The clearance of colloidal particles from
the blood circulation occurs by phagocytosis and/or endothelial cells, mainly in the
liver, spleen, and bone marrow. The relative distribution of the injected particles in
these organs is known to depend on various factors such as the size and the surface
properties of the particles, and the type of serum proteins adsorbed onto the surface
of the particles. The basic principles behind their distribution characteristics into the
RES, however, remain unclear (152). An interesting study reported by Ogawara et al.
(153) showed that precoating with serum albumin has reduced the receptormediated hepatic disposition of polystyrene nanospheres. This technique can be
used to prevent rapid clearance by the mononuclear phagocyte system in vivo.
SOME COMMERCIALLY AVAILABLE NANOPARTICLES
Melamine Nanospheres
The melamine (polymethylenemelamine) nanospheres and microspheres are made
from cross-linked melamine and have some advantages depending on the application compared with polystyrene particles. They have a higher density (1.51 g/cm3),
are very stable, can be stored indefinitely, can be resuspended in water, do not swell
or shrink in most organic solvents, and are heat resistant up to 300° C. These monodisperse (CV 1–2%) melamine microparticles are hydrophilic and can be suspended
in water and their refractive index is 1.68. The surface of plain melamine microparticles is terminated with methylol groups, which could be readily functionalized in
the desired manner (154).
Plain Polymethyl Methacrylate and Biodegradable
Polylactide Nanospheres
Plain polymethyl methacrylate particles are available as 10% suspension, when
higher concentrations for production are necessary (154). Polylactide (PLA) is a biodegradable thermoplastic derived from lactic acid. It resembles clear polystyrene,
provides good esthetics (gloss and clarity), but it is stiff and brittle and needs modifications for most practical applications (i.e., plasticizers increase its flexibility).
Nanoparticulate Drug-Delivery Systems
19
These particles are made from PLA with a density of 1.02 g/cm3. They are supplied
as 1% aqueous solution (10 mg/ml) and are stable at a neutral pH for at least three
months. Degradation starts through basic or acidic pH or enzymatic hydrolysis.
Magnetic Plain Dextran Nanospheres
The super paramagnetic NPs on the basis of dextran with a size of 250 nm have a
magnetite content of 90%. A permanent magnet can easily separate 50 and 100 nm
particles from 130 and 250 nm particles. Such smaller sizes can only be separated by
a “high gradient magnetic field” device.
Gold Nanospheres
Gold particles are of highest quality and can be used in the production of diagnostic
tests as well as conjugation studies of proteins and antibodies. The particles have a
very narrow size distribution (CV between 5% and 15% depending on size) and are
available from 2 to 250 nm. The number of particles/ml is given in the product/
ordering table. The solutions are stabilized with HAuCl4. Gold and silver colloids or
sols are available in a number of different sizes. There are 14 different gold colloid
sizes and are offered in four packing sizes. The products are best stored at room
temperature, although storage at 4°C is an option. However, temperatures too close
to freezing will destabilize the sol, causing aggregation and product loss (154).
Silver Nanospheres
Silver nanospheres are of highest quality and can be used in the production of diagnostic tests as well as conjugation studies of proteins and antibodies. The particles
have a very narrow size distribution (CV between 10% and 20% depending on size)
and are available from 2 to 250 nm.
Silica Nanospheres
These mono-disperse silica particles with a density of 2.0g/cm3 are simple to
dispense and to separate. Although polystyrene particles (d = 1.04) are difficult to
separate by centrifugation under a size of 500 nm, silica particles do set down easily
and are easy to resuspend. The silica particles are stable in water and organic
solvents, produced under a new dying method. Silica particles are easy to functionalize and available as fluorescent particles. They are useful for coupling of DNA,
oligonucleotides, oligopeptides, proteins, lectins, and antibodies. The silica particles
are also available with different functional groups as –NH2, and –COOH, albumin,
protein A, epoxy, NHS, NTA, and EDTA (154). Li et al. (155) have prepared and
characterized porous hollow silica NPs for controlled release applications. They
reported a novel method for preparing hollow silica nanospheres with a porous
shell structure via the sol–gel route and using inorganic calcium carbonate NPs as a
template with 100-nm diameter and a wall thickness 10 nm of the nanospheres.
Several factors were found to affect the drug release rate from the nanospheres (155).
Alumina Nanospheres
Alumina nanospheres and microspheres have been used in various applications
because of their size uniformity and high degree of spherical particles (as a result,
high flowability and high packing density). Properties of alumina, such as high
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thermal conductivity, heat resistance, hardness, and so on, are useful in formulation.
These particles are often used as fillers for thermal conductivity sheets. Alumina
nanospheres are available with plain, amino, and carboxyl functional groups. The
alumina microspheres and nanospheres can be prepared with many functional
groups such as albumin, protein A, epoxy, NHS, NTA, EDTA, and many others.
They are useful for coupling of DNA, oligonucleotides, oligopeptides, proteins,
lectins, and antibodies.
Carbon Nanotubes
In the last 15 years, it has been an exciting time for the field of carbon nanomaterials.
The discoveries of fullerenes and carbon nanotubes have attracted the attention of
many researchers all over the world (156). Iijima (157) first discovered these in 1991.
They are now commercially available and can be manufactured on large scales (158).
There are several companies who manufacture the carbon nanotubes and gradually
their prices are reducing. They have a great potential in the drug-delivery systems
and many other applications (159).
DIVERSE AND EMERGING TRENDS IN NANOTECHNOLOGY
APPLICATIONS
Nanotechnology is making a great impact in many areas and some of the major
areas are as follows.
Biological Analysis and Discovery
The basic science behind identifying the presence of a particular gene or protein has
been developing for some time. The introduction of nanofluidics offers an excellent
opportunity to work with smaller amounts of material. It can be used to segregate
proteins and nucleic acids (DNA and RNA) based on size and shape. Nanomembranes
offer a great potential in this area (160). There are several groups working on the
idea of passing a single DNA or RNA thread through a nanosized pore, forcing it to
straighten out and pass through a base part of fundamental coding element of
nucleic acid (161,162). Changing electrical gradients on either side of the structure
containing the pore, or quantum tunneling current across the pore, could be used to
identify the particular base that is passing through. The ability to sequence a whole
genome, the sum total of genes in an organism in a matter of hours, has been proposed as a potential application of this approach. The impact of these techniques in
formulation of drug-delivery systems and their therapeutic applications will be
worth watching during the coming few years.
Nanoparticles Tagging
Another boon to analysis will likely come from the attaching of NPs to molecules of
interest. NPs small enough to behave as quantum dots can be made to emit light at
varying frequencies. These attached molecules can be spectroscopically measured and
will allow many different molecules to be measured from the same single sample.
Another similar approach relies on getting the molecules to bind nanowires that
have stripes on them similar to bar codes. These detection technologies are getting
closer to commercialization. These can also further be applicable in the development
of drug-delivery systems (163).
Nanoparticulate Drug-Delivery Systems
21
Nanostructured Materials
Nanostructured materials coupled with liquid crystals and chemical receptors offer
the possibility of cheap, portable biodetectors that might, for instance, be worn as a
badge. Such a badge could change color in the presence of a variety of chemicals
and would have applications in hazardous environments (163). These can also be
explored as NPDDSs.
Single-Molecule Detection
There are methods to detect a single photon. If one can make a nanostructure such
as a quantum dot which will emit a photon in the presence of a particular molecule,
one can make a device that can detect a single molecule of a substance. Various
groups are working in this direction. The key commercial aspects of the use of microand nanotechnology in biodetection relate to portability, cost, and sensitivity (164).
Protective Nanoparticles Against Pathogens
NPs are disruptive to bacteria and viruses simply by virtue of their physical nature.
This has led to ideas of lacing fabrics in hospitals with such NPs or nanoparticulate
creams that can be spread on the body or sprays that can be inhaled, protecting
against various pathogens (163).
Nanotubes and Cellular Manipulation
Nanotubes and cellular manipulation hold a great promise of being useful in biological applications and as NPDDSs. There are tubes small enough to suck out a
nucleus from a cell and place it into another. This is the technique behind cloning
but nanotubes are finer still and offer the potential of making probes and delivery
mechanisms that can be even more precise (165,166).
Nanoengineered Prosthetics
There are several devices which can be plugged into parts of central nervous systems (CNSs), designed for processing visual or auditory information. Nanotechnology
is changing all that: there is a potential for replacing a few of our organs for seeing,
hearing, and touch, although connecting to CNS and avoiding classical problems of
rejection will be a challenge. Synergies between our application areas, materials
with greatly improved strength and designer surface properties offer potential for
use in all manner of implants, from artificial hearts to hip implants. All these will
have excellent potential in NPDDSs in the near future (163).
Thiomer Nanoparticles
Thiolated polymers designated as thiomers are thiol side-chain-bearing polymers.
Owing to the introduction of these functional groups, thiomers exhibit comparatively
stronger mucoadhesive, cohesive, enzyme inhibitory, and permeation-enhancing
properties. They have excellent potential in the formulation of the NPDDSs. Several
attempts are going on developing NPDDSs for insulin and calcitonin using thiomer
NPs. A nasal delivery system for human growth hormone, a pulmonary delivery
system for vasoactive intestinal peptide, a noninvasive delivery system for the KLH
antigen, and nasal delivery systems for amyloid-binding peptides are being generated and researched using thiomer polymeric NPs (154,167–169).
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Nanostructured Monoliths
Functional nanostructured monoliths are being developed and optimized for the
analysis of protein and peptides by nano-HPLC–MS/MS. Ring-opening metathesis
polymerization (ROMP) is employed to develop nanostructured separation media.
ROMP is a living polymerization technique capable of controlling polymerization at
a molecular level. These show a better performance of protein and peptide analysis
using nanocoupling techniques. Demand for nanoseparation phases is increasing
heavily as nanocoupling techniques, particularly nano-HPLC-MS, are becoming
more important.
Antibody-Coated Nanospheres
The 0.7 to 0.9 µm goat antimouse IgG-coated polystyrene particles are prepared by
using passive adsorption; the coated particles are stable for several years under
proper storage conditions. Antibody-coated polystyrene particles have antibody
contents 14 µg/mg solid, 0.2 µg/cm2, 1.5 × 104, IgG/particle, and binding capacity
to IgG-FITC 4 µg/mg. These can be used for targeted delivery of drugs (170).
Nanocrystallites
An interesting application of nanotechnology was reported by Yin et al. (171) where
they used pluronic in spray-drying to produce the nanocrystallites of the drug needing improved bioavailability or improved solubility. They utilized Pluronic F127
and developed several combinations of drug with pluronic, and spray-dried these
combinations using the organic solvent.
Nanohybrids
A novel method of delivering nonionic poorly water-soluble drugs such as
camptothecin was developed and reported by Tyner et al. (172). Camptothecin was
first incorporated into micelles derived from negatively charged surfactants. The
negatively charged micelles were then encapsulated in NPs of magnesium–
aluminum hydroxides by an ion-exchange process. The resulting structures were
termed nanohybrids. They released camptothecin rapidly with complete release
within 10 minutes at both 4.8 and 7.2 pH. The encapsulation process allowed almost
three times the drug loading and has excellent potential for drug-delivery applications. This complex was shown to provide similar cytotoxic characteristics to naked
drug but the nanohybrids can be administered in a dose-controlled fashion due to
good dispersion of the complexes in water. The ability to attach targeting biologically active molecules to the outside surface of the nanohybrids as well as the potential controlled release properties of the complexes indicate that these hybrids can be
used for specific delivery of poorly soluble nonionic drugs (172).
Another interesting study from the same group was reported for the application of the nanohybrids as a nonviral vector for gene delivery. The nanohybrids
were synthesized by the intercalation of a full gene and promoter encoding green
fluorescent protein between layers of an inorganic host. The nanohybrids were
delivered to 9L Glioma cells, JEG3 Choriocarcinoma placental cells, and cardiac
myocytes. All cells were able to internalize and tolerate the nanohybrids and expressed
the gene with some cell lines having up to 90% transfection efficiency (173). The
nanohybrids mimic key features of viral delivery systems. Unlike other reported
inorganic vectors, the host encapsulates the DNA molecule between the inorganic
Nanoparticulate Drug-Delivery Systems
23
layers, protecting it from premature recognition or degradation. In addition, the
presence of hydroxyl group on the host surface provides the means to link biologically active molecules to the exterior surface of the nanohybrids, suggesting the
possibility of targeting the drug-delivery systems. Other advantages of nanohybrid
systems include less foreign DNA and the ability to deliver multiple genes to cells.
In addition, use of the nanohybrids is expected to increase control of the expression
level of target genes by regulating the amount of DNA introduced into the cells.
Such dose control is difficult to achieve in viral vectors. Preparation of viral constructs can be time-consuming and tedious. As the nanohybrids are applicable to a
wide range of genes, they may be prepared through a simple one-step process, thus
greatly shortening the amount of time needed to produce a functional vector (173).
Nanocontainer Technology
Nanotechnology promises new avenues to medical diagnosis, treatment, and drug
delivery. In this respect, there are some special injectable nanovehicles that are programmable towards specific targets and are able to evade the immune defense and
are versatile enough to be suited as carriers of complex functionality. Biotinfunctionalized (poly-(2-methyloxazoline)-b-poly(dimethyl siloxane)-b-poly(2-methyl
oxazoline) triblock copolymers were self-assembled to form nanocontainers, and
biotinylated targeting ligands were attached by using Streptavidin as a coupling
agent. Specifically, fluorescence-labeled nanocontainers were targeted against the
scavenger receptor A1 from macrophages, an important cell in human disease. In
human and transgenic cell lines and in mixed cultures, receptor binding of these
generic carriers was followed by vascular uptake. Low nonspecific binding supported the stealth properties of the carrier while the cytotoxicity was absent. These
versatile carriers appear to be promising for diagnostic and therapeutic drugdelivery purposes (174).
Electrospun Nanofibers as Drug-Delivery Systems
Electrostatic spinning is a versatile technique applied to various micro- and
nanofabrication areas using numerous polymers. The technique involves jetting a
liquid stream of a drug/polymer solution to a potential between 5 and 30 kV;
fibers at submicrometer diameters can be formed when electrical forces overcome
the surface tension of the drug/polymer solution at the air interface (termed a
Taylor cone) such that a jet forms (175). As the jet accelerating through the electric
field, two possible outcomes have been hypothesized: (i) radial forces become
increasingly important leaning to splaying of the solution stream or (ii) a continuous single filament is generated based on bending instability (176). As the solvent
evaporates, fiber can be collected on the screen to give a nonwoven fabric or collected on a spinning mandrill. The fiber diameter is a function of solution surface
tension, dielectric constant of polymer solution, feeding rate, and electrostatic field.
The nature of the polymer can also direct the use of the electrospun fibers, with
water-soluble polymers giving rise to immediate release dosage forms and water
insoluble giving sustained release systems. A report by Verreck et al. (177) has
used these nanofibers for the release of itraconazole. The fiber diameter ranged
from 500 nm to a few micrometers. They showed that complete release of the
poorly soluble drug can be achieved and the rate of drug release can be tailored,
showing applications of these nanofibers for wound healing, buccal, and topical
applications (178).
24
Thassu et al.
FUTURE DIRECTIONS
The majority of commercial NP applications in medicine are geared towards drug
delivery. In biosciences, NPs are replacing organic dyes in applications that require
high photostability as well as high multiplexing capabilities. There are some new
developments in directing and controlling the functions of nanoprobes, for example, driving magnetic NPs to the tumor and then making them either release the
drug load or just heating them in order to destroy the surrounding tissue. The major
trend in further development of nanomaterials is to make them multifunctional and
controllable by external signals or by local environment, thus essentially turning
them into nanodevices (164).
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2
Nanosuspensions for Parenteral Delivery
Barrett E. Rabinow
Baxter Healthcare Corporation, Round Lake, Illinois, U.S.A.
HISTORICAL INTRODUCTION
Need
The need for nanosuspensions as a dosage form was recognized as a means to
administer therapeutic quantities of water-insoluble dosage forms (1). The continuing need for such a tool was reinforced with the uptake of high-throughput
receptor-based screening assays employed by the pharmaceutical industry during
the 1990s. This technique searches for drugs that exhibit strong binding to hydrophobic target receptor pockets and therefore produces drug leads that tend to be
poorly water soluble (2). It was not only the kind of molecule that changed, but also
the vast numbers of such drug leads that emerged from discovery efforts, that
compelled a solution to the resulting formulation conundrum. Typically, these in
vitro assays were conducted in dimethyl sulfoxide, to obviate problems of water
insolubility during the preliminary discovery phase. With progression to animal
studies, however, nontoxic vehicles were sought that would permit assessment of
the toxicity of the drug candidate itself, without interfering effects due to the vehicle
(3,4). Nanosuspensions commended themselves as suitable candidates because:
(i) solvents were not needed, (ii) small particulate size permitted intravenous
delivery, and (iii) the solid crystal phase permitted high loading, which permitted
the animal studies to be conducted at many multiples of the intended dose in man.
Predecessor Technology
Equally important as the demand of applications was the ready availability of
building blocks for implementation of the new technology. Homogenizers had been
in use since the late nineteenth century for homogenization of milk, and more
recently in the latter half of the twentieth century for commercial production of
intravenous lipid emulsions. Therefore, such issues of cleanability, sterilizability,
noncontamination of fluid processing streams with hydraulics, and so on, had been
long resolved. It remained for more efficient valve designs to be developed, transmitting more of the energy to the particle (5). Similarly, milling equipment for particulates, as, for example, the paint industry, was available and amenable to the
changes necessary for pharmaceutical use. This was accomplished by developing
cross-linked polystyrene grinding media that was smaller and relatively noncontaminating, resulting in smaller drug particles (6). Precipitation and crystallization
techniques were much older, but significant chemical engineering understanding
and optimization occurred during the last century (7). Supercritical fluid processing
(8) was commercially developed relatively recently. Substantial understanding of the
role of surfactants, both for formulation stabilization (9) and for impacting pharmacokinetics (10), became clarified. Here, lessons were learned from earlier developed
drug-delivery platforms, liposomes and emulsions, regarding prolonging circulation
times of particulate dosage forms.
33
34
Rabinow
Combination of Component Technologies
The component technologies described above were often combined to improve
performance and overcome individual deficiencies, in the development of nanosuspensions. Thus, nanosuspensions were lyophilized for greater stability (11), formulated with soluble molecular analogs (12), applied to liquid drugs to form
emulsions (13), and combined in matrix pellet formulations (14) and multivesicular
lipid systems (15). To reduce particle size to the nanometer range, supercritical fluid
techniques were combined with an ultrasonic vibrating surface to enhance mixing
and atomize the jet into nanodroplets (16), or sprayed into quenching surfactant
solutions (17). Taxol nanosuspensions were prepared by emulsion templating, that
is dissolving drug into a volatile solvent, homogenizing in an aqueous solution containing suitable stabilizers, and evaporating solvent to recover nanosuspensions (18).
Finally, rapid precipitation and homogenization were synergistically combined to
obtain stable nanosuspensions (19). Homogenization cracks crystals along their
defect planes, induced by rapid precipitation, to reduce their size further. Additionally,
the mechanical shock of homogenization often induces conversion of the unstable,
initially formed polymorph resulting from precipitation, to the more thermodynamically stable polymorph.
FORMULATION APPROACHES AND MANUFACTURING METHODS
Strategy
The preparation of stable nanosuspensions must recognize the thermodynamic forces
at work. For a given mass of drug substance, as particle size is reduced, surface area
is increased. In an aqueous medium, this significantly increases the surface free
energy of the drug system. Strong, and therefore stable bonds within the drug crystal lattice, on the one hand, and intermolecular hydrogen bonds of water molecules,
on the other hand, are disrupted. Instead, they are replaced by a large interfacial
area of hydrophilic water molecules in proximity with a hydrophobic drug surface.
Such a system will tend to reduce the energetically unfavorable area by particle
growth and by aggregation. Ostwald ripening represents one mechanism by which
this may occur. By the Ostwald–Freundlich equation, smaller particles have a higher
surface energy than larger particles (20). This leads to greater dissolution of smaller
particles with consequent increasing size of larger particles. As a result, the distribution of the suspension shifts to increasing particle size. To address this as well as
irreversible agglomeration, formulation strategy is designed to stabilize particle
size over time.
Utilization of surfactant excipients is an essential part of formulation strategy.
The particular surfactants are selected to be compatible with both phases, and so
interpose themselves between the hydrophobic drug surface on one side and the
aqueous media on the other. The effectiveness of their interaction is measured by
the reduction in surface free energy. An ionically charged surfactant will provide
electrostatic repulsion of neighboring particles, which will tend to inhibit aggregation. However, if the particles overcome this long-range r −2 repulsion (where r is the
interparticulate distance), as by addition of a salt to increase dielectric shielding,
they may still approach each other too closely. They may thus be subject to the shortacting (r −6), but much stronger, attractive London dispersion forces resulting from
instantaneous dipole polarization interactions (21). Under this circumstance, the
particles will aggregate. To forestall this likely possibility, a second type of surfactant, designed to prevent too close contact by nonelectrostatic means, is also used in
Nanosuspensions for Parenteral Delivery
35
combination with an ionic surfactant. This may be a neutral, block copolymer type
of molecule having hydrophobic and hydrophilic domains, for example, poloxamers. As particles bearing such surfactants approach each other, the movement of
their polymeric chains becomes constrained, entailing a loss of entropy. This steric
repulsion therefore is used synergistically with ionic repulsion to minimize
aggregation over time.
Precipitation
Precipitation is accomplished by dissolving the drug in a suitable solvent and then
mixing with a nonsolvent, effecting supersaturation with resultant precipitation.
This is a two-part process: initially, solid-phase nuclei must form from molecules in
solution, and subsequently crystals will grow. Conditions such as choice of solvent,
temperature, order, and speed of addition are chosen to achieve different goals. To
prepare drug crystals as small as possible, rapid creation of many nuclei is optimized while minimizing subsequent growth. To accomplish this, the drug solution
may be added rapidly to the nonsolvent, effecting nearly immediate supersaturation and growth of many nuclei (22). Because of the preponderance of nonsolvent
relative to drug solution, growth is limited. This order of addition is opposite to
what is typically accomplished in pharmaceutical technology, where the goal is to
achieve large, pure crystals by slower growth.
Despite attempts at optimization, it is found that precipitation alone is unable
to generate suitable particles. The particle size is usually too large for injectable
applications, and the distribution tends to be nonuniform. Furthermore, because of
the speed of formation, such suspensions are typically not thermodynamically
stable and tend to evolve over time. The metastable distribution may tend to grow
in size if crystalline, or evolve to a crystalline form if formed initially in the amorphous state. If crystalline, rapidly formed precipitates also tend to have many lattice
defects, and needle-like morphology.
Homogenization
Piston gap homogenization had been used for comminution of liquid-phase systems,
such as emulsions and liposomes. Haynes (1) first described the use of the technique
for size reduction of solid crystalline drugs in the presence of surfactant. The technique currently involves passage of a drug suspension under high pressure (5000–
50,000 psi) through a variably adjusted, narrow gap. As the velocity of the fluid
must increase through the stricture, the pressure drops in accordance with Bernoulli’s
principle. The pressure drop causes water vapor bubbles to form in the gap, but
these subsequently collapse as the fluid enters a higher-pressure area. The rapid
pressure collapse causes cavitations, which provides the energy to crack the particles. The greatest size reduction occurs in the first few passes of the suspension
through the gap, with diminishing benefit thereafter. This may result from initial
cracking of the crystals along defect planes, which become exhausted as the crystals
become smaller. It is observed that increasing pressure usually decreases the ultimate particulate size. Size reduction of the suspension, as by jet milling, and prewetting of the slurry with surfactant are helpful ancillary techniques that can be
used (23). The mean size of the particles attainable with this approach is typically
300 nm to 2 µm. Such suspensions may be terminally sterilized by autoclaving if the
drug withstands heat. If the drug is heat-sensitive, aseptic production starting from
sterile raw materials is possible.
36
Rabinow
Combined Precipitation/Homogenization
Precipitation may be combined with homogenization to overcome disadvantages
associated with either technique alone. Heat-sensitive drugs may be dissolved in
solvent, which is then aseptically filtered. Microprecipitation downstream of the
filter will provide sterile crystals which may then be passed through a homogenizer.
It is found that combining both techniques produces particles that are significantly
smaller than can be attained by either technique alone. Furthermore, with precipitation alone, it often happens that the initially obtained precipitate proves to be
unstable, undergoing further change. This may involve an amorphous to crystalline
conversion or simply crystal growth over time. It is found, however, that subsequent
homogenization arrests this growth by providing the energy necessary to overcome
the activation barrier between the initially formed precipitate and the thermodynamically stable material.
IN VITRO AND IN VIVO PREDICTION OF STABILITY
AND PHARMACOKINETICS
Formulation Selection
Rational decision-making for formulation development of a poorly water-soluble
compound proceeds sequentially by determining the most appropriate method for
the compound. If the simpler approaches of pH adjustment, salt preparation, use of
cosolvents, or use of lipids prove ineffective, then nanosuspensions may be considered. This approach is reserved for the most difficult compounds which have high
crystal energy (high melting temperature) (24), in addition to high log P, as well as
a requirement for high loading. For these compounds, successful formulation with
nanosuspensions can be expected, provided the appropriate surfactant package is
utilized, and sterilization issues can be addressed.
Insoluble (stable)/
slow crystallizer (many
Too soluble/
slow crystallizer
defects/smaller particle size)
(many defects/smaller
particle size)
Insoluble (stable)/
fast crystallizer (few
Too soluble/
fast crystallizer
defects/larger particle size)
Nucleation rate
Log (Free Energy, Ergs)
Molecular Determinants of Particle Size
Prior to embarking upon a specific nanosuspension program, the strategy for successful formulation can be determined by evaluation of the molecular structure
with regard to Figure 1. The log of the water solubility (measured or calculated,
(few
defects/larger particle size)
Ostwald ripening
Log (Solubility, M)
FIGURE 1 Planning formulation strategy of nanosuspensions based on molecular flexibility and
solubility. Suitable candidates are insoluble and stable to growth via Ostwald ripening. Additionally,
the nucleation rate is a measure of the defect number of the crystals, which determines friability and
ability to form smaller particles.
Nanosuspensions for Parenteral Delivery
37
with regard only to the structure, without need for melting point) is plotted along
the horizontal axis. Too high a water solubility will lead to Ostwald ripening
(increase of the particle size of the population). Along the vertical axis is plotted
entropy, a measure of the degrees of freedom of the molecule, as an indicator of the
disorder likely to be in the crystal. Crystal defects are an important consideration
both where nanosuspensions are manufactured by growth, for example, crystallization, as well as by attrition, by milling, or homogenization. In the case of crystallization processes, it is the smaller, more rigid molecules, with less entropy to lose,
that fit more readily, and therefore rapidly, into the growing crystal lattice (21). This
could lead to a too rapid crystallization rate, resulting in particles that are excessively large. For attrition processes, on the other hand, it is known that impact force
cracks crystals along their defect planes (25–28); smaller particles will result where
there are many defects. The molecular gas-phase entropy, as calculated by vibrational mode analysis of the structure, is indicative of the defect rate and is plotted
along the vertical axis. For molecules that are predicted to be too soluble (greater
than several hundred ppm solubility), other formulation approaches may be more
appropriate. If nanosuspensions must be used, then lyophilization should be considered for stability. Particular surfactant packages and more attention to processing
conditions should be anticipated for molecules that crystallize quickly, to ensure
successful formulation (29).
In Vitro Dissolution
It is also advantageous to be able to predict the in vivo pharmacokinetic behavior of
the suspension. This can be done experimentally after the formulation has been
made by injecting a bolus of the nanosuspension into a solution and plotting the
resulting light transmittance. Physical persistence of particles will decrease transmittance, indicating slow solubility rate. Besides the extreme examples of compounds that do not dissolve at all, and those that dissolve virtually instantaneously,
are those whose dissolution rate is dependent upon particle size or infusion rate.
A plasma-simulating solvent may better approximate the in vivo case. It is found
that the dissolution rate can also be calculated a priori by knowing only the surfactants used in the formulation and the structure of the molecule, by utilizing the
Stokes–Einstein and Ostwald–Freundlich equations (29,30). Thus, before undertaking development, the in vivo pharmacokinetics can be estimated to determine the
suitability of the approach.
Pharmacokinetic Profiles
Possible pharmacokinetic profiles of injectable drugs are shown in Figure 2. Based
upon in vitro dissolution, a fast or slow-dissolving nanosuspension may be indicated. A fast dissolving formulation will yield a pharmacokinetic profile essentially
identical to an a priori solution formulation (prepared with, e.g., high solvent level
or extreme pH) of the drug (31,32). As a result, the tissue distribution will also be
equivalent, as in the case for flurbiprofen (33). In contrast, the nanoparticles of a
slow-dissolving formulation may be phagocytized first by the fixed macrophages of
the liver and spleen (34). As they are subsequently dissolved within the cell, the
drug will be slowly released. The resulting IV depot effect will yield a relatively low
Cmax, an apparent dip in the plasma curve within approximately 30 minutes of injection, followed by a distinct rise in the plasma concentration, peaking at around six hours
for the rat, followed by a prolonged tail-off lasting over one hundred hours (Fig. 2A).
38
Rabinow
PK: Rat vs. Dog
(C)
Nanosuspension/ Drug
Metabolite
0.6
20
40
60
80
100
Time (hours)
120
140
Rat
0.2
0
0
Solution/Drug
Metabolite
0
Dog
0.4
10
Whole blood
Plasma
Time (h)
20
30
40
50
60
70
80
Time (h)
(B)
Drug Concentration (g/mL)
Drug Concentration (ug/mL)
(A)
Nanosuspension/Drug
Drug (µg/mL)
Drug Concentration
0.8
Solution/Drug
(D)
Nanoparticles
Microparticles
Time (h)
FIGURE 2 Schematic in vivo pharmacokinetic (PK) profiles of nanosuspensions. (A) Comparison
of pharmacokinetic profiles for slow-dissolving nanosuspension versus solution formulations for both
parent compound and metabolite. (B) Comparison of PK profile of slow-dissolving nanosuspension
for rat versus dog. (C) Comparison of PK profile of slow-dissolving nanosuspension in plasma
versus whole blood. (D) Comparison of plasma PK profiles of nanoparticles versus microparticles,
administered subcutaneously.
The half-life is noticeably longer in the dog (Fig. 2B). Where metabolites are also
tracked, it is observed that their growth is delayed, and their elimination is prolonged (Fig. 2A). In comparison with microparticulate suspensions administered
subcutaneously, nanosuspensions may demonstrate a higher Cmax and AUC (Fig. 2D).
This occurs because dissolution of injected particles in a depot site often constitutes
the rate-determining step for migration into the blood compartment. As particle size
is reduced, surface area is increased, which increases dissolution rate. Often, appreciably higher drug concentrations appear in a whole blood assay as compared with
plasma (Fig. 2C), assuming that the density of the nanosuspension is sufficiently
greater than that of plasma. This may suggest red cell binding, but could also be
indicative of uptake by circulating macrophages.
SAFETY OF INJECTABLE NANOSUSPENSIONS
The specification of safe levels of nanoparticulate suspensions that may be administered intravenously may be defined rationally in view of a wide database that has
by now been accumulated. To place this into perspective, 40 years ago concerns were
expressed about particulate matter in intravenous solutions, where extreme
examples with dire consequences were exaggerated for that relatively new dosage
form (35). Safety of injected particulates will be considered from two perspectives:
(i) potential vascular occlusion as a function of the size, number, and composition of
the particles and (ii) monocyte phagocytic system response.
Nanosuspensions for Parenteral Delivery
39
In Vivo Distribution as a Function of Particle Size
Pharmacokinetics of organ distribution is dependent on particle size and rate of
infusion. Nonmetabolizable particles larger than 7 µm are trapped in the pulmonary vasculature for extensive periods of time. In the lung, alveolar macrophages
provide a mechanism for passing particles less than 12 µm (36) through the capillary
walls permitting excretion into the sputum out of the lung. The extensive collateral
circulation of the pulmonary vasculature appears to mitigate the potential blockage
of capillaries by particles, with anticipated reduction of blood flow, if the particle
load is kept sufficiently low (37).
However, there is evidence of capillary occlusion in the lungs of recipients of
transfusion of unfiltered blood, which can contain particles of 20 to 500 µm in size.
This effect was eliminated with the use of Dacron® wool depth filters of 40 to 80 µm.
This suggests both an approach to deal with high particulate burdens of therapeutic
nanosuspension dosage forms, as well as acceptability of particles < 40 µm from
such products (37).
If not dissolved initially, particles smaller than 7 µm escape from their initial
lung sequestration rather quickly (38), within minutes, and undergo phagocytosis
by the fixed macrophage cells of the liver and spleen (39,40). This is a normal behavior of these cells when presented with microbes and foreign material of size less
than about 8 µm. In several rat studies, no evidence of an inflammatory reaction
was found. Histologically, a low incidence of focal myocardial degeneration was
found with 10 and 40 µm particles. Apparently safe levels of 8 × 106 particles/kg of
size 0.4 to 10 µm or 4 × 105 particles/kg of particles of size 40 µm could be
administered.
By way of comparison, Optison™ (Amersham Health) is an approved albumin
microsphere suspension for echocardiographic imaging. Although the residence
time in the body is very short because of the ultrasound-induced disruption of the
particles, there is a load of smaller particles resulting from disintegration of the primary particles that must be cleared by the monocyte phagocytic system (MPS). The
particle size mean diameter is 2.0 to 4.5 µm, with 93% less than 10 µm, but with a
range extending upward to 32 µm. The concentration is 5 × 108 to 8 × 108 microspheres/mL, with a maximal recommended dose of 8.7 mL per contrast study (41).
The maximum number of particles that can be injected for this approved product is
therefore 7 × 109, or 9.9 × 107 kg−1.
Additionally, macroaggregated albumin injection is an approved product,
routinely used in diagnostic imaging. The aggregated particles are formed by denaturing human albumin in a heating and aggregation process. Each vial contains four
to eight million particles. By light microscopy, more than 90% of the particles are
between 10 and 70 µm, whereas the typical average size is 20 to 40 µm; none is
greater than 150 µm. The suggested range of particle numbers for a single injection
is 200,000 to 700,000 with the recommended number being approximately 350,000
(42). By way of comparison, the USP <788> microscopic test for particulate matter
in small-volume parenteral intravenous solutions permits 300 particles >25 µm.
Therefore, conformance of IV drug nanosuspensions to the limits contained within
the USP <788> standard will ensure significant safety factors relative to the current
practice of pulmonary perfusion of radiographic particulate injections.
The above analysis addresses the high size end of the particle size distribution
of a nanosuspension drug formulation. The mean value is far smaller, below 1 µm.
Table 1 summarizes maximal IV levels of tolerated doses, reported in the
literature.
40
Rabinow
TABLE 1 Maximal Levels of Injected Particles with Outcomes
Particle size (μm)
Protocol (particle dose/kg)
1.3
0.5–1.17
0.4,4,10
3.4
3.7
3.4
2.0–4.5
0.4
Bolus, 6 ×
Bolus, 1.6 × 1012
Rats, bolus, 8 × 106
Dogs, bolus, 1 × 1010
Dogs, repet. bolus, 2.4 × 108
Dogs, 2 min bolus, 8.9 × 107
Humans, bolus, 9.9 × 107
Rats, bolus, 2.5 × 1012
0.4
Dogs, infusion, 1.3 × 1012
109
Outcome
PK study (43)
PK study (45)
Well tolerated (39)
Well tolerated (48)
Well tolerated (40)
Well tolerated (36)
Optison, approved product (41)
Well tolerated (Jerome Gass, Baxter
Healthcare, personal communication)
Well tolerated (Jerome Gass, Baxter
Healthcare, personal communication)
Response of the Monocyte Phagocytic System
The majority of the animal studies performed in the literature involved inert, nonmetabolizable polystyrene, cross-linked styrene divinyl-benzene, or latex microspheres. A metabolizable drug nanoparticulate will be processed through the
phagolysosomes of the macrophages much faster than will inert particles (51). This
poses much less burden on the macrophages and enables them to cycle faster. If the
reticuloendothelial system becomes overloaded by phagocytic activity, then reticuloendothelial blockage could occur (52), but only if the phagocytic overload is
continued and heavy (53,54), because these cells can digest all biodegradable
substances (55). Clinically, administration of liposomal doxorubicin (a cytotoxic
agent and macrophage targeter) did not result in more frequent opportunistic infections in patients with AIDS-related Kaposi’s sarcoma compared to patients treated
with combinations of doxorubicin, bleomycin, and vincristine (56). This probably
results from the compensatory increase in macrophage numbers and activity when
subjected to high phagocytic loads (57).
APPLICATIONS OF PARENTERAL FORMULATIONS
Regional Anesthetics
Pursuing a local or regional anesthetic agent for post-op or chronic pain, Boedeker
et al. (58) investigated the anesthetic effect of a lecithin-coated tetracaine-HI nanosuspension. It was produced by sonication, having 80% of the particle size distribution between 100 and 500 nm (by Coulter counter) and no particles >5 µm. About
0.3 cm3 of sample was infiltrated subcutaneously in rats’ tails, and local anesthesia
was detected with a hemostat. Animals served as their own controls by testing both
proximally and distally to the injection site. Rats injected with 10% lecithin-coated
tetracaine nanocrystals showed a tail block distal to the injection site lasting 43.4 ± 1.2
hours (n = 9), whereas those injected with 1% tetracaine solution had a mean tail
block time of 8.5 ± 1.8 hours (n = 5). Both these groups regained a positive response
to tail clamping distal to the injection site, thus ruling out nerve injury as the mechanism of local anesthesia. Those rats receiving 10% tetracaine solution either died
within 10 minutes or developed wet gangrene of the tail. Negative controls of lecithin membranes without drug and 5% dextrose showed no anesthetic effect. No
evidence of gross local tissue damage or systemic toxicity was observed with the
nanosuspensions. Comparison of inflammation scores, based on neutrophilic
accumulation, revealed no statistically significant difference in the level of tissue
Nanosuspensions for Parenteral Delivery
41
inflammation for the 10% nanosuspension versus 1% solution. All agents reached
the noninjected control level of inflammation by 14 days (59). The nanosuspension
thus provided sustained release without peaks in delivery of what would otherwise
be a toxic concentration of the drug.
Because of the potential for neurotoxicity from the nanosuspension depot if
placed proximal to nerves, a neurotoxicity study was conducted. Exposure of the rat
sciatic nerve, followed by extra fascicular administration of the test and control articles to the fascia surrounding the nerve, and reclosure of the wound was performed.
At 72 hours, the sciatic nerves were excised, fixed, stained, and scored microscopically for edema. There was no statistical difference between the 10% tetracaine
nanosuspension, the 1% tetracaine solution which is used clinically, or 5% dextrose.
All of the agents, however, caused statistically significant, although minimal, neural
edema when compared to noninjected nerve (60).
Intradermal toxicity of a lecithin-coated suspension of dezocine, prepared
similarly to that for tetracaine above, was evaluated in a rat model. It was compared
with that for Dalgan®, a commercial solution formulation of the analgesic, consisting
of 1.5% dezocine in a 30/70 propylene glycol/water mixture (61). Twenty rats were
studied in each of the four groups, receiving 0.05 mL ID injections onto the shaved
midback. Skin reactions were evaluated for 72 hours using the following criteria:
discoloration 0 to 2; blanching 0 to 2; ulceration 0 to 2; eschar formation 0 to 2, with
0 being no effect and 2 the most severe. A cytotoxic index was calculated by multiplying the total score above by the area of injury. The index for 250 µg microcrystal
dezocine at 6.0 ± 0.7 was significantly ( p < 0.01) less than that for 250 µg Dalgan
injection, which had an index of 24.2 ± 1.4. Negative control groups of lecithin membranes and 10% dextrose had still lower cytotoxicity indexes at 1.8 ± 0.3 and 0.6 ± 0.2,
respectively. Despite having less cytotoxicity, the microcrystal dezocine formulation
resulted in a significantly ( p < 0.01) extended plateau of analgesia of 334 ± 16.9
minutes versus 48 ± 7.5 minutes for Dalgan. Thus, the intensity and duration of
analgesia were prolonged while rendering the drug more tissue-compatible.
Intravenous
Malignant Hyperthermia
Phospholipid-coated nanosuspensions of dantrolene and sodium dantrolene were
studied for treatment of malignant hyperthermia (62). The currently available
dosage form, Dantrium® Intravenous, is a lyophilized formulation which must be
reconstituted slowly to produce a 0.33 mg/mL solution, thus requiring a large
administration volume of about 600 mL, at an alkaline pH of 9.5, as well as including mannitol (63). The administration-related issues are particularly onerous in the
setting of an intraoperative emergency maneuver. To increase loading of this poorly
water-soluble drug in a readily reconstituted format, either dantrolene or its sodium
salt was coated with egg phospholipid prepared by homogenization using a microfluidizer. Following lyophilization, rapid reconstitutability within one minute to
produce 10% to 15% suspensions was enabled for both the dantrolene nanosuspension (NS-D) and the sodium dantrolene nanosuspension (NS-NaD). Particle
sizes of 300 to 800 nm for NS-NaD and 500 to 800 nm for NS-D were measured.
Dissolution of both nanosuspensions was rapid. A formal comparison for
NS-NaD yielded pharmacokinetics equivalent to that for the solution formulation
of the drug. A strain-gauge transducer of forelimb adduction measured the dose–
response curves. These were also comparable, and are summarized as the effective
42
Rabinow
dose necessary to produce 50% and 95% of the plateau response, as well as the magnitude of the plateau response. Similar values were also found in normal swine.
Although the NS-D formulation gave a slightly higher ED50 twitch depression than
Dantrium in malignant hyperthermia-susceptible swine (1.0 ± 0.2 vs. 0.6 ± 0.1 mg/kg),
the more important ED95 was statistically equivalent (3.5 ± 0.4 vs. 2.7 ± 0.5). ED95 is
more significant because the drug is dosed to near-plateau effect. The nanosuspensions successfully treated or prevented malignant hyperthermia in swine models.
With NS-NaD, 2.5 mg/kg IV bolus in swine was observed to cause pulmonary
artery (PAP) hypertension as well as systemic hypotension. This was similar to the
response seen with the injection of undissolved dantrolene powder. The systemic
hypotension was eliminated with addition of a 6 µm filter, and the PAP increase was
significantly reduced but not eliminated. NS-D, on the other hand, produced only
minimal PAP increase when injected at the same rate with a 2 µm filter. It was suspected that postinfusion aggregation of particles was responsible for these effects, 19%
of the NS-NaD being observed in aggregates of greater than 3 µm diameter, following
200× dilution. In contrast, no aggregation was seen for NS-D following dilution.
Earlier, it had been demonstrated that agglomeration of injected particles
could sometimes be seen after fast (1 mL/5 sec) but not after slow (1 mL/min) injection of nanoparticles, essentially leading to plug flow (64,65). Additional explanation was provided by Ward and Yalkowski (66) who found that plug flow occurred
where the injection rate matched the blood flow rate. If one injected either slower or
faster than this rate, faster dissipation of the bolus would occur. The safer strategy
is slower infusion (66). In dogs, rapid IV bolus up to 10 mg/kg of unfiltered NS-D
could be administered with no PAP change, suggesting that swine may be a more
sensitive model than the dog.
Antifungal
Preclinical Studies with Drug-Susceptible Fungal Strain
When formulated as a nanosuspension, the antifungal agent itraconazole could be IV
dosed to the rat with no mortality at nearly 10 times the LD50 of the commercially
available hydroxypropyl-β-cyclodextrin-solubilized drug, Sporanox® ( Janssen
Pharmaceutical Products, L.P.). This may occur because of reduced Cmax, and therefore reduced Cmax-induced toxicity, inasmuch as the slowly dissolving itraconazole
particles are initially sequestered by the MPS. Subsequently, the drug is released over
prolonged therapeutic, yet apparently, subtoxic levels. This safety profile permitted a
much greater dose to be administered to rats and dogs. In Candida albicans-challenged
immuno-suppressed rat models, the greater dose resulted in higher drug levels at the
site of infection in the kidney which significantly reduced colony counts in that organ.
In fact, colony counts of zero were found with sufficiently high doses of the nanosuspension. Treatment with Sporanox® resulted in a decrease in the colony-forming units,
but they did not decline to zero. These results are consistent with the work of Andes
(67), who found that for drugs of the azole class, the AUC/MIC ratio is the critical
pharmacokinetic/pharmacodynamic parameter associated with treatment efficacy.
Here, AUC is the area under the drug plasma curve and MIC is the minimum inhibitory concentration. Efficacy of azoles is independent of peak concentration.
Preclinical Studies with Drug-Resistant Fungal Strain
Animal results suggest that higher dosing, achievable with nanosuspensions, may
cause a reconsideration of the in vitro–in vivo correlation of antifungal susceptibility
43
Nanosuspensions for Parenteral Delivery
testing. Current guidelines for interpretive breakpoints of MIC for mucosal Candida
infections are: susceptible ≤0.125 µg/mL; susceptible, dependent upon dose (S-DD):
0.25 to 0.5 µg/mL; and resistant ≥1.0 µg/mL (68). That is to say, the fungal infection
should be clinically treatable if the results of the susceptibility testing indicate an
MIC ≤ 0.125 µg/mL; and itraconazole should not be clinically effective if the
MIC ≥ 1.0 µg/mL. Between 0.25 and 0.5 µg/mL, the organism may be considered susceptible, dependent upon dose (S-DD). It is clear, however, that these guidelines were
formulated with assumptions made about the dose of itraconazole that can be administered, dependent upon current approved drug labeling. If greater dosing becomes
clinically achievable through nanosuspensions, then the interpretive breakpoints of
MIC, corresponding to what is considered treatable, may be revised upward.
Preliminary animal results suggest that efficacy toward fungal strains, conventionally considered resistant to itraconazole, can be demonstrated for itraconazole
nanosuspensions. Thus, a prednisolone immuno-compromised rat model was challenged with a Candida strain considered resistant to itraconazole by the above guidelines (MIC = 16 µg/mL). Survival of the majority of the nanosuspension-treated
animals was observed by the end of the 10-day experiment, whereas all of the
Sporanox®-treated animals had died (69).
Clinical Study
The pharmacokinetics of itraconazole nanosuspension has been studied clinically in
allergenic hematopoietic stem cell transplant recipients over a 14-day intravenous
course of treatment (70). On days five and six prior to the transplant, 200 mg of
nanosuspension was given IV every 12 hours followed by 200 mg every 24 hours for
the next 12 days. Steady state was not reached, and the therapeutic level of 500 µg/L
was maintained in five of the six cases for at least nine days after treatment cessation. The pharmacokinetic parameters found are compared in Table 2 with those of
the commercial Sporanox® solution, as listed in the Physicians’ Desk Reference (71)
and other literature.
It is seen that in comparison with the solution formulation Sporanox®, the
nanocrystal suspension occupied a larger volume of distribution and was cleared
more slowly to give a longer half-life and larger area under the plasma concentration curve for the first 24 hours, AUC24. This pharmacokinetic behavior is consistent
with sequestration in the MPS depot, causing reduced clearance and increased
volume of distribution. Subsequent release results in greatly prolonged delivery for
the nanosuspension, as shown by half-life and more efficient utilization of the drug,
as manifested by higher AUC.
The clinical results confirmed the animal data and suggest that the daily
dosing established for Sporanox® may be reduced to a frequency commensurate
with the prolonged half-life of the nanosuspension. The maintenance of therapeutic
TABLE 2 Comparison of Clinical Pharmacokinetics Between Itraconazole
Nanosuspension and Sporanox®
Itraconazole nanosuspension
Vss (L)
Cl (L/hr)
t½ (hr)
1677 ± 827
3.35 ± 1.8
346 ± 225 terminal
AUC24 (μg hr/L)
51558 ± 10635
Sporanox
796 ± 185
22.9 ± 5.7
35.4 ± 29.4 mean and
30 terminal (72)
30605 ± 8961
44
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levels for nine days following termination of treatment is especially interesting, and
suggests utility when patients migrate from intravenous to oral dosing, following
hospital discharge, for example. Having a reliable internal depot of drug would be
useful during the transition process to ensure continuity of delivered dose. Further
clinical studies would be needed to confirm this point.
Anticancer: Paclitaxel
Characterization and Manufacture
Abraxane™ for Injectable Suspension is an approved, commercialized nanosuspension formulation. As shown by transmission electron microscopy, it consists of a
core of paclitaxel, surrounded by an albumin shell with an overall mean size of
130 nm, spanning a range of about 100 to 200 nm (73). It is manufactured by adding
a methylene chloride solution of paclitaxel to an aqueous solution of human serum
albumin, using low-speed homogenization. The albumin migrates to the aqueous–
solvent interface of the emulsion that is created. Application of high-pressure
homogenization reduces the particle size and cross-links the albumin coating via
the disulfide bonds, thus stabilizing the particle. The methylene chloride is then
volatilized, leaving an aqueous suspension of nanoparticles. The particles consist of
an amorphous paclitaxel core coated by a 25-nm thick shell of albumin with bound
paclitaxel. The particle size is sufficiently small so as to be sterile-filtered (18).
Importantly, Abraxane is formulated without Cremophor EL, which has been associated with a host of problems including sensitivity reactions, the need to infuse
over prolonged periods of time, leaching of plasticizer from infusion sets, and so on.
Pharmacokinetics
Clinical pharmacokinetic studies for Abraxane have defined parameters relative to
Taxol® summarized in Table 3 (74,75). The shorter 30-minute duration of infusion of
Abraxane understandably results in higher plasma Cmax than that for Taxol® administered with a three-hour infusion. The near-equivalence of half-lives, although not
observed earlier for mice, reflects a similar terminal elimination rate of drug from
the tissue compartment. Ibrahim et al. have attributed the reduced AUC of the nanosuspension to possibly several factors. The first is faster partitioning of the nanosuspension paclitaxel out of the vascular compartment, in comparison with Taxol®. It is
known that Cremophor micelles reduce the free paclitaxel plasma fraction available
for cellular partitioning, thus reducing tissue distribution from the central blood
compartment. Not being formulated with Cremophor, Abraxane might be expected
to exit the circulation faster. The second reason involves increased endothelial transcytosis via albumin-mediated gp60 receptor uptake of the particles (76).
Clinical Trial
Abraxane and Taxol® were studied head to head in a Phase III trial involving
460 breast cancer patients. Taxol® was administered using the standard protocol of
TABLE 3 Comparison of Clinical Pharmacokinetics Between Abraxane Paclitaxel
Nanosuspension and Taxol®
Drug/dose
Abraxane (135 mg/m2/
30 min)
Taxol (135 mg/m2/3 hr)
Cmax (ng/mL)
AUC(0–∞) (ng hr/mL)
t½ (hr)
Cl (L/hr/m2)
6100
6427
15
21
2170
7952
13
18
Nanosuspensions for Parenteral Delivery
45
175 mg/m2 by 3-hour infusion, including premedication with steroid and antihistamines to inhibit Cremophor-related hypersensitivity. In contrast, Abraxane was
given at 260 mg/m2 over a shorter duration 30 minutes, without premedication or
G-CSF support. Despite the more aggressive protocol for Abraxane, the toxicity was
no worse: there were no hypersensitivity reactions; neutropenia decreased; whereas
neuropathy increased somewhat. This is significant given the correlation established
between the duration of plasma paclitaxel concentrations exceeding 0.1 µmol/L
with decline of absolute white blood cell count (74,75,77). In this trial, Abraxane also
produced a higher tumor response rate versus paclitaxel (33% vs. 19%) and a longer
time to tumor progression (21.9 wk vs. 16.1 wk) (78).
The improved efficacy of the nanosuspension may be surprising in view of the
purported benefit of Cremophor EL in inhibiting the P-glycoprotein efflux pump,
thereby enhancing drug level in tumor cells (79). The alleged benefit of this is
probably overrated, inasmuch as Cremophor is retained in the central blood compartment, and therefore does not enter the tumor tissue (80). There are ancillary
benefits to not formulating with Cremophor. As corticosteroids do not have to be
taken as premedication, there is the possibility for combining paclitaxel with IL-2 or
interferon for treatment of metastatic melanoma, renal cell carcinoma, and so on.
The current Cremophor-containing formulation cannot be used because steroids,
used for premedication, lyse lymphokine activated killer (LAK) cells, thus mitigating the benefits of the cytokines (18).
Intrathecal Delivery
Regional delivery of water-insoluble drugs offers the possibility of increasing local
therapeutic concentrations while decreasing systemic side effects. In early work,
epidural injection of a 10% butamben suspension intended for chronic cancer pain
was well tolerated in dogs and humans (44,81). As a treatment modality for intractable brain tumors, the technique of direct injection into the ventricles of the brain is
known as convective enhanced delivery. In addition to simply placing the drug
within the central nervous system, this pressure gradient microinfusion of drug
overcomes diffusion barriers associated with high intratumoral interstitial pressures
and disordered tumor vasculature. As a result, drug is more rapidly distributed
throughout the target volume (46). As an example of this application, intrathecal
delivery of nanosuspension busulfan to a mouse model of neoplastic meningitis led
to a significant increase in survival (47,49). The pharmacokinetics was determined
in patients afflicted with neoplastic meningitis, who received the drug both by an
Ommaya reservoir for intraventricular delivery and via lumbar puncture. The drug
was well tolerated and resulted in delayed progression of disease (50).
CONCLUSIONS
Growth in the applications of nanosuspension technology has occurred in response
to the voluminous number of water-insoluble drug candidates that have emerged
from discovery programs. The inherent high loading of this dosage form distinguishes it from liposomes, emulsions, cyclodextrins, and polymeric nanoparticles,
permitting dosing in animal toxicity studies at the required multiples of anticipated
human exposure. Parenteral applications for subcutaneous, intramuscular, intradermal, intravenous, epidural, and intrathecal delivery have been studied in animals,
with enhanced efficacy. At the same time, the safety profile has been observed to be
46
Rabinow
improved in many cases when compared to conventional solution forms of the drugs.
This occurs due to deletion of noxious excipients, change in the pharmacokinetic
profile, or regional delivery, thus minimizing systemic toxicity. These therapeutic
and safety benefits have been demonstrated for several drugs for different disease
indications in clinical trials. On the basis of the successful clinical applications, it is
anticipated that growth of this formulation tool will accelerate.
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Antimicrobial Agents and Chemotherapy, Chicago, IL, December 16–19, 2001.
71. SPORANOX® (Itraconazole) Injection. Approved labeling. In: PDR 58 ed Physicians’
Desk Reference®. Montvale, NJ: Thomson PDR, 2004:1772.
72. Willems L, van der Geest R, de Beule K. Itraconazole oral solution and intravenous formulations: a review of pharmacokinetics and pharmacodynamics. J Clin Pharm Therap
2001; 26:159.
73. Blum JL, Beveridge R, Robert N, et al. ABI-007 Nanoparticle paclitaxel: demonstration of
anti-tumor activity in taxane-refractory metastatic breast cancer. Abstract 64. Proc Am Soc
Clin Oncol 2003.
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74. Ibrahim NK, Desai N, Legha S, et al. Phase I and pharmacokinetic study of ABI-007, a
cremophor-free, protein-stabilized, nanoparticle formulation of paclitaxel. Clin Cancer
Res 2002; 8:1038.
75. Taxol® (paclitaxel) Injection. Direction insert. Princeton, NJ: Bristol-Myers Squibb
Oncology 2003.
76. Desai N, Trieu V, Yao R, et al. Increased endothelial transcytosis of nanoparticle albuminbound paclitaxel (ABI-007) by endothelial gp60 receptors: a pathway inhibited by Taxol.
Poster AO 066. In: 27th San Antonio Breast Cancer Symposium, San Antonio, TX,
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77. Huizing MT, Keung ACF, Rosing H, et al. Pharmacokinetics of paclitaxel and metabolites
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78. Lee D. Current trials of a nanoparticle albumin-bound taxane formulation in metastatic
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79. Woodcock DM, Jefferson S, Linsenmeyer ME, et al. Reversal of the multidrug resistance
phenotype with cremophor EL, a common vehicle for water-insoluble vitamins and
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80. Sparreboom A, Verweij J, Van der Burg MEL, et al. Disposition of Cremophor EL in
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Reg Anesth 1990; 15:142.
3
Nanoparticles Prepared Using Natural
and Synthetic Polymers
Sudhir S. Chakravarthi and Dennis H. Robinson
Department of Pharmaceutical Sciences, University of Nebraska Medical Center,
Omaha, Nebraska, U.S.A.
Sinjan De
Research and Development, Perrigo Company, Allegan, Michigan, U.S.A.
INTRODUCTION
Ideally, biologically active agents should be encapsulated within nanoparticles
using polymers with well-defined physical and chemical properties. These polymers
protect the active ingredient and, after local or systemic administration, potentially
target and then release the drug in a controlled, predictable manner. The use of
polymeric nanoparticle for drug delivery is a strategy that aims to optimize therapeutic effects while minimizing adverse effects. The purpose of this chapter is to
highlight the major factors related to the use of nanoparticles fabricated from
synthetic or natural polymeric materials that have been used in drug delivery and
imaging. A comprehensive review of this area of research is beyond the scope of this
chapter and hence the readers are referred to other reviews in the areas for additional information (1–3).
POLYMERIC CARRIERS USED TO PREPARE NANOPARTICLES
Polymers used in controlled drug delivery, including nanoparticles, may be classified as either (i) natural and synthetic, or (ii) biodegradable and nonbiodegradable.
Examples of naturally occurring biodegradable and biocompatible polymers used
to prepare nanoparticles include: cellulose, gelatin, pullulan, chitosan, alginate, and
gliadin. The characteristics and performance, particularly in vivo, of nanoparticles
prepared using natural polymers may be less predictable as these polymers may
vary widely in chemical composition and hence, physical properties. In addition,
natural polymers are often mildly immunogenic. Conversely, it is possible to synthesize polymers with precise chemical composition, resulting in highly predictable
physical properties such as solubility, permeability, and rates of biodegradation.
As a result, synthetic polymers are also more easily designed for specific applications, such as controlled rates of dissolution, permeability, degradation, and erosion,
as well as for targeting. Examples of synthetic biodegradable polymers used to prepare nanoparticles include: polylactide (PLA), poly-(lactide-co-glycolide) (PLGA),
polyanhydrides, poly-ε-caprolactone, and polyphosphazene. Biodegradability and
biocompatibility are important properties of polymeric materials that are to be
injected or implanted into the body. Nonbiodegradable polymeric nanoparticles
may be used for controlled drug delivery and also in the complimentary field of
diagnostic imaging. Examples of nonbiodegradable, synthetic polymers used in
drug delivery include polymethyl methacrylate while polystyrene particles have
been used as diagnostic agents.
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Chakravarthi et al.
Natural Biodegradable Polymers Used to Prepare Nanoparticles
Alginates
Alginates are linear, unbranched polysaccharides composed of random chains of
guluronic and mannuronic acids (4). In aqueous media, the sodium ions from salts
of these anionic, heteropolymers exchange with divalent cations, such as calcium, to
form water-insoluble gels (5). Because of the favorable conditions during manufacture, alginates are ideal carriers for oligonucleotides (6), peptides (7), proteins (7),
water-soluble drugs, or drugs that degrade in organic solvents. Alginates are nonimmunogenic and available in a wide range of molecular weights as characterized
by their inherent viscosity. Alginate nanoparticles are prepared by extruding an
aqueous sodium alginate solution through a narrow-bore needle into an aqueous
solution of a cationic agent, such as calcium ions, chitosan, or poly-l-lysine. These
cations cross-link the guluronic and mannuronic acids to form an egg-box structure
that forms the core of the gel matrix. In vivo, therapeutic agents are released when
the matrix redissolves due to the reversible exchange of divalent cations with monovalent ions, especially sodium present in physiological fluid. A disadvantage of the
use of alginates is that this reversible ion exchange may result in the rapid release of
the therapeutic agent. However, an example of the use of alginate nanoparticles to
sustain antibacterial drug levels above the minimum inhibitory concentration in the
liver, lungs, and spleen after pulmonary administration was demonstrated using
isoniazid, rifampicin, and pyrazinamide (8). One method to prolong release from
alginate particles is to coat them with a cationic polymer, for example, poly-l-lysine
or chitosan. In this application, the mass ratio of alginate to cationic polymer is
critical in terms of release characteristics and particle size (9).
Chitosan
Chitosan is a natural polymer obtained by deacetylation of chitin, a component of crab
shells. It is a cationic polysaccharide composed of linear β(1,4)-linked d-glucosamine.
The various methods used to prepare chitosan-based nanoparticles and their
applications have been extensively reviewed (10). Chitosan can entrap drugs by
numerous mechanisms including chemical cross-linking, ionic cross-linking, and
ionic complexation (11).
Gelatin
Gelatin is a natural, biodegradable protein obtained by acid- or base-catalyzed
hydrolysis of collagen. It is a heterogenous mixture of single- or multi-stranded
polypeptides composed predominantly of glycine, proline, and hydroxyproline
residues and is degraded in vivo to amino acids. Gelatin nanoparticles are prepared
by a two-step, desolvation process (12). Briefly, this coacervation procedure involves
the addition of either another more water-soluble polymer or, a water-miscible
nonsolvent for gelatin, to an aqueous gelatin solution above its gel temperature of
about 40°C. The concentrated gelatin liquid particles are isolated and hardened by
chemical cross-linking with glutaraldehyde. Alternately, these particles can be prepared using a simple o/w emulsion or w/o/w microemulsion method. Gelatin
nanoparticles have been used to deliver paclitaxel, methotrexate, doxorubicin,
DNA, double-stranded oligonucleotides, and genes. PEGylation of the particles
significantly enhances their circulation time in the blood stream (13) and increases
their uptake into cells by endocytosis. Antibody-modified gelatin nanoparticles
have been used for targeted uptake by lymphocytes (14).
Nanoparticles Prepared Using Natural and Synthetic Polymers
53
Pullulan
Similar to dextran and cellulose, the glucans in pullulan are water-soluble, linear
polysaccharides that consist of three α-1,4-linked glucose molecules polymerized
by α-1,6 linkages on the terminal glucose (15). Pullulan is a fermentation product of
the yeast Aureobasidium pullulans. When made hydrophobic by acetylation, these
polymers will self-associate to form nanoparticles with a hydrophobic core that will
encapsulate hydrophobic drugs. Pullulan nanoparticles have been prepared by
dialysis of an organic solution against water. In one method, a reverse micellar
solution of the anionic surfactant, Aerosol OT, in n-hexane was prepared and an
aqueous solution of the drug and pullulan added (16). The nanoparticles are stabilized by cross-linking with glutaraldehyde. These delivery systems have been used
in delivering cytotoxic drugs, genes, and as pH-sensitive delivery systems.
Gliadin
Gliadin is a glycoprotein that, as a component of gluten, is extracted from glutenrich food such as wheat flour. They are slightly hydrophobic and polar. Bioactive
molecules of variable polarity can be encapsulated into gliadin nanoparticles.
Gliadin nanoparticles can be prepared by a desolvation method that exploits the
insolubility of this polymer in water (17). Briefly, gliadin nanoparticles are precipitated when an ethanolic solution of gliadin is poured into an aqueous solution.
Gliadin nanoparticles have been used to deliver trans-retinoic acid, α-tocopherol,
and vitamin E. Lectins have been conjugated to the surface of gliadin nanoparticles
to target the colon and treat Helicobacter pylori infections (17).
Synthetic Biodegradable Polymers Used to Prepare Nanoparticles
Polylactide and Polylactide-co-Glycolide
The hydrophobic PLA may be used alone or copolymerized with poly-glycolic acid
to form a range of PLGA of widely varying polymeric ratios and hence physicochemical properties. These FDA-approved polymers have been widely used in drug
delivery including nanoparticles. PLA and PLGA polymers degrade by random
bulk hydrolysis that is catalyzed in acidic media. The methods used to prepare PLA
and PLGA nanoparticles as well as their range of applications, physical properties,
biological fate, and targetability have been comprehensively reviewed (2).
Polyanhydrides
Polyanhydrides are biodegradable polymers with a hydrophobic backbone and a
hydrolytically labile anhydride linkage. They are synthesized by ring-opening
polymerization and degrade by surface hydrolysis (3). The application of polyanhydrides has been limited to film and microsphere formulation for sustained release
of a drug or protein at the site.
Poly-є-Caprolactones
Methods used to prepare nanoparticles using poly-ε-caprolatones have been previously reviewed (18) and include emulsion polymerization, solvent displacement,
dialysis, and interfacial polymer deposition. These semicrystalline polymers are
chemically stable, possess a low glass transition temperature, and degrade slowly.
Hence, they have the potential for long-term drug delivery. Poly-ε-caprolactone
nanoparticles have been used as vehicles to deliver a wide range of drugs including
tamoxifen, retinoic acid, and griseofulvin.
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Polyalkyl-Cyanoacrylates
Polyalkyl-cyanoacrylate (PACA) nanoparticles are prepared by the conventional
emulsion-evaporation technique. In addition to sustaining drug release, PACA
nanoparticles have the ability to overcome multidrug resistance at both the cellular
and subcellular levels (19). The potential for targeted delivery of PACA nanoparticles
to cells has been demonstrated by conjugation of polysaccharides to the surface (20).
Nonbiodegradable Polymers Used to Prepare Nanoparticles
Polymethacrylate (PMA) and polymethyl methacrylate (PMMA) have been widely
used in a variety of pharmaceutical and medical applications. Specifically, PMMA
Eudragit® nanoparticles can be prepared by nanoprecipitation method (21) that
involves adding hydroalcoholic solution of the polymer to an organic solvent.
Incorporation of poly-acrylic acid into nanoparticles increased the transfection efficiency of DNA. The side chain of PMMA can be modified to make these polymers
possess pH-dependent solubility and has been used to prepare pH-sensitive
nanoparticles to increase the oral bioavailability.
Thermosensitive Nanoparticles
Polyvinyl caprolactone nanoparticles exist in swollen state when dispersed in water
at room temperature. However, they shrink upon increasing temperature above the
volume phase transition temperature, expelling water (22). Poly(N-isopropyl acrylamide) (NIPAAM) has been extensively used to prepare thermosensitive hydrogelbased nanoparticles. This polymer has a low critical solution temperature of 32°C to
34°C, above which it shrinks to release the drug.
Solid–Lipid Nanoparticles
Although not polymers, purified triglycerides and waxes that are solid at ambient
and body temperatures can also be used as nanoparticulate carriers. These may be
prepared by various methods including high-speed homogenization, microemulsions, emulsion-solvent evaporation, and ultrasonication (1). It is possible to prepare
lipid nanoparticles with high drug loading, particularly with lipophilic drugs that
sustain release due to their hydrophobicity and low drug diffusivity in the matrix.
CHARACTERIZATION OF NANOPARTICLES
Physical Properties of Polymers
The following physicochemical properties of polymers greatly influence the properties, method of preparation, and performance of nanoparticles.
Molecular Weight
The molecular weight of a polymer influences physical properties such as glass
transition temperature, viscosity in solutions, solubility, crystallinity, degradation
rate, and mechanical strength. In general, polymers with a lower molecular weight
exhibit a lower viscosity and tensile strength, and degrade more rapidly. Hence,
selection of a polymer with an appropriate molecular weight is important for the
intended application. For example, if a polymer degrades by acid-catalyzed, bulk
hydrolysis, a low-molecular-weight polymer will degrade faster due to autocatalysis by the greater proportion of oligomers formed. Inefficient polymeric synthesis
Nanoparticles Prepared Using Natural and Synthetic Polymers
55
may form polymers with high polydispersity that degrade more rapidly than homopolymers of similar molecular weight.
Degree of Crystallinity
The mechanical properties of polymers can be altered by the degree of crystallinity.
For example, because of the uniform arrangement of its chains within the lattice
structure, a crystalline polymer will degrade slowly than an amorphous form.
However, pure crystalline polymers are brittle and usually less suited to drugdelivery applications. Further, amorphous polymers possess poor mechanical
toughness. Therefore, the polymers used in drug delivery are usually a mixture of
crystalline and amorphous forms.
Hydrophobicity
Factors that influence the hydrophobicity of the polymer include molecular weight,
aqueous solubility of the monomers, and the degree of branching. Although increase
in molecular weight increases the hydrophobicity, an increase in the degree of
branching results in a more water-soluble polymer. Hence, nanoparticles prepared
using a hydrophobic polymer exhibit decreased water penetration and wettability,
resulting in relatively slower drug release and polymer degradation times than
similar hydrophilic forms. However, the incorporation of a hydrophilic polymer or
additives into nanoparticles prepared using hydrophobic polymers will form pores
in aqueous media to increase the rate of polymer degradation and erosion as well as
drug release. It is important to realize that the method used to prepare nanoparticles
will be influenced by the relative hydrophilicity and hydrophobicity of both the
polymer and the drug.
Copolymer Ratio
The choice of a polymer and the method of polymerization directly affect the type
of copolymer as well as its molecular weight, crystallinity, and hydrophobicity. In
general, copolymers used to prepare nanoparticles typically contain both hydrophobic and hydrophilic segments which facilitate greater flexibility in preparation
and more predictable physical properties such as release characteristics.
Biodegradability and Biocompatibility
A biodegradable polymer must degrade into physiologically inert products that
are eliminated by the body. For example, PLGA polymers hydrolyze to form lactic
and glycolic acids that are further degraded into normal constituents of the
body. Biocompatible polymers are defined as those that do not elicit immune reaction or inflammation, are stable for the duration of action, and are completely
metabolized in the body. Most biodegradable polymers used in drug delivery are
specifically intended for parenteral administration, which is not a prerequisite for
oral delivery.
Solubility
The solubility spectrum of a polymer influences the method of preparation and in
vivo performance of the nanoparticles. In general, if both the drug and the polymer
are soluble in organic solvents, a simple o/w emulsion technique or phase separation
can be used to prepare the nanoparticles. For protein and peptide drugs, a milder
aqueous environment is preferable and therefore, the choice of the polymer is critical and a multiple emulsion technique may be used.
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Drug–Polymer Interactions
Chemical and physical interactions that occur depend on the chemical nature of
both the polymer and the drug as is well documented in the literature. Drug–polymer interactions are chiefly determined by charge, solubility, and hydrophobicity
and may greatly alter the properties of the polymer, for example, glass transition
temperature and degree of crystallinity, as well as the properties of the nanoparticles, for example, release characteristics. The interactions that occur between a drug
and a polymer can be identified, and in some cases quantified, using differential
scanning calorimetry (DSC), thermogravimetric analysis (TGA), solid-state 13C NMR,
Fourier transform infrared spectroscopy, Flory Huggins interaction parameters, and
comparison of total or partial solubility parameters (23). For example, isothermal
calorimetry has been used to determine the binding affinity between alginates and
chitosan as well as poly l-lysine (9). Differences in the solubility of the drug in the
polymer results in the microphase separation of the drug within the matrix (24).
Quantifying the binding affinity using isothermal calorimetry can identify an electrostatic interaction between the polymer and the drugs.
Preparation of Polymeric Nanoparticles
The main methods used to prepare polymeric nanoparticles include emulsionsolvent evaporation, phase separation, and use of supercritical fluid technology.
These methods will not be described here as they have been extensively reviewed in
the literature (1,2,4,6–14,16,17–22,24,25).
Properties of Nanoparticles that Affect Biological Performance
Route of Administration
Nanoparticles have been administered by the following routes: oral, intravenous,
subcutaneous, intrathecal, intraocular, and pulmonary. When selecting a route to
administer nanoparticles, it is important to consider the stability of the drug in biological fluids, as well as anatomical and physiological characteristics of the route of
administration and at the target site. Oral delivery of nanoparticles may significantly
increase the bioavailability of poorly soluble drugs. However, after intravenous
administration, the bioavailability of drugs may decrease as macrophages will internalize hydrophobic unmodified nanoparticles. This problem may be circumvented
by coating the surface of nanoparticles with hydrophilic polyethylene glycol, which
confers “stealth” properties to these particles. It is important to note that, after pulmonary administration, nanoparticles may drain into the lymphatic system and, as
with all nebulizers and aerosols, the site of deposition of the particles will depend
on the type of the device used. Although not widely employed, intrathecal delivery
of nanoparticles may sustain the release of the encapsulated drug.
Particle Size and Particle Size Distribution
The particle size and particle size distribution are critical factors in the performance
of nanoparticles, as batches with wide particle size distribution show significant
variations in drug loading, drug release, bioavailability, and efficacy. Particle size
and particle size distribution can be determined using light scattering techniques
and by scanning or transmission electron microscopy (Fig. 1). Formulation of
nanoparticles with a narrow size distribution will be a challenge if emulsion cannot
be produced with a narrow droplet size distribution. As nanoparticles are internalized into cells by endocytosis, an increase in particle size will decrease uptake and
Nanoparticles Prepared Using Natural and Synthetic Polymers
57
FIGURE 1 Scanning electron
microscope image of poly-(lactideco-glycolide) nanoparticles.
potentially, bioavailability of the drug. The extent of endocytosis is dependent on
the type of the target cell.
Zeta Potential
The charge on the surface of the nanospheres will influence their distribution in the
body and extent of uptake into the cells. Because cell membranes are negatively
charged, there is greater electrostatic affinity for positively charged nanoparticles.
Therefore, the surface of cationic or neutral nanoparticles may be modified to confer
a positive charge to enhance efficacy. For example, positively charged tripalmitin
nanoparticles containing etoposide prolonged residence time in the blood, produced higher blood concentrations, decreased clearance by the liver, and increased
distribution into the brain and bone (26).
Drug Loading and Loading Efficiency
Although drug loading expresses the percent weight of active ingredient encapsulated to the weight of nanoparticles, drug loading efficiency is the ratio of the
experimentally determined percentage of drug content compared with actual, or
theoretical mass of drug used for the preparation of the nanoparticles. The loading
efficiency depends on the polymer–drug combination and the method used.
Hydrophobic polymers encapsulate larger amounts of hydrophobic drugs, whereas
hydrophilic polymers entrap greater amounts of more hydrophilic drugs. Several
formulation parameters, such as emulsifier type, weight ratio of polymer to drug,
and organic to aqueous phase ratio, will influence the extent of drug loading.
Dissolution Profile
The in vitro release of drugs from nanoparticles may approximate the drug release
profile inside the body although the rate is usually faster in vivo due to the presence
of enzymes and surfactants in biological fluids. An in vitro dissolution medium
mimics the pH and salt concentrations in the body. Particularly for hydrophobic
drugs, it is critical during dissolution testing that sink conditions are maintained
and pH and salt concentrations of biological fluid are approximated. If solubility of
the drug in the media is a limitation, it may be necessary to add surfactants to the
dissolution media. The release data must be evaluated using well-known equations
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to determine parameters such as the mechanism(s) of release, order of release, dissolution rate constant, extent of release from the nanoparticles, and duration.
Surface Modification
The surface of nanoparticles may be modified to conjugate targeting ligands or alter
biodistribution. Coating of nanoparticles with hydrophilic polymers such as
polyethylene glycol, heparin, or dextran, protects them from being engulfed by the
macrophages or kupffer cells, thereby increasing their circulation time and enhancing
drug bioavailability. Surface modification of nanoparticles also involves alteration
of the surface charge as discussed previously.
Biocompatibility
Tissue response to polymeric particles occurs in three phases. In phase I, mild inflammation is observed with monocytes, lymphocytes, and leucocytes surrounding the
particle. In Phase II, monocytes differentiate into macrophages, which further
differentiate into giant body cells and engulf the particles. In some cases, fibrous
tissue may be formed. Phase III involves further degradation of the particles. The
type of immune response elicited by the nanoparticles will depend on the route of
administration, the type of therapeutic agent encapsulated, and the choice of the
polymer. Because of dynamic flow, minimal immune responses are generally
observed after intravenous administration.
IN VITRO CELL CULTURE
Internalization of Nanoparticles into Cells
The different mechanisms by which the nanoparticles are taken up by the cells include
endocytosis, pinocytosis, fluid-phase diffusion, carrier-mediated, and facilitated
transport. Although endocytosis and carrier-mediated transport are ATP-dependent,
facilitated transport and diffusion are energy-independent. Nanoparticles undergo
endolysosomal escape which increases accumulation of particles inside the cell (27).
Receptor-mediated transport involves the internalization of nanoparticles through
specific cell surface receptors on the cell membrane, and is exploited in the design
of nanoparticles to specifically target these receptors (Fig. 2). Nanoparticles are
Nanosphere
bound to
receptor
Ligand
conjugated to
nanosphere
Recycling
of receptor
Receptor
mediated
endocytosis
of nanospheres
NUCLEUS
Internalization
of nanosphere
FIGURE 2 (See color insert.)
Schematic diagram of specific
receptor targeting of nanospheres using ligands.
Nanoparticles Prepared Using Natural and Synthetic Polymers
59
primarily internalized by endocytosis. Once the particles are internalized, they transcytose across cells deeper into the tissue. Recently, it was demonstrated that the
drug content in the cells increased with particle size, suggesting that in addition to
particle internalization, diffusion of free drug released outside the cells may also
play a role in enhancing the total drug content of the cells (28).
Protection from Efflux Pumps
Trans-membrane pumps, such as p-glycoprotein and multidrug-resistance-associated
protein, are present on the apical side of the cells and actively efflux foreign substances, including drugs, out of the cell, significantly reducing the drug uptake.
Nanoparticles can protect the encapsulated drug from these efflux pumps.
Targeted Delivery
Some limitations of drugs delivered using nanoparticulate vehicles include nonspecific uptake into cells, accumulation of particles in nonspecific regions, and
inability to differentiate between diseased and normal tissues. Targeted delivery of
nanoparticles increases the extent of drug uptake at the site of action while potentially reducing adverse effects. Targeted delivery is based on: (i) physiological or
externally induced phenomena (pH and thermo-sensitive nanoparticles, ultrasound-triggered nanoparticles, magnetic nanoparticles); (ii) cell- or tissue-specific
targeting (nanoparticles targeting transferrin, folate, epidermal growth factor); and
(iii) permeability-enhancing targets [nanoparticles targeting trans-activating transcriptional factor (TAT) peptide, integrin]. Targeted delivery of nanoparticles can
significantly improve the therapeutic efficacy and safety of drugs.
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4
Nanofiber-Based Drug Delivery
Matthew D. Burke
Department of Pharmaceutical Development, GlaxoSmithKline, Research Triangle
Park, North Carolina, U.S.A.
Dmitry Luzhansky
Department of Corporate Technology, Donaldson Company, Inc., Minneapolis,
Minnesota, U.S.A.
INTRODUCTION
Electrospinning is a process that was originally developed in the early 1930s, but
did not receive much attention until recent decades. Most likely the increased interest is due to the refocusing of more research groups on nanotechnology. Although
electrospinning has existed for a significant period of time and is relatively easy to
execute, the physics of electrospinning nanofibers is only understood to a limited
extent. A typical electrospinning process involves dissolving the drug of interest
and a polymer in an appropriate solvent. The solution is then placed in a syringe,
and a high voltage is applied. A small amount of the polymer solution is drawn out
of the syringe, forming a Taylor cone. Increasing the applied voltage further results
in the initiation of a charged fluid jet, which follows a chaotic trajectory of stretching
and bending until it reaches the grounded target. A stable jet is formed when the
charge is increased above a critical voltage, and there is a balance between the
surface tension of the fluid and the repulsive nature of the charge distribution on
the surface of the fluid. The presence of molecular entanglements in the polymer
solution prevents the jet from breaking into droplets (electrospraying), and when
combined with the electrical forces results in a whip-like motion of the jet, known as
bending instability. This process typically results in the drawing of a virtually
endless fiber with a nanometer-sized to micrometer-sized diameter. The final product is a three-dimensional nonwoven mat of entangled nanofibers with a high
surface-area-to-volume ratio (1). Scanning electron microscopy (SEM) is a typical
method to evaluate the nanofibers produced through electrospinning, as shown
in Figure 1.
Electrospun nanofibers have a very large surface-area-to-volume ratio, as
large as 1000 times that of a microfiber. This physical property has generated a
significant amount of interest in the biomedical and pharmaceutical industries
particularly for drug delivery of poorly soluble drug substances. In the pharmaceutical industry, recent trends in drug discovery have led to the development of a
significant number of highly potent compounds with extremely low solubility, thus
requiring advanced methods to create efficacious pharmaceutical products that
overcome the solubility issues of the drug. Currently, the pharmaceutical industry
uses methods such as milling to reduce the particle size of a drug substance, but this
is a high-energy method which can lead to stability issues and often cannot produce
truly nanosized drug particles. In conventional dry milling, the limit of fineness is
reached in the region of 100 µm. Wet grinding or wet bead milling produces further
reduction in the particle size, but often not below the micrometer range.
61
62
Burke and Luzhansky
FIGURE 1 Sample SEMs of various types of GRAS polymer nanofibers produced by electrospinning. (A) Cellulose acetate, (B) PVAc, (C) polyethylene oxide, and (D) Kollidon® SR (BASF AG,
Ludwigshafen, Germany). Abbreviations: GRAS, generally regarded as safe; PVAc, polyvinylacetate; SEM, scanning electron microscopy.
The utility of the electrospinning process for pharmaceutical products is the
single-step creation of nanosized drug particles with a low-energy process. The
selection of the polymer also controls the drug-release properties. Therefore, immediate-release nanofibers can be created by water-soluble polymers, enteric-release
nanofibers can be created by enteric polymers such as methacrylic acid copolymers,
and sustained-release nanofibers can be created by polylactic acid or polyvinyl acetate polymers. Although the diameter of electrospun fibers is often characterized
using SEM as proof that they are nanosized, it is important to note that the size of
the drug particles embedded in the nanosized fiber is significantly smaller than the
diameter of the fibers themselves.
Further utility may be found with electrospinning by embedding it at the end
of the chemical synthesis of the drug substance, which would streamline the transfer of material from chemical development to pharmaceutical development. The
last step in the chemical synthesis of a drug is usually a purification/precipitation
step to create a drug powder. Then the powder is transferred to pharmaceutical
development teams, which often mill the powder to a particular particle size and
further granulate/process the material into a tablet. This powder handling and processing, which often requires safety controls, can be avoided if the last step in the
63
Nanofiber-Based Drug Delivery
chemical synthesis is transformed into a step where the drug and a preferred polymer
are in an appropriate solution for electrospinning. Then the drug/polymer solution
can be directly electrospun into a final product, thus creating a seamless process from
chemical synthesis through creation of an appropriate final pharmaceutical product.
This would eliminate powder handling of the drug substance, and possibly reduce
variations in the final particle size of the drug substance during manufacturing.
DISSOLUTION ENHANCEMENT FOR IMMEDIATE-RELEASE
DOSAGE FORMS
Pharmaceutical tablets or capsules which immediately release their drug cargo when
orally ingested are the most common dosage form in the pharmaceutical industry.
Creating a product which performs this task is challenging with low-solubility
drugs. Inappropriate formulation of low-solubility drugs can result in slow or limited drug dissolution in the gastrointestinal (GI) fluids and subsequently minimal
drug being absorbed into systemic circulation. As mentioned above, currently a
common technique to overcome this low bioavailability issue is to reduce the particle
size of the drug substance to increase the dissolution of the drug into the GI fluids.
The particle size of the drug is directly related to the rate that it dissolves. The effect
of particle size on the drug dissolution process can be mathematically described by
the Nernst and Brunner diffusion layer model (2):
dQ
D
=
S(Cs–Cg)
dt
h
where Q is the amount of the drug dissolved, t the time, D the diffusion coefficient
of the drug in the solubilizing fluids of the GI tract, S the effective surface area of the
drug particles, h the thickness of a stationary layer of solvent around the drug particles, Cs the saturation solubility of the drug in the stationary layer h, and Cg the
concentration of the drug in the bulk fluids of the GI tract.
Based on the drug dissolution equation described in the previous paragraph,
it is clearly evident that the surface area of the drug particle is a key physicochemical property which can be used to increase the rate of drug dissolution. The particle
size of the drug substance is often used as a surrogate marker for surface area; therefore, a decrease in the particle size will increase the surface area and thus increase
drug dissolution. As mentioned earlier, the current particle size reduction technologies often reach a limit around 1 micron. In contrast, by first intent electrospinning
produces nanofibers less than 1 micron, thus electrospinning presents a mechanism
to allow further reduction of the particle size beyond the current technologies. The
electrospun nanofibers are preferably a homogeneous mixture of polymer and drug,
thus the particle size of the drug should be significantly less than the diameter of the
nanofiber. Although the exact particle size of the drug in the nanofiber remains
challenging to quantify, the dissolution rate can provide indirect evidence that
significant particle size reduction has occurred. Reduction of the particle size of the
drug substance through electrospinning with a rapidly dissolving polymer causes
the drug dissolution rate to be very rapid, as shown in Figure 2.
Another advantage of electrospinning as a dissolution-enhancement tool is
the ability to control the morphology of the drug substance (3,4). By selecting a
polymer such as hydroxypropylmethylcellulose acetate succinate (HPMC-AS),
which has an amorphous character, one can create an amorphous drug substance
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Burke and Luzhansky
FIGURE 2 (A) Dissolution of a model low-solubility drug from polyethylene oxide electrospun fibers
in a USP 2 dissolution apparatus. (B) Scanning electron microscopy image of the nanofibers.
through electrospinning. Amorphous drug substances are at a higher energy state,
therefore, in general have higher solubility and higher dissolution rate compared
with crystalline materials. Thus, the creation of an amorphous drug substance
represents another technique for dissolution enhancement. Amorphous drugs are
generally less stable physically and chemically than corresponding crystalline
materials, so appropriate stability assessment needs to be performed (5). Although
stability issues can limit the use of amorphous drug substances, there are several marketed oral products such as Vancocin® (Viropharma, USA), Plendil ER® (AstraZeneca
Ltd., UK), Cesamet® (Eli Lilly and Co., USA), and Certican® (Novartis AG, Basel,
Switzerland) that contain amorphous drug substances.
In addition to increasing the solubility and dissolution by creating an amorphous drug substance with HPMC-AS, it has also been shown to reduce in vivo precipitation of the drug by maintaining the drug as a supersaturated solution in the GI
tract. Spray-drying with the drug substance and HPMC-AS has been performed to
create amorphous particles to increase the drug exposure for rapid screening of
drugs in preclinical models (6). Electrospinning could be used in a similar fashion
and it has two additional benefits: first, the ability to create smaller-diameter fibers
than the typical particle size of spray-dried material, and second, process collects
the product onto a grounded surface, which results in very high efficiencies (99%+)
and simplified recovery.
An example of the use of electrospinning to create amorphous materials for
dissolution enhancement was performed by Verreck et al. (7). In this case, the
researchers used hydroxypropylmethylcellulose as the electrospinning polymer,
and itraconazole as the model drug. On the basis of differential scanning calorimetry,
they generated data supporting the conclusion that itraconazole was in the amorphous form, and performed dissolution studies to evaluate the rate of release. These
data also highlighted the fact that by optimizing the electrospinning conditions to
produce fibers with a diameter of 300 to 500 nm versus 1 to 4 µm, the dissolution
rate can be further increased (7).
SUSTAINED-RELEASE ELECTROSPUN FIBERS
There is significant interest in the use of electrospinning for dissolution enhancement of poorly water-soluble drugs; however, the majority of the literature on
Nanofiber-Based Drug Delivery
65
FIGURE 3 (A) Dissolution of a model low-solubility drug from polyvinylacetate electrospun fibers in
a USP 2 dissolution apparatus. (B) Scanning electron microscopy image of the nanofibers.
electrospinning of pharmaceutical or biomedical products tends to be for wound
dressings or other products which have sustained release of drug substance. This
reveals the large range of drug-release profiles which can be obtained by careful
selection of the polymer. For release of a drug substance for multiple days to months,
a polymer in the biodegradable family, such as polyglycolide, polylactic acid, or
polycaprolactone, can be used. A copolymer of ε-caprolactone and ethyl ethylene
phosphate was used to sustain the release of human β-nerve growth factor for at
least three months (8). However, selection of the polymer needs to be carefully
screened because the polymer drug compatibility has been shown to play a critical
role in the distribution of drug within the fiber (9).
Polymers which hydrate or swell but are insoluble can also be used to create
sustained-release nanofibers. Verreck et al. (10) electrospun segmented polyurethane itraconazole fibers to produce a sustained release of the drug substance. This
type of release can be achieved through the use of generally regarded as safe (GRAS)
polymers for oral drug delivery such as polyvinylacetate (PVAc) nanofibers. An
example of drug release of a poorly water-soluble drug substance from electrospun
fibers of PVAc is shown in Figure 3.
As shown by the examples above, various polymer-based methods to control
drug release such as enteric polymers, biodegradable polymers, or even polysaccharide can be utilized with electrospinning. The intrinsic properties of the polymer
form the basis for the type of drug release that occurs from the nanofibers. Although
proof of concept has been achieved with electrospun nanofibers for dissolution
enhancement and controlled drug release, the use of electrospinning in the pharmaceutical industry is still in its infancy. More advances such as functionalized nanofibers
and nanotube structures through co-axial electrospinning are likely to occur (11,12).
LARGE-SCALE MANUFACTURING
Although electrospinning is still at an early stage in the pharmaceutical industry,
commercial use of nanofibers produced by electrospinning in other industries has
already been established. Donaldson Company’s nanofiber filter media production
has increased well beyond 10,000 m2/day during the last 20 years. An SEM example of the filter utilization of a nanofiber web layer on the surface of a spunbond
nonwoven is shown in Figure 4.
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Burke and Luzhansky
FIGURE 4 Examples of nanofiber composites electrospun by Donaldson. Nanofibers on the
surface of (A) filter paper and (B) spunbond nonwoven.
By utilizing many of the current manufacturing practices used in the filtration
industry, the large-scale manufacturing of electrospun nanofibers for pharmaceutical
applications can progress at a rapid rate, including the adoption and implementation of appropriate quality control tools for nanofiber production such as automated
fiber sizing and mechanical integrity testing (13).
Automated Fiber Sizing
Full process control of the nanofiber production requires measurement and control
of fiber size, fiber size distribution, and quantity of fibers. Therefore, there is a need
for a tool that can directly measure mean fiber size and fiber size distribution.
Routine sampling and measurement of these properties are achieved using an SEM
and a developed analysis methodology. Measuring nanofiber diameters usually
consists of manually comparing the diameter of fibers in a photomicrograph to a
known scale. The process is very time-consuming, and operator consistency and
fatigue can reduce the accuracy. Automating the activity of sizing fibers is a natural
solution to the problem.
For drug-delivery applications, fiber diameter and fiber diameter distribution
will be important parameters to measure. It is expected that fiber diameter is one of
the variables that can be controlled to tune the rate of fiber dissolution and control
drug release (i.e., a bigger fiber has lower surface area and would lead to a delayed
release compared with a smaller fiber). A proprietary algorithm has been developed
at Donaldson Company to overcome the limitations of commercially available software. In the automated fiber sizing process, the SEM image is first cropped to a
desired size, and unwanted details are eliminated. The calibration bar from the SEM
image is used to set the number of pixels per micrometer. Next, the program converts the image to black and white using a gray scale function. The black (nonfiber)
areas are sorted according to size. Starting with the largest black area, a straight line
is defined on the border pixels and a diameter is drawn from one pixel across the
white area at 90°. The diameter drawing stops when another black pixel is intersected. The process is repeated around the black shape at an interval of approximately every five pixels, and then around each black area in order of size (Fig. 5).
All fiber diameter lengths are recorded and a running histogram is generated
(Fig. 6). The process that used to take hours of painstaking measurements for an
operator can now be completed in seconds. Characterization of fiber size, fiber size
distribution, and quantity of fibers is critical to ensure a quality product. In addition,
Nanofiber-Based Drug Delivery
67
FIGURE 5 Nanofiber scanning electron microscopic photograph with fiber diameter measurement
lines shown.
the mechanical integrity of a nanofiber structure can be important for sustainedrelease drug-delivery applications such as wound dressings, tissue engineering,
and regenerative medicine. Knowledge of the strain–stress characteristics is important in understanding the performance of nanofiber composites under dynamic
stress such as the gastric compression and peristaltic action of the GI tract. Owing to
the small size of fibers and extremely low weight of the layer, traditional measurement methods do not give useful results. It is anticipated that the mechanical integrity of nanofibers should be engineered, measured, and controlled.
Mechanical Integrity Testing
The DL Bending Tester was developed to study the strain properties and failure
mechanisms of nanofibers applied to the surface of another material. First, a sample
FIGURE 6
(See color insert.) Results of automated fiber size analysis.
68
Burke and Luzhansky
FIGURE 7 (A) Specimen before deformation: 1, sample; 2, jaws; 3, cylinder. (B) Specimen after
deformation.
is secured in the tester and positioned in the optical microscope. A motor then bends
and extends the sample around a cylinder with known diameter. The strain condition
approximates plane strain because the sample is thin compared with the relatively
large radius of the cylinder. Thus, differences in the strain condition between the
upper and lower surfaces of the sample are minimal. An area of approximately
6 mm × 4 mm (4 mm in the direction of stretch) is observed and can be varied. An
angle of deformation, α, is measured using a scale on the apparatus.
The basic schematic representation of the tester is shown in Figure 7A and B.
The general view of the tester is shown in Figure 8.
The strain can be calculated from the measured angle:
e=
4pαR
360L0
where ε is the calculated strain, α the measured angle, R the cylinder diameter, and
L0 the initial length of the specimen.
A camera mounted on the microscope sends a dynamic image to a monitor
that is used to observe the sample throughout the test. The first sign of relative
FIGURE 8 Dmitry Luzhansky
bending tester.
Nanofiber-Based Drug Delivery
69
movement between the components of the composite structure is an indicator of
critical strain. An operator records the relative movements and angles of the first
destruction in the nanofiber layer and full destruction of nanofiber layer. Comparison
of the angles for different composites gives a measure of integrity of the material.
CONCLUSION
Electrospun nanofibers have shown utility in a range of pharmaceutical and medical applications for immediate and controlled drug release, which will increase the
need for future process refinement and large-scale manufacturing capabilities to
convert these novel concepts into commercial products. Electrospinning represents
a nanotechnology that is steadily maturing and its use will grow as companies
implement this technology as a platform drug-delivery technique.
REFERENCES
1. McKee MG, Wilkes GL, Colby RH, Long TE. Correlations of solution rheology with
electrospun fiber formation of linear and branched polyesters. Macromolecules 2004;
37:1760.
2. Hoener BA, Benet LZ. Factors influencing drug absorption and drug availability. In:
Banker GS, Rhodes CT, eds. Modern Pharmaceutics. 4th ed. New York: Marcel Dekker,
2002 (chapter 4).
3. Ignatious F, Baldoni JM. Electrospun pharmaceutical compositions. 2001; WO 01/54667 A1.
4. Ignatious F, Sun L. Electrospun amorphous pharmaceutical compositions. 20041; WO
2004/014304.
5. Yu L. Amorphous pharmaceutical solids: preparation, characterization and stabilization.
Adv Drug Del Rev 2001; 48:27.
6. Shanker RM. Drug–polymer systems for the supersaturation of GI luminal fluid. In:
Presented at AAPS Annual Meeting, Baltimore, MD, November 7–14, 2004.
7. Verreck G, Chun I, Peeters J, Rosenblatt J, Brewster ME. Preparation and characterization
of nanofibers containing amorphous drug dispersion generated by electrostatic spinning. Pharm Res 2003; 20:810.
8. Chew SY, Wen J, Yim EKF, Leong KW. Sustained release of proteins from electrospun
biodegradable fibers. Biomacromolecules 2005; 6:2017.
9. Zheng J, Yang L, Liang Q, et al. Influence of the drug compatibility with polymer solution
on the release kinetics of electrospun fiber formulation. J Control Release 2005; 105:43.
10. Verreck G, Chun I, Rosenblatt J, et al. Incorporation of drugs in an amorphous state into
electrospun nanofibers composed of a water-insoluble, nonbiodegradable polymer.
J Control Release 2003; 92:349.
11. Casper CL, Yamaguchi N, Kiick KL, Rabolt JF. Functionalizing electrospun fibers with
biologically relevant macromolecules. Biomacromolecules 2005; 6:1998.
12. Huang Z-M, Zhang Y-Z. Micro-structures and mechanical performance of co-axial
nanofibers with drug and protein cores and polycaprolactone shells. Gaodeng Xuexiao
Huaxue Xuebao 2005; 26:968.
13. Luzhansky D. Quality control in manufacturing of electrospun nanofiber composites. In:
Presented at International Nonwovens Technical Conference, Baltimore, MD, September
15–18, 2003.
5
Drug Nanocrystals—The Universal
Formulation Approach for Poorly
Soluble Drugs
Jan Möschwitzer and Rainer H. Müller
Department of Pharmaceutical Technology, Biotechnology, and Quality Management,
Freie Universität Berlin, Berlin, Germany
INTRODUCTION
During the last two decades, many modern technologies have been established in
the pharmaceutical research and development area. The automation of the drug
discovery process by technologies such as high-throughput screening, combinatorial chemistry, and computer-aided drug design is leading to a vast number of drug
candidates possessing a very good efficacy. Unfortunately, many of these drug candidates are exhibiting poor aqueous solubility. Long before one of these compounds
can reach the market, it needs to be formulated for the pharmacological activity
tests and for the preclinical studies. The great challenge for the pharmaceutical
development is to create new formulation approaches and drug-delivery systems to
overcome solubility problems of these drug candidates which are also often associated with poor oral bioavailability (1,2).
The dissolution velocity (low solubility in general is correlated with low
dissolution velocity, law by Noyes–Whitney) and intestinal permeability are key
determinants for the bioavailability, particularly for perorally administered drugs.
To evaluate and characterize pharmaceutical compounds with respect to their
aqueous solubility and intestinal permeability, a biopharmaceutics classification
system has been developed (3,4). The system divides the drug compounds into four
classes. Poorly soluble compounds can belong to either class II or class IV. Class IV
means that the drug shows simultaneously poor solubility and low permeability.
A solubility enhancement cannot necessarily solve the bioavailability problems of
class IV drugs in any case. Drug candidates for a successful improvement of their
bioavailability by a solubilization technique belong to class II which means that
their bioavailability is only limited by their poor aqueous solubility/dissolution
velocity. The term “solubilization techniques” in the present context means technologies which increase the dissolution velocity dc/dt and—ideally—also the saturation
solubility cs.
There are many conventional approaches for the solubilization of poorly
soluble drugs. Salt formation and pH adjustment are the first attempts if the
molecule is ionizable, because in general the ionized species has a higher aqueous
solubility compared to the neutral one. If this approach fails, often cosolvents such
as propylene glycol, are used, especially for parenteral or liquid oral dosage forms.
Systemic toxicity or pain on injection are typical drawbacks associated with cosolvents (5). Other systems contain a large amount of surfactants to solubilize drugs by
an increased wetting of the hydrophobic compound. However, the extended use of
surfactants can also cause side effects. A typical example is the hypersensitivity
reaction caused by the Cremophor EL® in Taxol® (6). Another example of surfactants
71
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Möschwitzer and Müller
is mixed micelles, for example, Valium® MM (7). In case of lipophilic drugs, a lowmelting point emulsification system such as microemulsions, self-emulsifying
DDSs, or self-microemulsifying DDSs, could be used. The drawbacks of these
systems include batch-to-batch variability, chemical instabilities, and high surfactant concentration.
Another approach is the use of liposomes to incorporate hydrophobic drugs
in phospho-lipid bilayers of uni- or multilamellar vesicles. One example is the drug
Amphotericin B, which is marketed as liposomal formulation Ambisome®.
A specific approach is the formation of inclusion complexes, for example, with
cyclodextrins. Cyclodextrines are cyclic oligomers of dextrose or dextrose derivatives, which can form a reversible, noncovalent association with poorly soluble drugs
to solubilize them. Especially, the more water soluble and less toxic derivatives,
such as sulfobutylether-β-cyclodextrin (Captisol™, CyDex, Inc.) and hydroxypropyl-β-cyclodextrin (HP-β-cyclodextrin), are used in different pharmaceutical
formulations (8). Sporanox® by Johnson and Johnson/Janssen (Itraconazole/HPβ-cyclodextrine) and Zeldox® by Pfizer (Ziprasidone/SBE-β-cyclodextrine;
Captisol®) are examples of marketed products. In order to build this complex, it is
in general required that the drug molecule fits into the cyclodextrin cavity. For that
reason, this promising specific approach can be used only for a limited number of
drugs. Another drawback of this technology is the high excipient level of the
resulting product.
To sum up, the prementioned technologies are successfully applied for a number
of drugs. The marketed products prove their acceptance and applicability. The success
of a technology can be rated in two areas:
1. time between developing the technology and first market products and
2. number of products on the market in total (or more precisely, number of products
launched per year).
Based on these success criteria, the performance of most specific formulation
approaches appears rather poor. Liposomes—rediscovered by Bingham in 1968—
needed about 20 years to come to market. The number of products is relatively low,
definitely distinctly behind expectations. Similar is the situation for cyclodextrines,
currently moving forward with the new, better tolerated derivatives. These technologies are commonly used as primary strategies. However, they are more or less
specific approaches for the solubilization of a certain drug candidate. They can be
used only in compliance with certain requirements determined by the drug and
route of administration; there is no universal formulation approach.
Much smarter are nonspecific formulation approaches applicable to almost
any drug molecule (apart from a few exceptions). Particle size reduction has been a
nonspecific formulation approach for many years. The micronization of drugs is
applied to increase their surface area. Increasing the surface area will proportionally
increase the rate of dissolution and the rate of diffusion (absorption). Micronization
means transfer of relatively coarse drug powder to micrometer crystals using colloid mills or jet mills. The mean diameter of such micronized drug powders is in the
range of approximately 2 to 5 µm, corresponding to a size distribution of approximately 0.1 to 20 µm (9). Owing to their particle size distribution, such formulations
in general cannot be used for intravenous (IV) injections. Micronization cannot
improve the saturation solubility of a drug substance. In cases of practically insoluble pharmaceutical compounds or compounds of very low solubility, the effect
of micronization on the bioavailablity is not sufficient. For that reason, the next
Drug Nanocrystals
73
consequent step was to go down one further dimension in size, which means to
reduce the particle size in the nanometer range.
Since the 1980s, when drug nanoparticles were produced by List and Sucker
(10) via precipitation, various techniques for the production of drug nanocrystals
have been developed. This chapter will give an overview of the beneficial features
of drug nanocrystals, will discuss the most important production techniques, and
point out some reasons why drug nanocrystals can be seen as a universal formulation approach for poorly soluble drugs.
DEFINITION
Drug nanocrystals are pure solid drug particles with a mean diameter below
1000 nm. A nanosuspension consists of drug nanocrystals, stabilizing agents such as
surfactants and/or polymeric stabilizers, and a liquid dispersion medium. The dispersion media can be water, aqueous solutions, or nonaqueous media. The term
“drug nanocrystals” implies a crystalline state of the discrete particles, but depending on the production method they can also be partially or completely amorphous.
Drug nanocrystals have to be distinguished from polymeric nanoparticles, which
consist of a polymeric matrix and an incorporated drug. Drug nanocrystals do not
consist of any matrix material.
PHYSICOCHEMICAL PROPERTIES OF DRUG NANOCRYSTALS
The increased saturation solubility and the accelerated dissolution velocity are the
most important differentiating features of drug nanocrystals. In general, the
saturation solubility (cs) is defined as a drug-specific constant depending only on
the solvent and the temperature. This definition is only valid for drug particles with
a minimum particle size in the micrometer range. A particle size reduction down to
the nanometer range can increase the drug solubility.
The saturation solubility of solid particles depends on their particle radius
and their lattice structure according to the Ostwald–Freunlich equation and the
Kelvin equation:
⎛ S ⎞ 2υγ 2 Mγ
ln ⎜ ⎟ =
=
⎝ S0 ⎠ rRT ρ rRT
(1)
where S is the drug solubility at temperature T, S0 the solubility if r = ∞, M the
molecular weight of the compound, υ the molar volume, γ the interfacial surface
tension, and ρ the density of the compound.
From the Ostwald–Freundlich equation, it can be concluded that a drug shows
higher solubility if the particle radius is decreased. This effect is not substantial for
larger particles but will be more pronounced for particles below 1 to 2 µm, especially well under 200 nm. Another important factor influencing the solubility is the
crystalline structure of the drug. The higher the solid density and the melting point
are, the lower the solubility in general is. In contrast, a polymorph form with a lower
packaging shows a higher molar volume and lower solid density (11).
The Kelvin equation can also be used to describe the correlation of increased
saturation solubility by decreased particle size. The Kelvin equation describes the
vapor pressure as a function of the curvature of liquid droplets in a gas phase.
The vapor pressure increases with increasing curvature (decreasing particle size).
74
Möschwitzer and Müller
This can be transferred to solid drug particles in a liquid medium: the dissolution
pressure increases with decreasing particle size (12).
dcx DA
=
( cs − c x )
h
dt
(2)
where dcx/dt is the dissolution velocity, D the diffusion coefficient, A the surface of
the drug particle, h the thickness of diffusional layer, cs the saturation solubility of
the drug, and cx the concentration in surrounding liquid at time x.
The increased dissolution velocity is the characteristic feature of drug nanocrystals. The Noyes–Whitney equation [Equation (2)] describes now an increase in
dissolution velocity is proportional to an increase in surface area. For example,
when moving from a spherical 50 µm particle to micronized 5 µm particles, the total
surface area enlarges by a factor of 10, moving to 500 nm nanocrystals by a factor of
100. The decrease in the diffusional distance h is an additional factor accelerating the
dissolution velocity. According to the Prandtl equation [Equation (3)], the diffusional
distance h is reduced with increasing curvature of ultrafine particles. Together with
the increased saturation solubility of ultrafine particles, the concentration gradient
in the Noyes–Whitney equation is significantly increased. For that reason, nanonization can distinctly increase the dissolution velocity of poorly soluble drugs:
⎛ L1/ 2 ⎞
hH = k ⎜ 1/ 3 ⎟
⎝V ⎠
(3)
where hH is the hydrodynamic boundary layer thickness, V the relative velocity of
the flowing liquid against a flat surface, k a constant, and L the length of the surface
in the direction of flow.
Drug nanocrystals can be used for a chemical stabilization of chemically labile
drugs. The drug paclitaxel can be preserved from degradation when it is formulated
as a nanosuspension (13,14). The same result was found for the chemically labile
drug omeprazole. When formulated as a nanosuspension, the stability was distinctly
increased in comparison to the aqueous solution (15). The increased stability can be
explained by a shield effect of the surfactants and the drug protection by a monolayer made of degraded drug molecules which reduce the accessibility for destructive agents (16).
POTENTIAL CLINICAL ADVANTAGES OF DRUG NANOCRYSTALS
Application for Oral Delivery
The oral route is the most important and preferred route of administration. The
formulation of drug nanocrystals can impressively improve the bioavailability of
perorally administered poorly soluble drugs. In 1995, Liversidge and Cundy (17)
reported an increase in bioavailability for the drug Danazol from 5.1 ± 1.9% for the
conventional suspension to 82.3 ± 10.1% for the nanosuspension. The increased
dissolution velocity and saturation solubility lead to fast and complete drug
dissolution, an important prerequisite for drug absorption.
Whenever a rapid onset of a poorly soluble drug is desired, the formulation of
drug nanocrystals can be beneficial, for example, in case of analgesics. The analgesic
naproxen, formulated as a nanosuspension, has shown a reduced tmax but simultaneously approximately threefold increased AUC in comparison to a normal suspension (Naprosyn®) (18).
Drug Nanocrystals
75
Besides the faster onset of action, the naproxen nanosuspension has also shown
a reduced gastric irritancy (19,20). If absorption windows limit the drug absorption
or by food effects, drug nanocrystals have advantages in comparison to conventional suspensions. Wu et al. have reported reduced fed-fasted ratio and an improved
bioavailability for nanocrystalline aprepitant (MK-0869), the active ingredient in
Emend®, in beagle dogs.
Another important advantage of drug nanocrystals is their adhesiveness and
the increased residence time, which can positively influence the bioavailability. The
mucoadhesiveness can be raised by the use of mucoadhesive polymers in the
dispersion medium (21,22). Additionally the utilized mucoadhesive polymers can
prevent the drug from degradation. The reduced particle size can be also exploited
for improved drug targeting, as reported for inflammatory tissues (23) or the
lymphatic drug uptake (24).
Parenteral Administration of Drug Nanocrystals
The parenteral application of poorly soluble drugs, particularly intravenous (IV)
administration of practically insoluble compounds, using cosolvents, surfactants,
liposomes, or cylcodextrines, is often associated with large injection volumes or
toxic side effects. Carrier-free nanosuspensions enable potential higher loading
capacity compared to other parenteral application systems. Using nanosuspensions, the application volume can be distinctly reduced compared to solutions (15). To
fulfil the distinctly higher regulatory hurdles, the drug nanocrystals need to be
produced in an aseptic process. Alternatively, nanosuspensions can be sterilized by
autoclaving (25) or alternatively by gamma irradiation as well as sterile filtration
(26). When a drug is administered as a nanosuspension, the rapid dissolution of the
nanocrystals will mimic the plasma concentration profile of a solution. Drug nanosuspensions can be formulated with accepted surfactants and polymeric stabilizers
for IV injection. In contrast, solutions of poorly soluble drugs require the use of
cosolvents and/or high surfactant contents (e.g., Chremophor EL in Taxol®), which
can cause undesired side effects (6). Comparing the toxicity of Taxol® with a paclitaxel nanosuspension, the latter has shown a distinctly reduced toxicity. The
nanosuspension was much better tolerated, resulting in an approximately doubled
LD50 value (27). The same effect of increased tolerated dose was found for the
antifungal drug itraconazole. Itraconazole is marketed as Sporanox IV® by Janssen
Pharmaceutica Products, L.P., an inclusion complex of itraconazole and 2-hydroxypropyl-β-cyclodextrine (HP-β-CD). The product exhibits a significant acute toxicity above 10 mg/kg and an LD50 value lower than 40 mg/kg when administered
as a bolus in the caudal vein of rats. In contrast, a 1% nanosuspension of itraconazole could be administered up to 320 mg/kg without animal mortality. Besides
the decreased acute toxicity, the nanosuspension has also shown a prolonged effect,
whereby the administration intervals could be extended almost three times in comparison to the daily administration of Sporanox IV® (11). Comparing a clofazimine
nanosuspension with a liposomal formulation, both are similarly effective in the
treatment of artificially induced Mycobacterium avium infections. The targeting to
the reticuloendothelial system, the lung, liver, and spleen was comparable to the
liposomal formulation (28). Furthermore, a special targeting can be achieved by a
surface modification using the concept of “differential protein adsorption.” A surface
modification of drug nanocrystals with the surfactant Tween 80 leads to a preferential
adsorption of apolipoprotein E. This protein adsorption enables a targeted delivery
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of drug nanocrystals to the brain. Atovaquone drug nanocrystals modified with
Tween 80 have shown an excellent efficacy in the treatment of Toxoplasmosis (29).
Drug Nanocrystals for Pulmonary Drug Delivery
Delivery of water-insoluble drugs to the respiratory tract is very important for the
local or systemic treatment of diseases. Many important drugs for pulmonary delivery show poor solubility simultaneously in water and nonaqueous media, for
example, important corticosteroids such as budesonide or beclomethasone dipropionate. In the past, most of these drugs were administered as aerosols, but in compliance
with the Montreal Protocol of 1987 the use of chlorofluorocarbon (CFC) must be avoided. Therefore, alternatives such as dry powder inhalers or metered dose inhalers
without CFC (MDI) were developed. These systems are filled with micronized drug
powders produced by jet-milling. The mean particle size in the lower micrometer
range (3–25 µm) results in a significant oropharyngeal deposition of larger particles
leading to increased occurrence of candidasis. Additionally, the oral deposition of the
drug leads to gastrointestinal (GI) drug absorption followed by systemic side effects
(30). Nanosuspensions can be successfully applied to overcome these problems. The
nebulization of nanosuspensions generates aerosol droplets of the preferred size
loaded with a large amount of drug nanocrystals. Using nebulized nanosuspensions,
the respirable fraction is distinctly increased in comparison to conventional MDIs (30).
The smaller the particle size of the drug nanocrystals, the higher the drug loading of
the aerosol droplets (31,32). Therefore, the required nebulization time is distinctly
reduced (33). Besides this, drug nanocrystals show an increased mucoadhesiveness,
leading to a prolonged residence time at the mucosal surface of the lung.
Other Administration Routes
Dermal nanosuspensions are mainly of interest if conventional formulation approaches fail. The use of drug nanocrystals leads to an increased concentration gradient
between the formulation and the skin. The increased saturation solubility leads to
“supersaturated” formulations, enhancing the drug absorption through the skin. This
effect can further be enhanced by the use of positively charged polymers as stabilizers
for the drug nanocrystals. The opposite charge leads to an increased affinity of the
drug nanocrystals to the negatively charged stratum corneum (unpublished data).
The ocular delivery of nanoparticles, including drug nanocrystals, is also of
high interest. The development of such colloidal delivery systems for ophthalmic
use aims at dropable dosage forms with a high drug loading and a long-lasting drug
action. The adhesiveness of the small nanoparticles, which can be further increased
by the use of mucoadhesive polymers, leads to a more consistent dosing. Blurred
vision can be reduced by the use of submicron-sized drug particles (34,35).
PARTICLE SIZE REDUCTION TECHNIQUES
Nanoparticles Produced by Media Milling Processes
The use of media mills for the production of ultrafine particles is very common.
From the first half of the twentieth century, ball mills were known for the production
of ultrafine particle suspensions (36). In 1991, Liversidge et al. (37) have adapted
this technique for the production of surface-modified drug nanoparticles. In order
to produce nanocrystalline dispersions by the NanoCrystals® technology, a milling
chamber is charged with milling media, dispersion medium (normally water), stabilizer, and the drug. The drug particles are reduced in size by shear forces and
Drug Nanocrystals
77
forces of impaction generated by a movement of the milling media. Small milling
pearls or larger milling balls are used as milling media. With a reduction in the size
of grinding media in a media mill, the number of contact points is increased exponentially, resulting in improved grinding and dispersing action (i.e., leading to smaller
particles). The pearls or balls consist of ceramics (cerium- or yttrium-stabilized zirconium dioxide), stainless steel, glass, or highly cross-linked polystyrene resin-coated
beads. A problem associated with the pearl milling technology is the erosion from the
milling material during the milling process. Buchmann et al. (38) reported the formation of glass microparticles when using glass as milling material. In order to reduce the
quantity of impurities caused by an erosion of the milling media, the milling beads
were coated with highly cross-linked polystyrene resin (39). A perpetual problem is the
adherence of product to the large inner surface area of the milling system. The inner
surface area is made up of the surface area of the chamber and of all milling beads
together. Even in recirculation systems, this product adherence causes a product loss.
Of course, this undesirable drug loss can be an issue in very expensive drugs, especially when very small quantities of new chemical entities (NCEs) are processed.
In general, there are two milling principles: either the complete container is
moved in a complex movement leading consequently to movement of the milling
material or the milling medium is moved by an agitator. Assume that 76% of the
milling chamber volume will be filled with milling material (larger batches are difficult to produce in a batch mode). In a 1000-L mill, this corresponds to 760-L milling
material, based on the apparent density of zircon oxide pearls being 3.69 kg/L
which corresponds to 2.8 tons of milling material. For that reason, agitator bead mills
in recirculation mode are preferred for the production of larger batches. Figure 1
shows a cross-section of an agitator bead mill in a horizontal arrangement. The
suspension is pumped vertically through the bead mill. In this case, the product is
separated from the milling media by a dynamic separator gap. In other cases,
separator cartridges are used. The recirculation mode prolongs the required milling
time, because the residence time of the drug particles under impaction of shear
forces is reduced. The milling time depends on many factors, such as hardness of
the drug, surfactant content, temperature, viscosity of the dispersion medium,
specific energy input, and size of the milling media. Milling periods from 30 minutes
up to several days are reported (18). Milling equipment is now available from the
labscale (41) to the large production scale. Besides other factors, this is an important
prerequisite for the commercial production of nanocrystalline drugs. In 2000,
FIGURE 1 Agitator bead mill with
dynamic separator gap (continuous
mode): 1, product inlet; 2, product
outlet; 3, cooling jacket; 4, milling
pearls; 5, rotor with grinding disks;
6, dynamic separator gap. Source:
From Ref. 40.
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Möschwitzer and Müller
Rapamune® was launched by Wyeth as the first product containing sirolimus
NanoCrystals®. The coated Rapamune® tablets are more convenient and show a
27% increased bioavailability compared to the Rapamune® solution (42). This is an
example to compare two formulation strategies. The oral solution shows the principles of cosolvents and surfactants, whereas the tablets shows the nice performance of
a particle size reduction technique. Emend® is the second product incorporating the
NanoCrystal® technology. It was introduced to the market in 2003 by Merck.
Emend® is a capsule containing pellets of nanocrystalline aprepitant, sucrose, microcrystalline cellulose, hyprolose, and sodium dodecylsulfate (43). The third product
is TriCor, a nanocrystalline fenofibrate tablet marketed in 2004 by Abott. Megaace
ES, an oral suspension containing megestrol acetate for the treatment of HIV-associated anorexia and cachexia, was launched as a fourth product late in the middle of
the year 2005.
To sum up, the wet media milling is a viable particle size reduction technology. The performance has been proven by four FDA-approved products. But other
nanonization technologies also can benefit from the success of the NanoCrystal®
technology by the attraction of the size reduction principle in general.
Precipitation Methods
The classical precipitation process, known as “via humida paratum,” is actually a
very old pharmaceutical procedure. Later, this basic idea was applied for the
production of nanocrystalline drug particles (44). The first application of the precipitation technique was developed by List and Sucker (10). It is known as hydrosol
technology, and the IP is owned by Sandoz (now Novartis). A poor water-soluble
drug is dissolved in an organic medium, which is water-miscible. A pouring of this
solution into a nonsolvent, such as water, will cause a precipitation of finely dispersed
drug nanocrystals. As simple as the particle formation process is, the preservation
of the nanocrystalline particle size is difficult. The fine particles tend to grow up,
driven by a phenomenon called “Ostwald ripening.” This is a process where small
particles are dissolved in favor of larger particles. Sucker (45) suggested immediate
lyophilization to preserve the particle size.
The crystalline state of the particles obtained by the precipitation process can
be controlled. Depending on the employed method, amorphous drug nanoparticles
can also be generated (46). Beta-carotene is dissolved in a water-miscible organic
solvent together with digestible oil. This solution is admixed to an aqueous solution
of a protective colloid (gelatine) causing a precipitation of amorphous nanoparticulate beta-carotene. After an annealing step and spray-drying, a stable amorphous
product can be obtained. This NanoMorph® technology, invented by Auweter et al.,
is used by the company Soliqs.
Another approach to preserve the size of the precipitated nanocrystals is the
use of polymeric growth inhibitors, which are preferably soluble in the aqueous
phase. The increased viscosity of the aqueous phase can reduce particle growing.
The resulting suspension is subsequently spray-dried to obtain a dry powder with
a relatively high drug loading (47). Using this technique, a tremendous increase in
dissolution rate (from 4% to 93% within 20 minutes) was shown for a poor watersoluble drug ECU-01 (48). Although the feasibility of preparing drug nanocrystals
by precipitation has been shown by many groups, no commercial drug product
using this technology has entered the market. To use the prementioned methods, it
is required that the drug is soluble in at least one water-miscible solvent. This is
Drug Nanocrystals
79
often not the case for NCEs. Many drugs are simultaneously poorly soluble in
aqueous and nonaqueous media. Even if there is a suitable solvent available, it is
difficult to remove this solvent completely. Solvent residues can be potential risk
factors for drug alteration and toxic side effects. In addition, in cases of amorphous
drug nanoparticles, it is seen as very critical to preserve the amorphous character
throughout the shelf life of a product. Recrystallization would impair the oral
bioavailability. This effect is less critical in food products because of less strict
regulatory requirements that allow more tolerance.
Nanoparticles Produced by High-Pressure Homogenization
High-pressure homogenization is another universal approach to reduce the particle
size of poorly soluble compounds. Considering the homogenization equipment and
the homogenization conditions, it has to be divided between three technologies.
Microfluidizer Technology (IDD-PTM Technology)
Particles can be generated by a high shear process using jet-stream homogenizers,
such as Microfluidizers (Microfluidizer®, Microfluidics, Inc.). A frontal collision of
two fluid streams under pressures up to 1700 bar leads to particle collision, shear
forces, and also cavitation forces. To preserve the particle size, stabilization with phospholipids or other surfactants and stabilizers is required. A major disadvantage of
this process is the required production time. In many cases, 50 to 100 time-consuming
passes are necessary for a sufficient particle size reduction (49,50). SkyePharma
Canada, Inc. (previously RTP, Inc.) applies this principle for its IDD-P™ technology
to produce submicron particles of poorly soluble drugs (51).
Piston-Gap Homogenization in Water (Dissocubes® Technology)
Drug nanocrystals can also be produced by high-pressure homogenization using
piston gap homogenizers. Depending on the homogenization temperature and the
dispersion media, there is a difference between the Dissocubes® technology and the
Nanopure® technology. The Dissocubes® technology was developed by Müller et al.
in 1994 (52) and later acquired by SkyePharma PLC. It involves the production of
nanosuspensions in water at room temperature. The trade name Dissocubes®
already indicates the improved dissolution behavior and the cubic shape of the
resulting drug nanocrystals (53). A drug powder is dispersed in an aqueous surfactant solution. The resulting macrosuspension is subsequently forced by a piston
through a tiny homogenization gap applying pressures up to 4000 bar.
Depending on the viscosity of the suspension and the homogenization pressure, the width of the homogenization gap is in the range of 5 to 20 µm (54). Figure
2 shows a cross-section through a piston-gap homogenizer. According to Bernoulli’s
law, the resulting high streaming velocity of the suspension causes an increase in
dynamic pressure that is compensated by a reduction in static pressure below the
vapor pressure of the aqueous phase. The water starts boiling and the formation of
gas bubbles occurs. These gas bubbles collapse immediately when the liquid leaves
the homogenization gap, resulting in cavitation-caused shock waves. The enormous
power of these shock waves, turbulent flow, and shear forces break the drug
particles (55).
The detailed illustration shows the principle of diminution mechanisms in the
homogenization gap (Fig. 3). At the beginning, the particles are broken at crystal
imperfections; with continuing the homogenization process, the number of
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Möschwitzer and Müller
FIGURE 2 Cross-section of a piston-gap homogenizer. A macrosuspension is forced through a
very tiny homogenization gap in order to produce drug nanocrystals. Source: From Ref. 40.
imperfections decreases and almost perfect small crystals will remain. The required
number of cycles is mainly influenced by the hardness of the drug, the finesse of the
starting material and the requirements of the application route or the final dosage
form, respectively. In general, 10 to 20 homogenization cycles are sufficient to obtain
a unimodal size distribution in the nanometer range (56–58).
The use of water as dispersion medium is associated with some disadvantages. Hydrolysis of water-sensitive drugs can occur, as well as problems during
drying steps. In cases of thermolabile drugs or drugs possessing a low melting
point, a complete water removal requires relatively expensive techniques, such as
lyophilization. For these reasons, the Dissocubes® technology is particularly suitable if the resulting nanosuspension is directly used without modifications, such as
drying steps.
FIGURE 3 Diminution principles
during high-pressure homogenization
(detailed illustration of a homogenization gap in cross-section): 1, implosion
area (cavitation); 2, boiling area and
crystal collision; 3, shear forces; 4,
turbulent flow. Source: From Ref. 40.
Drug Nanocrystals
81
Nanopure® Technology
In 1999, Müller et al. (59) found that a similar effective particle diminution can also be
obtained in nonaqueous or water-reduced media. The proprietary technique is known
as Nanopure® technology developed and owned by PharmaSol GmbH/Berlin. By
using dispersion media with a low-vapor pressure and performing the homogenization process at low temperatures (e.g., 0°C), the cavitation in the homogenization
gap is distinctly reduced or does not exist at all. It could be shown that even in the
absence of cavitation, a sufficient particle size diminution was obtained. The turbulent flow and shear forces during the homogenization process are strong enough to
break the drug particles and to produce drug nanocrystals. The high-pressure homogenization in nonaqueous or water-reduced media is particularly beneficial if the
nanosuspension has to be transferred into a traditional final dosage form. By reducing
the water content in the dispersion medium, the required energy is minimized for
drying steps, such as spray-drying, fluidized bed drying, or upon suspension layering onto sugar spheres. The evaporation processes can be performed under milder
conditions, which is beneficial for temperature-sensitive drugs. Production of nanosuspensions at 0°C or even below can prevent temperature labile drugs from
degradation (15). If the high-pressure homogenization is carried out completely in
nonaqueous media, even water-sensitive drugs can be processed without hydrolysis.
Nanosuspensions produced in liquid polyethylene glycol (PEG) or hot-melted PEGs
can be directly filled into gelatine or hydroxy propyl methyl cellulose (HPMC) capsules (60). Depending on the requirements of the final formulation, the water content
can be varied from water-free to isotonic conditions for intravenous suspensions.
Irrespective of the employed technique (Dissocubes® or Nanopure®), the
production of drug nanocrystals by high-pressure homogenization is a productionfriendly process. Homogenization is a low-cost process; approved production lines
are already in use for the production of pharmaceutical products, such as emulsions
for parenteral nutrition (12). The process can be easily transferred from the labscale
to the large production scale. High-pressure homogenization can be performed
starting from 0.5 mL (Avestin EmulsiFlex-B3, Avestin, Inc., Canada) up to large
batch sizes of 2000 L/hr (Rannie 118, APV homogenisers, Denmark) (26,61).
When producing nanosuspenions by high-pressure homogenization even at
hard homogenization conditions, only a noncritical product contamination below
1 ppm was observed (62). Suspensions with drug concentration up to 30% and more
can be easily processed by high-pressure homogenization (63).
Combination Technologies
Nanoedge® Technology
Baxter’s NANOEDGE® process relies on the combination of a microprecipitation
technique with a subsequent annealing step by applying high shear and/or thermal
energy (64). A fine suspension is formed by adding an organic solution of the waterinsoluble drug to an antisolvent, for example, aqueous surfactant solution.
Depending on the precipitation conditions, either small amorphous or crystalline
drug particles in the nanometer range or friable needle-like crystals in the micrometer range are formed. Consequently, the following high-energy input can have two
effects on the preformed particles. Small amorphous or crystalline drug particles
will be preserved in size by an annealing step without changing the mean diameter.
It could be shown that the tendency to crystal growth can be reduced by energy
input after the precipitation step. In case long friable needle-like crystals are obtained, they will be reduced in size by the high-energy input using high-pressure
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Möschwitzer and Müller
homogenizers. According to the patent (64), particle sizes in the range of 400 to
2000 nm can be obtained. The organic solvent utilized has to be carefully removed
from the final nanosuspension without changing the particle size of the drug nanocrystals. Otherwise, crystal growth can be promoted by an increased solubility of
the drug. Any content of the organic solvent dissolved in the aqueous phase can act
as a “cosolvent,” leading to an increased tendency to Ostwald ripening. Also, toxic
effects can be caused by potential solvent residues, especially if the nanosuspension
is the final product. For these reasons, the NANOEDGE® process is particularly suitable for drugs that are soluble in nonaqueous media possessing low toxicity, such
as N-methyl-2-pyrrolidinone.
Nanopure® XP Technology
Considering the commonly used particle size reduction techniques, the production
of nanosuspensions is in most cases associated with a high-energy input (highspeed media mills, high-pressure homogenization) and a relatively long period
between the drug synthesis and the final product. A micronized drug material (size
10–100 µm) is recommended for milling and high-pressure homogenization processes (37,55). Therefore, the drug often has to be jet-milled before the nanonization.
Impurities caused by an abrasion from the milling material or solvent residues from
the precipitation process are undesirable. They can causes side effects especially if
the drug is administered for the treatment of chronic diseases. The minimal achievable particle size is significantly determined by the hardness of the drug. In cases
of very hard drugs, an increasing energy input (by extending the milling time or the
number of homogenization cycles) will not lead to a smaller particle size. In general,
the smallest achievable size of nanocrystals is around 200 nm; only under special
conditions can about 100 nm be produced. However, especially crystals below
100 nm would show an extremely fast dissolution and simultaneously a great
increase in saturation solubility (11). Therefore, such particles are of high commercial interest.
In 2005, Möschwitzer (65) developed a new combination method for the
production of drug nanosuspensions, which is now owned by PharmaSol GmbH.
The Nanopure® XP technologogy (process variant: H42) enables extension of the
performance of the Nanopure® technology to very hard and crystalline materials.
Modification of the starting material by an evaporation process before the subsequently performed high-pressure homogenization can significantly reduce the
number of homogenization cycles (66). Owing to the reason that the solvent will be
removed completely before homogenization by the evaporation process, various
solvents can be used for the modification process without restrictions due to toxicity
reasons. Additionally, excipients can be added to the drug solution to increase the
number of crystal imperfections upon drying. Figure 4 makes clear the effectiveness
of the new combination method in comparison to the classical high-pressure homogenization. The application of the H42 technology leads to distinctly reduced particle
size by a simultaneous reduction of the required number on homogenization cycles.
This will consequently reduce the production costs and the wear on the homogenization equipment.
H96 is another technology belonging to the Nanopure® XP platform. The technology, developed in 2005 by Möschwitzer and Lemke (67), is a modified process
based on high-pressure homogenization. By using this undisclosed technology, it
was shown that drug nanocrystals below 100 nm can be produced by high-pressure
homogenization.
Drug Nanocrystals
83
FIGURE 4 Comparison of the new
homogenization technology H42 (left )
with the conventional homogenization
in water (right ), performed in pistongap homogenizer, influence of cycle
numbers, results represent the laser
diffractometry diameters (volumeweighted, Coulter Ls 230, BeckmanCoulter, Germany). Source : From
Ref. 40.
H96 results in a translucent nanosuspension in comparison to another nanosuspension produced using the Dissocubes® technology. Figure 5 shows two nanosuspensions. The left nanosuspension was produced applying the conventional
high-pressure homogenization in water; the right nanosuspension was produced by
using the H96 technology. Although both formulations are being composed similarly,
the right formulation is translucent due to the significantly reduced particle size.
The particle size analysis also provides the evidence of the performance of the H96
method. Almost 99% of the particles were smaller than 100 nm [laser diffractometry
(LD) D99%, volume size distribution, LS 230, Beckman-Coulter, Germany unpublished data with permission from Ref. 40]. A very small particle size and a narrow
range of particle size distribution can be obtained by using the H96 technology.
Other Techniques for the Production of Drug Nanocrystals
Milling, high-pressure homogenization, and precipitation are the main methods
employed for the production of drug nanocrystals. The importance for improvement
of the bioavailability of poorly soluble drugs by the production of drug nanocrystals
is widely accepted. The intensive research for new technologies led to many other
FIGURE 5 (See color insert.) Two
nanosuspensions composed similarly,
produced by high-pressure homogenization, conventional method (left )
versus translucent nanosuspension
(particle size well below 100 nm)
resulting from H96 technology. The
red laser beam is reflected by the tiny
nanocrystals. Source : From Ref. 40.
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Möschwitzer and Müller
approaches for the production of drug nanocrystals. Even nonpharmaceutical companies, such as Dow Chemical, are entering the market of poorly soluble drugs with
solubility-enhancing technologies. Spray-freezing into liquid (68) and evaporative
precipitation into aqueous solutions (69) are examples for such new technologies.
FINAL FORMULATIONS FOR DRUG NANOCRYSTALS
In order to show their advantages in vivo, the drug nanocrystals need to be
transferred into the right dosage form. Nanosuspensions can be directly used as
oral suspensions to overcome the difficulties of swallowing tablets by pediatric or
geriatric patients. One example is Megace® (Bristol Meyers Squibb), an oral suspension of megestrol acetate, used for the treatment of HIV-associated anorexia and
cachexia. The application of these nanosuspensions can improve the solubility of
the drug and the dissolution rate; additionally, suspensions can be applied for reasons of taste-masking. Nanosuspensions can also be used directly for parenteral
drug administration. Although nanosuspensions have shown a sufficient long-term
stability without Ostwald ripening, for intravenous products a lyophilization step is
recommended in order to avoid aggregation or caking of settled drug nanocrystals.
The lyophilized product can be easily reconstituted before use by adding isotonic
water, aqueous glucose solution, or other reconstitution media (27,70).
Without question, both the patients and the marketing experts prefer the oral
administration of traditional dosage forms. Hence, to enter the pharmaceutical market
successfully in most cases drug nanocrystals have to be formulated as traditional products, such as tablets or capsules. A perfect solid dosage form should preserve the in
vivo performance of drug nanocrystals. When reaching the target part of the GI tract,
the dosage form should release the drug nanocrystals as a fine, nonaggregated suspension. Otherwise, self-agglomeration or aggregation can impair the drug release (71).
Using nanosuspensions as granulation fluid for a further tablet production is
a very simple approach. The nanosuspension is admixed to binders and other excipients, and the granules are finely compressed to tablets. This dosage form is limited in the maximum achievable drug content. A maximum drug content of about
50% or less is suggested in order to ensure a complete disintegration into a finely
dispersed suspension (72). Nanosuspensions can also be used for the production of
matrix pellets (Fig. 6) or as layering dispersions in a fluidized bed process. After the
pellet preparation, the cores can be coated with several polymers in order to modify
the release profile of the final formulation (73–75).
A very smart formulation approach is the Nanopure® technology. Nanocrystals
produced in nonaqueous media, such as liquid PEG or oils (e.g., Miglyol), can be
directly filled into gelatine or HPMC capsules. The production of drug nanocrystals
in melted PEGs is a new strategy for the production of final dosage forms containing drug nanocrystals. After performing the high-pressure homogenization in
melted PEG at about 60°C, the mixture can be solidified. The resulting matrix, fixing
the drug nanocrystals in separated state, can be compressed to tablets or directly
filled into capsules (76).
Spray-drying of the nanosuspensions is another cost-effective approach to
transfer nanosuspensions into dry products. The drug nanocrystals can directly be
produced by high-pressure homogenization in aqueous solutions of water-soluble
matrix materials, for example, polymers [polyvinylpyrrolidone, polyvinylalcohol or
long-chained PEG, sugars (saccharose, lactose) or sugar alcohols (mannitol, sorbitol)].
Afterwards, the resulting nanosuspension can be spray-dried under appropriate
Drug Nanocrystals
85
FIGURE 6 SEM photograph of uncoated matrix core containing drug nanocrystals: (left ) overview
(magnification 60×), (right ) detailed magnification (1000×) showing drug nanocrystals combined
with the binder material.
conditions. The dry powder, composed of drug nanocrystals embedded in a watersoluble matrix, can be filled in hard gelatine capsules for oral administration (77).
Another attractive approach using the spray-drying principle is described as
“direct compress technology” (78). Lactose and other matrix-forming materials,
such as micronized polymer powders or lipids, are admixed to the prior-produced
nanosuspension. The resulting suspension is transferred into a drug–matrix–compound by spray-drying. Subsequently, the free-flowable powder can be used for
direct compression of fast dissolving or prolonged release tablets. Alternatively, the
powder can also be filled into hard gelatine capsules.
CONCLUSION
Poor aqueous solubility is clearly recognized by the pharmaceutical industry as a
major problem. The use of drug nanocrystals is a universal formulation approach to
increase the therapeutic performance of these drugs in any route of administration.
Almost any drug can be reduced in size to the nanometer range. Owing to their
great formulation versatility drug nanocrystals are no longer only the last chance
rescue for a few drugs. Many insoluble drug candidates are in clinical trials formulated as drug nanocrystals (at present about 10).
Various nanonization techniques are available. Production facilities are available to produce tons of nanosuspensions. Currently, attention is turned to improving
the diminution performance to produce drug nanocrystals well below 100 nm, also in
cases of very hard drugs. First approaches were already successful. New technologies
are in development to produce final dosage forms with higher drug loadings, better
redispersability at their site of action, and an improved drug targeting.
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6
Lipid-Based Nanoparticulate
Drug Delivery Systems
Jun Wu and Xiaobin Zhao
Division of Pharmaceutics, College of Pharmacy, The Ohio State University,
Columbus, Ohio, U.S.A.
Robert J. Lee
Division of Pharmaceutics, College of Pharmacy, NCI Comprehensive Cancer
Center, NSF Nanoscale Science Engineering Center for Affordable
Nanoengineering of Polymeric Biomedical Devices, The Ohio State University,
Columbus, Ohio, U.S.A.
INTRODUCTION
Phospholipids, upon hydration, spontaneously form bilayer membrane vesicles
(liposomes) or may act as surfactants in forming micro- or nanoemulsions or
solid–lipid nanoparticles. Phospholipids, triglycerides, and cholesterol are the main
ingredients of liposomes and lipid nanoparticles. They are natural components of
biological membranes and lipoproteins and are, therefore, presumed to be highly
biocompatible (1). Drugs and cell-targeting ligands can be incorporated into these
structures by encapsulation (for hydrophilic molecules), lipid-phase solubilization
(for lipophilic molecules), conjugation to a lipid anchor (as a lipid-derivatized prodrug), or electrostatic complexation (for poly-anionic molecules such as nucleic
acids), depending on their specific physicochemical properties. Liposomes and lipid
nanoparticles smaller than 300 nm are potentially suitable for systemic administration. In this chapter, various aspects of these nanoparticulate systems are discussed
in the context of drug delivery.
LIPOSOMES
Liposomes are phospholipid bilayers with an entrapped aqueous volume. On the
basis of the number of layers (lamellarity) and diameter, liposomes are classified
into multilamellar vesicles (MLVs, diameter >200 nm), large unilamellar vesicles
(diameter 100–400 nm), and small unilamellar vesicles (diameter <100 nm). On the
basis of surface charge (zeta potential), they are classified into cationic, neutral, and
anionic liposomes. In addition, liposomes can be deliberately engineered to possess
unique properties, such as long systemic circulation time, target cell specificity, pH
and reductive environmental sensitivity, and temperature sensitivity. These are
achieved by selecting the appropriate lipid composition and surface modification
for the liposomes.
METHODS OF LIPOSOME PREPARATION AND CHARACTERIZATION
Liposomes form spontaneously when phospholipids are hydrated. Additional steps
are often necessary to modify the size distribution and lamellarity of liposomes.
Several methods have been established for liposome preparation based on the scale
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of the preparation and other considerations, such as drug encapsulation efficiency.
The first step in liposome preparation is to dissolve lipid ingredients in a suitable
solvent; this is then followed by lipid hydration and particle size reduction. For
example, lipids are first dissolved in chloroform/methanol and dried into a thin
film on a rotary evaporator and are then hydrated in an aqueous buffer above the
phase-transition temperature (2). This process will result in the formation of heterogeneous MLVs. The lamellarity of the MLVs can be reduced by repeated cycles of
freezing and thawing. If phospholipids are dissolved in a water-miscible solvent
such as ethanol and rapidly diluted into an aqueous buffer, liposomes with relative
small particle sizes can be generated directly. This is then followed by removal of
the solvent from the liposome preparation and/or further size-modifying steps (3).
Particle size can then be reduced by a number of methods, including sonication,
extrusion, and homogenization. Sonication, typically using an ultrasonic probe
sonicator, can reduce liposome particle size by inducing cavitation (4). Extrusion is
performed using a high-pressure filter unit, such as the Lipex™ extruder by Northern
Lipids, Inc., containing a track-etched polycarbonate membrane at a temperature
above the phase-transition temperature of the liposomal bilayer (5). The polycarbonate membrane has straight through cylindrical pores of precise diameters and can
withstand pressure of 3000 psi with proper support. Liposomes extruded through the
polycarbonate membranes typically have narrow particle size distribution. Largescale production of liposomes is possible using the extrusion method employing
high-surface-area extrusion filter units. Another potentially scalable method for
liposome particle size reduction is high-pressure homogenization usually at pressures
above 5000 psi (2). This method can be used for continuous production of liposomes
or nanoparticles at large scale. In the labscale, liposomes can be synthesized by detergent dialysis in which lipids are first solubilized in a dialyzable detergent, such as
octylglucoside and then dialyzed against a buffer (6). Alternatively, reversed-phase
evaporation method may be used to form a water-in-oil emulsion in a volatile organic
solvent followed by phase inversion upon solvent removal. This method was designed
to maximize entrapment efficiency of water-soluble agents (7).
Methods suitable for drug loading into liposomes depend on the properties of
the drug. Lipophilic drugs can be codissolved with the lipids during liposome preparation. Hydrophilic drugs can be passively entrapped into liposomes during
liposome formation. Alternatively, drugs can be incorporated into liposomes by a
pH gradient-driven remote loading procedure. For example, an inward-directed pH
gradient could be established by entrapment of a low pH buffer (e.g., pH 4 sodium
citrate) or by entrapping ammonium sulfate followed by external buffer exchange
resulting in reduction of intraliposomal pH. Addition of a weakly basic drug, such
as doxorubicin and vincristine, results in near-quantitative loading of the drug into
the liposomes due to intraliposomal protonation of the drug molecules and complexation with entrapped counterion (8,9). For polyelectrolytes such as DNA, loading
into liposomes can be achieved by electrostatic complexation with incorporation of
cationic lipids into the liposome composition (10,11). Structure of DNA complexes
of cationic liposomes may not be similar to the structure of typical liposomes and is
highly dependent on lipid head group structure, cationic-to-anionic charge ratio,
and kinetics of complex formation (12).
Liposomes can be purified by a number of methods. At the labscale, liposomes
can be purified based on their size by high-speed centrifugation, size exclusion
chromatography, or dialysis (3). At a larger scale, liposomes can be purified by
tangential flow diafiltration (13). Liposomes can also be lyophilized in the presence
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of a cryoprotectant, typically a disaccharide such as sucrose, lactose, or trehalose,
which can prevent vesicle fusion and particle size increase, although significant
leakage of aqueous content may occur upon rehydration of the liposomes (14).
A liposomal formulation can be characterized by a number of established methods. First, particle size distribution can be measured by dynamic light scattering, by
cryo- or negative-staining electron microscopy, or by atomic force microscopy. Surface
electrical property of liposomes can be measured by zeta potential measurement. Other
useful analyses include colloidal stability and rate of drug release in storage and in
plasma by dialysis, kinetics of uptake, and internalization of fluorescence labeled liposomes in cultured cells by fluorescence microscopy and flow cytometry. Cytotoxicity of
drug-carrying and empty liposomes can be studied in tissue culture. Furthermore,
plasma clearance kinetics, tissue biodistribution, toxicity, and therapeutic efficacy of
drug-carrying liposomes can be assessed in appropriate animal models.
LIPOSOMES AS DRUG CARRIERS
The application of liposomes as a drug-delivery system has become more popular
over the last decades, because of their biocompatibility and versatility in carrying systemically administered drugs such as chemotherapeutics and antibiotics with narrow
therapeutic windows. A variety of therapeutic agents have been incorporated into
liposomes. Several have reached clinical use. These include liposomal doxorubicin
(15) (Doxil™), daunorubicin (16) (Daunoxome™), amphotericin B (17) (Amphotec™,
Ambisome™, and Abelcet™), cytarabine (18) (Depocyte™), and verteporfin (19)
(Visudyne™). Numerous liposomal formulations are in clinical trial, including those
for vincristine, all-trans retinoic acid, topotecan, and cationic liposome-based
therapeutic gene transfer vectors. Many more are in preclinical evaluation including
liposomal formulations of chemotherapeutics, neutron capture agents, oligonucleotides, plasmid DNA, photosensitizers, antibiotics, and vaccines (20). Besides
potential use in systemic gene delivery, cationic liposomes are routinely used as
transfection reagents for plasmid DNA and oligonucleotides in the laboratory.
Liposomal delivery of anticancer drugs has been shown to greatly extend
their systemic circulation time, reduce toxicity by lowering plasma free drug concentration, and facilitate preferential localization of drugs in solid tumors based on
increased endothelial permeability and reduced lymphatic drainage, or enhanced
permeability and retention (EPR) effect (21–23). For example, liposomal entrapment
of doxorubicin greatly reduces its dose-limiting cardiotoxicity. Clearance of drugs in
a liposomal formulation is mediated by phagocytic cells of the reticuloendothelial
system (RES), primarily located in liver and spleen. Liposomal entrapment can
protect drugs such as nucleic acids from rapid metabolism by plasma enzymes (24).
In addition, liposomal delivery of drugs appears to mediate bypassing of
P glycoprotein (Pgp)-related multidrug resistance in tumor cells. Liposomes also
present a platform for delivery of drug combinations. For example, Wang et al. (25)
coencapsulated doxorubicin and verapamil (a Pgp inhibitor) into liposomes and
studied their in vitro cytotoxicity. The result demonstrated effective reversal of multidrug resistance in doxorubicin-resistant cell lines.
FORMULATION STRATEGIES FOR LIPOSOMES
Lipid Composition for Increased Stability In Vivo
As drug permeability over the lipid bilayer is reduced in the “gel” state compared
to “fluid” state, stability of liposomal entrapment can be maximized by selecting
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high-phase-transition phospholipids, for example, phosphatidylcholines (PCs) with
long and saturated fatty acyl chains, such as distearoyl PC and hydrogenated soy
PC, which remain in a gel state at physiological temperature. Addition of 30 to
50 mol% of cholesterol can further improve stability of the lipid bilayers by filling in
gaps between PC molecules. Having a tight bilayer also reduces insertion of plasma
proteins and reduces RES clearance of the liposomes.
Sterically Stabilized Liposomes
RES clearance of liposomes is mediated by adsorption of plasma proteins to the
bilayer surface. Incorporation of 3 to 10 mol% of polyethylene glycol (PEG)conjugated lipid, such as monomethoxy-PEG (molecular weight 2000)–distearoyl
phosphatidylethanolamine (mPEG2000–DSPE) has been shown to greatly extend the
circulation time of liposomes by providing steric hindrance on the bilayer surface
(21,22). PEGylated liposomes exhibit circulation half-life of up to two days compared to several hours for non-PEGylated liposomes. The prolongation in mean
residence time of PEGylated liposomes is due to slower clearance of these liposomes by RES organs (26). This can, in addition, increase EPR effect-mediated tumor
localization and antitumor therapeutic efficacy.
pH-Sensitive Liposomes
Although high stability of liposomes prior to reaching the cellular target is generally
desirable, efficient release of liposomal drug in the target tissue and/or cell is essential for its therapeutic activity. Environmentally sensitive liposomes are designed to
take advantage of the differences in the microenvironment of a solid tumor or inside
the cell and to undergo destabilization on reaching their target. pH-sensitive liposomes are designed to destabilize at mildly acidic pH found in solid tumors and in
endosomal compartments. These are typically composed of dioleoyl phosphatidylethanolamine (DOPE), which has a cone-shaped geometry that favors transition
from bilayer to HII phase, and a weakly acidic amphiphile, such as oleic acid or
cholesteryl hemisuccinate, which stabilizes the bilayer structure at neutral pH but
not at mildly acidic pH (27–29). pH-sensitive liposomes have been shown to be
much more effective in facilitating endosomal release of membrane-impermeable
drugs from internalized liposomes in cells. Other nonbilayer-favoring lipids, such as
oleyl alcohol and diolein, are also effective in forming pH-sensitive liposomes
(30,31). Alternatively, a pH-sensitive oligopeptide that undergoes coil to α-helix
conformational transition, such as glutamic acid-alanine-leucine-alanine (GALA),
influenza fusion peptides, and pH-responsive polymer poly-2-ethyl-acrylic acid
linked to a lipid anchor, can also mediate intracellular disruption of the endosomal
membrane (32–34).
Fusogenic and Endosomolytic Liposomes
These liposomes can be constructed by reconstitution of envelope proteins of viruses
into liposomes or encapsulation of hemolysins from bacteria with varying degrees
of pH dependence (35). In addition to pH, the reducing and enzymatic environment
inside the endosomal compartment can be utilized to trigger the cleavage of disulfide or enzyme-sensitive linker that may be incorporated into a bilayer-stabilizing
lipid (36,37). For example, gelonin, a type I plant toxin, was coencapsulated inside
pH-sensitive liposomes with listeriolysin O (LLO), the pore-forming protein that
mediates escape of the intracellular pathogen listeria monocytogens from the
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endosome into the cytosol (38). Such a strategy resulted in a dramatic improvement
on the cytotoxicity of encapsulated gelonin against the murine B16 melanoma cell
line, over free gelonin or gelonin encapsulated in non-LLO-containing pH-sensitive
liposomes. In another study, Mastrobattista et al. (35) have demonstrated that
coencapsulation of a pH-dependent fusogenic peptide (diINF-7) and diphtheria
toxin A chain (DTA) in non-pH-sensitive immunoliposomes promotes efficient cytosolic delivery of DTA.
Temperature-Sensitive Liposomes
Local release of liposomal drug can also be triggered by hyperthermia by adopting
a bilayer composition with phase-transition temperature slightly above 37°C, such
as dipalmitoyl phosphatidylcholine or conjugation to a thermosensitive polymer
(39,40). Temperature-sensitive liposomes that show phase transition at 40°C can be
synthesized by incorporating lipid-conjugated copolymers of N-isopropylacrylamide and N-acryloylpyrrolidine (41). A thermosensitive liposome formulation
entrapping doxorubicin (ThermoDox) is currently undergoing clinical evaluation.
Cationic Liposomes
These liposomes can form electrostatic complexes with plasmid DNA and facilitate
gene transfer (42). A wide variety of cationic lipids have been synthesized including
those with a monovalent headgroup, such as 1,2-dioleoyl-3-trimethylammoniumpropane, N-[2,3-(dioleyloxy)propyl]-N,N,N-trimethylammonium chloride, and
3-β-[N-(N’,N’-dimethylaminoethyl)carbamoyl]-cholesterol, and those with a multivalent headgroup, such as 2,3-dioleyloxy-N-[2(spermine-carboxamido)ethyl]-N,Ndimethyl-1-propanaminium trifluoroacetate. DOPE is often used as a helper lipid
in these liposomes to enhance their fusogenicity (43). Cationic liposomes with
multivalent cationic lipids form particles with condensed structure with plasmid
DNA, whereas those with monovalent cationic lipids have been shown to form
extended spaghetti-like structures. Although cationic liposomes exhibit efficient
gene transfer activity in tissue culture and are currently commonly used reagents for
in vitro transfection, only low-level transfection in select tissues, typically the lung,
can be obtained upon systemic administration of cationic liposome/DNA complexes in murine models (44,45). This might be due to the trapping of the cationic
complexes in the capillary blood vessels in the lung, which is the first-pass organ
encountered by intravenously administered liposome complexes. In addition to
plasmid DNA, cationic liposomes have been used in the delivery of antisense oligodeoxyribonucleotides (ODNs) and siRNA into cells. Liposomes with a weakly basic
cationic lipid, such as 1,2-dioleyl-3-dimethylammonium-propane, can efficiently
incorporate plasmid DNA or ODNs at acidic pH, where the lipids are largely cationic and return to a mostly neutral zeta potential at pH 7.4, where the lipids
are mostly deprotonated (46). These liposomes have longer circulation time than
cationic liposomes carrying permanently charged cationic lipids and may be useful
for systemic delivery of DNA to solid tumors. Cationic liposome complexes with
plasmid DNA, ODNs, or siRNA can also activate the innate immune system and
may play a role in immunotherapy (47). In addition to nucleic acid delivery, cationic
liposomes carrying chemotherapy agent paclitaxel have been shown to preferentially target tumor endothelium, suggesting a possible role for these liposomes
in tumor-targeted drug delivery. Schmitt-Sody et al. (48) showed that cationic
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liposomal paclitaxel exhibits high selectivity for tumor endothelium and is highly
efficacious in tumor growth inhibition.
Targeted Liposomes
Liposomes can be targeted to specific cell populations via incorporation of a targeting moiety. The targeting ligand can consist of a lipid-anchored antibody or antibody
fragment, transferrin, folate, or carbohydrate. Immunoliposomes are synthesized
by conjugation of the liposome to an antibody [e.g., anti-HER2 (49), antitransferrin
receptor (50,51), anti-CD20 (52), and anti-CD19 (53)] or an antibody fragment such
as Fab (54) and scFv (55). HER2 is a receptor tyrosine kinase, a product of the HER2
(c-erbB2) proto-oncogene, which has been shown to play an important role in the
development and progression of breast and other cancers. Park et al. (56,57) reported
that anti-HER2 immunoliposomes, with encapsulated doxorubicin, displayed
significantly enhanced therapeutic efficacy in four different breast cancer xenograft
models when compared to nontargeted liposomes, free drug, or free antibody. For
targeting PEGylated liposomes, it is helpful to incorporate a long PEG-based linker
between the targeting ligand and the lipid anchor. Incorporation of the targeting
moiety can be accomplished during liposome formation by detergent dialysis, or
postliposome formation by conjugation to reactive lipids or postinsertion of ligands
from micelles of lipid-derivatized antibodies (58,59). The last method seems highly
promising for future clinical development. In addition to antibodies, other targeting
moieties such as transferrin (51,60) and folic acid (61–63) have been frequently
evaluated in targeting liposomes to tumor cells that overexpress the transferrin or
the folate receptor. Targeted liposomes are specifically taken up by target cells and
have been shown to be highly efficient in drug delivery and to bypass multidrug
resistance in cell lines (64). Tumor localization of targeted liposomes often does not
greatly exceed that of nontargeted liposomes because biodistribution of liposomes
is dictated by vascular permeability and the associated EPR effect. Furthermore,
there is concern that liposomes, due to their size, cannot penetrate into solid tumors,
which typically have high interstitial pressure. Nevertheless, targeted liposomes
such as anti-HER2 immunoliposomes and folate receptor-targeted liposomes have
shown improved antitumor efficacy in murine models over nontargeted control
liposomes (51,63). Leukemias, which have greater accessibility from circulation, are
also potentially good targets for targeted liposomes, as suggested by recent reports.
The advantage of using immunoliposome for MAb-based targeted therapy in leukemia exists in: (i) liposomes containing high payload of cytotoxic agents have unrestricted access to malignant cells and (ii) application of a chemotherapeutic agent
that has already shown clinical efficacy can potentially bring synergistic effect with
therapeutic MAb, based on a different killing mechanism. Pan et al. (65) studied the
therapeutic efficacy of folate receptor-targeted liposomes in combination with
upregulation of FR-β expression in an ascitic xenograft model of acute myelogenous
leukemia using all-trans retinoic acid. In vivo antitumor activity of folate receptortargeted liposomal daunorubicin in the L1210-JF ascitic murine leukemia model has
also been reported by Pan and Lee (66). The result showed that folate receptortargeted liposomes could effectively target receptor-positive leukemia cells in vivo.
Lipid Nanoparticles
Lipid nanoparticles are nanoscale spherical particles composed of lipids with a
lipidic core. These are suitable for delivery of lipophilic therapeutic agents. The
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molecule of interest can be formulated to lipid nanoparticle matrix through lipidphase dissolution. They have considerable utility as controlled delivery system for
drugs and vaccines. Lipid nanoparticles can be synthesized by combining an oil
phase (e.g., triolein) with phospholipids as emulsifiers. The oily core can be used to
incorporate lipid-soluble drugs such as paclitaxel (67,68), hematoporphyrin (69),
and lipid-conjugated prodrugs (70). They can be synthesized by similar methods as
those used in liposomes, such as high-pressure homogenization. Like liposomes,
these particles are cleared by RES, localized in tumors via EPR effect, and can be
made long circulating by incorporation of PEGylated lipid.
CONCLUSIONS
Lipid-based nanoparticulates are versatile drug carriers with significant potential for
clinical applications. Technological advances such as introduction of remote loading
methods, PEGylated liposomes, and targeted liposomes provided additional advantages. In addition to modulating toxicity, pharmacokinetics, and biodistribution,
liposomal delivery has shown promise as a mechanism to overcome multidrug resistance. Furthermore, liposomes are promising delivery vehicles for novel therapeutic
agents such as siRNA and drugs that lack aqueous solubility. Particularly promising
for future development are targeted liposomes, which have yet to be thoroughly
evaluated in clinical studies. Given current trends, lipid-based nanoparticulates are
likely to have an expanding role in drug delivery in the clinical setting.
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7
Nanoengineering of Drug Delivery Systems
Ashwath Jayagopal and V. Prasad Shastri
Biomaterials, Drug Delivery, and Tissue Engineering Laboratory, Department of
Biomedical Engineering, Vanderbilt University, Nashville, Tennessee, U.S.A.
INTRODUCTION
Challenges in drug delivery include the attainment of tunable release profiles, drug
solubility and structural stability, biocompatibility, and the confinement of therapeutic action to diseased sites. Increased attention has been directed toward the
nanoscale manipulation or nanoengineering of drug-delivery systems (DDSs) for
conferring unique properties on the drug or drug vehicle often not achievable by
either conventional- (e.g., free-form) or microscale-drug carriers. Nanoengineering
may be achieved at various levels during formulation of a drug within a carrier.
Examples include modulation of the drug environment at the molecular level, alteration of the physicochemical characteristics of the drug (e.g., solubility and structural integrity), control of drug-diffusive mobility, and modification of the bulk and
surface chemistry of the drug nanocarrier. These approaches have potential in the
achievement of favorable therapeutic endpoints, such as enhanced drug efficacy for
prolonged intervals, minimal drug toxicity, reduced dosage and cost burden, and
improved patient compliance. In this chapter, we discuss recent significant advances
in DDSs in which nanoengineering is the enabling technology for such objectives,
with an emphasis on the application of nanoengineering principles to the design of
controlled- and triggered-release DDSs, enhancement of drug activity and stability,
and smart tissue-targeted therapies. Specific attention is given to the unique effects
of nanoscale modifications on critical drug-delivery parameters. We evaluate the
promise of such approaches in drug-delivery technology, and discuss future considerations relevant to their clinical implementation.
CONTROLLED- AND TRIGGERED-RELEASE SYSTEMS
To develop suitable therapies for a diverse range of diseases, a host of natural and/
or synthetic biomaterials have been selected based on their physicochemical characteristics which enable the development of sustained- and triggered-release DDSs.
Sustained-release systems are administered for the maintenance of drug concentration within optimal therapeutic windows over extended periods ranging from hours
to months. In other cases, for instance where instant drug activity is required, DDSs
can be tuned for pulsatile release in response to varying physiological stimuli or
external triggers, such as acidic pH and acoustic pulses. Nanoscale modulation of
drug-transport characteristics, with or without entrapment within a nanocarrier,
can be achieved by formulation with novel drug-complexation agents for enhancement of both controlled- and triggered-release applications.
Polymeric Drug Delivery Systems for Sustained and Triggered Release
The engineering of drug carriers to bear one or more polymeric agents, with
each block having a specific role in the solubility and degradative susceptibility of
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the resulting vehicle, allows for nanoscale design of controlled-release DDSs.
Biodegradable polymer-based DDSs feature labile groups, the presence of which
can be tuned to control the extent of degradation and hence drug release over time.
The well-characterized and biocompatible poly(d,l-lactide-co-glycolide) (PLGA) drug
carriers are an example of a system which undergoes bulk erosion by hydrolysis,
enabling its utilization in controlled-release applications. The PLGA block copolymer is composed of poly(lactic acid) and poly(glycolic acid), the former of which is
slower to degrade due to increased crystallinity and steric hindrance to scission by
water. Both are safely eliminated by the body. Varying the relative composition of
the two blocks on a PLGA nanoparticle surface provides a means of tailoring drugrelease rates via modulation of the hydrophilic–lipophilic balance. PLGA can also
be incorporated into nanoparticulate DDSs for sustained intracellular drug delivery.
This is a beneficial consequence of two events: first, the nonspecific endocytic
internalization of PLGA nanoparticles into subcellular compartments and second,
the escape of PLGA nanoparticles into the cytoplasm made possible by surface
cationization within acidic endolysosomal compartments (1). These features suggest the versatility of PLGA nanoparticles as intracellular sustained-release depots
for gene or drug delivery, which bypass previous challenges such as nucleic acid
degradation in the cytoplasm, or drug efflux pumps which can be a barrier to
passive drug diffusion. A more recent DDS utilizing polyester-based hydrolytic
degradation has been reported by Ahmed and Discher (2). In this application, an
amphipathic polymersome composed of a water-soluble polyethylene glycol (PEG)
and hydrolytically labile polylactic acid or poly(caprolactone) block is synthesized
by self-assembly and can be utilized to encapsulate hydrophobic or hydrophilic
therapeutics, or imaging agents. Release is tunable from hours to weeks by adjustment of the hydrophilic–lipophilic balance, which controls hydrolytic-dependent
poration of the membrane. Although encapsulation efficiencies are similar to
well-characterized drug-delivery carriers such as liposomes, polymersomes exhibit
enhanced mechanical strength due to increased membrane thickness and extended
circulation half-life, suggesting the utility of this carrier for applications requiring
prolonged circulation in vivo with a resistance to destabilization mechanisms. In a
fundamentally different approach, therapeutics can be directly incorporated into a
polymer backbone (trade name PolymerDrug, Polymerix Corp.) (3). Adjustment of
the physical properties of the resulting polymer offers a pharmacokinetic tuning
mechanism, and biodegradation products are inert. This technology offers a platform
for the low-risk incorporation of therapeutics with known safety profiles that can be
implemented in implantable or degradable DDSs.
Polymeric DDSs have also been applied as ad hoc triggered-release systems.
One route to this goal is the blending of polymers to create polymer chain compositions which are more susceptible to rapid degradation. By this approach, sustainedrelease polymers can be modified for rapid release of encapsulated contents when
needed. For example, although PLGA is a well-characterized copolymer for biodegradable sustained-release systems, its intracellular release profile may not be
rapid or efficient enough for some gene-delivery applications. Little et al. addressed
this challenge by coblending PLGA with a pH-sensitive poly( β-amino ester) which
is insoluble at physiologic pH, with a solubility transition occurring within the
acidic pH range of lysosomes (4). Taking advantage of the rapid solubility of the
polymer below pH 6.5, poly( β-amino ester) blends with PLGA in microparticlecontaining genetic vaccines rapidly released their contents within dendritic cells (5),
and were also noted to have higher plasmid DNA-loading efficiencies (6). The same
Nanoengineering of Drug Delivery Systems
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polymer was also incorporated within poly(ethylene oxide) (PEO)-engineered
nanoparticles for the enhanced, pH-selective delivery of Taxol within tumors (7). In
another triggered-release strategy, micelles consisting of a hydrophilic PEO and a
hydrophobic polypropylene oxide (PPO) in a triblock copolymer formulation
(PEO–PPO–PEO) were observed to rapidly release doxorubicin and ruboxyl upon
nanoparticle perturbation induced by ultrasonic cavitation (8). In addition, upon
removal of the ultrasonic pulse, released drug was observed to become reencapsulated within micelles, and intracellular presence of 10% Pluronic micelles expedited
the sealing of ultrasound-induced cell membrane poration, resulting in minimal
drug efflux (9,10). Together, these data indicate that Pluronic micelles can tightly
bear their cargoes (in this case, due to PPO–hydrophobic interactions with drug)
with a spatially and temporally controlled acoustic-release mechanism. Taking
advantage of reencapsulation upon removal of acoustic pulses, therapeutic action
can be confined solely within the sonophoretic range, and due to the micelle-induced
membrane “sealing” effect, which increases intracellular drug retention, a lower
dosage may be used. These findings reinforce the concept that nanoscale delegation
of polymer block roles in drug release, solubility of the nanocarrier, and interactions
with the surrounding environment and external stimuli can be harnessed for diverse
and powerful drug-delivery approaches.
Lipid-Based Drug Delivery Systems for Sustained and
Triggered Release
Controlled-release DDSs have also been formulated using the unique properties
conferred by lipid-based materials. The fluidity of lipid domains on the nanoscale is
a parameter which, when tuned, allows for a mechanism controlling water influx.
Palm oil, a biocompatible vegetable oil, was used as a hydrophobic excipient along
with the phospholipid dipalmitoyl phosphatidylcholine for entrapment and controlled release of the highly hydrophilic drug terbutaline sulfate for pulmonary
administration (11). Hydrogenation of palm oil excipient was utilized to tightly
pack hydrophobic domains together for a tighter microsphere barrier. Degradation
of the spray-dried hydrophobic coat resulted in the sustained release of drug
nanoparticles without burst effects observed with free-drug nanoparticles. This
approach may serve as a potential high-compliance method of delivering drugs of
varying solubility through the pulmonary route featured by noninvasive circulatory access as spray-dried microspheres with near-optimal aerodynamic diameters,
as suggested by Edwards et al. (12) for improved deep-lung deposition with reduced
clearance by alveolar macrophages.
Using a judicious selection of components with appropriate physicochemical
characteristics, lipid-based DDSs can also be formulated for instant trigger-based
release applications. By altering the phase transition of the lipid chains from closely
packed arrangements (crystalline) to loose, disordered fluid chains, ideally at an
achievable, clinically relevant temperature, diffusion of drug from internal lipid
nanocarrier compartments can be achieved. Such phase transitions can occur within
sharp, predictable temperature ranges on the order of a few degrees celsius, and are
primarily a function of hydrocarbon length, extent of saturation, charge, and the
presence of other structures which affect lipid packing, such as cholesterol or double
bonds. In low-temperature thermosensitive liposomes (LTSLs), for example, the gel
phase rapidly transitions to the liquid crystalline phase at 42°C upon mixture
formulation of a conventional phospholipid combination with lysolecithin, at which
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temperature excessive efflux of drug is observed, with full release of doxorubicin
within 20 seconds (13). The rapid release of high quantities of drug from LTSLs has
important consequences on triggered chemotherapeutic applications. Peak intratumoral doxorubicin concentrations upon hyperthermia of LTSLs were measured to
be 30-fold higher than free-drug and twofold higher than previous thermosensitive
liposome designs. Manganese-guided loading of LTSLs with doxorubicin has been
carried out to achieve similar loading concentrations as nontemperature-sensitive
liposomes (14), suggesting that hyperthermia-based chemotherapy may be a viable
clinical option to conventional liposomes for delivering more toxic payloads to the
tumor space. Given the current clinical usage of liposomal formulations such as
Doxil, similar liposomal formulations triggered by hyperthermia may be readily
accepted if thermal dosimetry protocols can be standardized. Novel nanoengineering approaches involving the incorporation of transition-temperature-lowering
components as well as metallic drug-loading schemes to achieve high encapsulation efficiencies are helpful in facilitating the implementation of liposomes with
enhanced trigger functionality in the clinic.
Modulation of Drug Environment and Physicochemical Properties
for Sustained and Triggered Release
In many cases, nanoscale modulation of the actual drug structure is needed to
enhance its solubility or otherwise increase its efficiency of transport. In alternative
administration routes, such as the transdermal pathway, this type of modification is
gaining increased attention. The benefits of transdermal administration are several
fold: (i) first-pass liver metabolism associated with oral and systemic therapy is
avoided, (ii) there is generally increased patient compliance (e.g., wearing a patch
vs. repeated injections), and (iii) adverse reactions associated with conventional
administrations of drug may be avoided in this form. However, the skin has low
permeability to most therapeutics, especially high-molecular-weight compounds,
and thus is not by itself an efficient administration site for most therapies. Several
approaches have been reported to address this challenge, and include physical skin
barrier perturbation and/or the promotion of drug penetration using ultrasound
(sonophoresis), electroporation, drug molecule electrophoresis (iontophoresis), and
chemically assisted cotransport (15). The latter involves the nanoscale modulation
of drug-diffusive mobility by the transient induction of skin permeability, and/or
the improvement of drug portioning within the skin. Known permeation enhancers
include fatty acids, ethanol, and other skin-permeable solvents (15). Lee et al. (16,17)
reported the synergistic effect of n-methyl pyrrolidone and oleic acid in the enhancement of hydrophobic and hydrophilic drug flux (Fig. 1).
N-Methyl pyrrolidone is perfectly miscible with water for comixing in aqueous transdermal formulations, and is thought to exert its effects by promotion of
hydrogen bonding with the therapeutic, as well as the transient disruption of lipid
bilayers, enabling drug partitioning in the stratum corneum for enhanced flux
(16,17). Given enhancements in drug transport observed with physical perturbation methods, the combination of chemical permeation enhancers with ultrasound,
a widely available technology, is likely to have a significant impact on the transport
of a wide spectrum of drugs for sustained release.
The alteration of drug solubility is often critical to its therapeutic efficacy.
A technique of modulating drug solubility and thus its partitioning dynamics is
supramolecular complexation. In this approach, the therapeutic is combined with a
Nanoengineering of Drug Delivery Systems
103
FIGURE 1 Top: Flux enhancement of hydrophilic HCl salt drugs from 1:1 H2O/n-methyl pyrrolidone cosolvent through stripped human cadaver skin. Bottom: Transport of lidocaine and n-methyl
pyrrolidone from H2O/n-methyl pyrrolidone 1% oleic acid systems. Abbreviation: NMP, N-methyl
pyrrolidone. Source: From Ref. 17.
stabilizing agent to form an inclusion complex, with the overall structure then inheriting the physicochemical properties of the complexing agent. Cyclodextrins (CDs),
cyclic polysugars, have been known to perform this task very efficiently in oraldosage formulations, enabling powerful sustained-release applications which are
not achievable with the free-drug alone. For example, the poorly water-soluble antimicrobial agent chlorhexidine was complexed with β-CDs of varying lipophilicity
(18). The CD–drug complex was loaded into PLGA-implantable DDSs, creating a
concentration gradient. Enhanced solubility conferred by the complex allowed for
improved drug diffusivity, whereas the lipophilicity of the drug–CD inclusion
complex could be altered to modulate sustained-release kinetics. Furthermore, the
antimicrobial activity of the chlorhexidine was enhanced by complexation with CD,
through a plausible mechanism that involves enhancement of the interaction of the
drug with the glycocalyx (polysaccharide coat) on the bacteria (US Patent 6,699,505).
Thus, the direct alteration of drug physicochemical properties is a powerful approach
in extending the scope of drugs that can be administered by DDSs with high efficacy.
ENHANCEMENT OF DRUG STRUCTURAL STABILITY
AND DURATION OF ACTIVITY
Several diverse approaches have been reported for the preservation of drug structural
integrity and enhancement of therapeutic activity. Therapeutics such as proteins
and peptides are especially prone to in vivo degradation and clearance, which are
barriers to their bioavailability and duration of action. Other therapies may be poorly
soluble, or degrade within harsh nanocarrier environments over time in circulation
or in storage prior to administration. Approaches to address these challenges on the
nanoscale include the maintenance of controlled drug microenvironments within
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the vehicle, polymeric surface engineering to modulate drug interactions with the
surrounding environment, and the development processing techniques which
utilize suitable temperatures, solvents, and/or complexation agents which confer
stability upon labile drugs, thus enhancing their efficacy and applications.
Formulation Processing Strategies for Enhanced Drug
Stability and Activity
Many promising therapeutics are limited in clinical applications by poor solubility.
Such drugs also have low dissolution rates which hinder bioavailability even in
cases of rapid drug uptake. Previous approaches focused on increasing dissolution
rates by increasing the surface area of the drug powder using milling of the drug to
create microparticles. However, for low-solubility drugs, this is often not a sufficient
surface area enhancement to promote adequate dissolution. Nanoengineering of the
drug particle formulation itself has been performed to improve drug solubility and
dissolution. For example, a “nanonization” process utilizing high-pressure homogenization techniques was developed by Keck and Muller (19), which avoids
undesirable effects associated with solvent precipitation and pearl milling (trade
name: DissoCubes). The decrease in particle size and diffusion distance afforded by
DissoCubes provides for an enhanced drug dissolution rate and solubility. The latter
is strictly a consequence of the fact that the drug is less than 1 µm in diameter (above
which size, temperature and solvent characteristics are primary determinants of
particle solubility), and increased curvature of the drug particle increases its
dissolution pressure, which thus enhances solubility. Owing to the homogeneity of
the particles within the nanosuspension, which may include electrostatic and steric
stabilizing agents, potentially degradative processes such as Ostwald ripening are
not observed, providing drug stability in storage.
To attain suitable intracarrier environments for drug stability within DDSs
such as PLGA nanoparticles, antacid salts have been introduced during encapsulation to neutralize the acidic internal microclimate, reducing protein aggregation
and degradation (20). Saccharides such as trehalose have been used as a cryoprotectant in drugs encapsulated within solid–lipid nanoparticles (21). The CDs
previously discussed are known to protect labile drugs against hydrolysis, oxidation, and photodegradation (22). The utility of CD complexation for the protection
of a thermally labile drug, rhodium (II) citrate, was reported by Sinisterra et al. (23),
highlighting the potential of CD complexation to stabilize drugs from hightemperature compression and injection molding processes used to manufacture
polymeric DDSs (Fig. 2).
Hydrophilic CDs with hydrophobic cavities, such as HPβCD, can serve as
nanoscale shields which protect hydrophobic residues on proteins, for reduced denaturation and aggregation, and reduce unwanted polymer–drug interactions within the
carrier. These are examples of nanoscale-drug-stabilization strategies which can be
utilized to enhance solubility and structural conformation of drugs for incorporation
into nanoscale DDSs, or simply for improved transport parameters of the drug itself.
Polymeric Strategies for Prolongation of Drug Action
Polymeric surface functionalization of DDSs is now commonly used to modulate
the interaction of DDSs with the surrounding environment, for enhanced drug
activity through enhanced half-life in the body, and reduced biodegradation. PEG is
well known for its ability to diminish carrier clearance by the reticuloendothelial
Nanoengineering of Drug Delivery Systems
105
FIGURE 2 (A) Thermogravimetric curves of hydroxypropyl-β-cyclodextrin (HPβCD) (—), Rh(II)
citrate (-· -·), and association complex between Rh(II) citrate and HPβCD (- - -). (B) Differential
thermogravimetric (DTG) curves of HPβCD (—), Rh(II) citrate (-· -·), and association complex
between Rh(II) citrate and HPβCD (- - -). The thermogravimetric and DTG curves of the Rh(II)
citrate–HPβCD complex indicate only one thermal transition at higher temperature, around 310°C,
which was accompanied by an 80% mass loss. This data suggests that HPβCD complexation with
Rh(II) citrate results in enhanced thermal stability. Source: From Ref. 23.
system and inhibit protein surface adsorption. This is a consequence of the exclusion
volume exerted by the freely mobile polymer chain of the carrier surface. The biocompatible PEG imparts water solubility to the conjugated (PEGylated) therapeutic.
This strategy has been recently utilized for the protection of therapeutic proteins,
as reinforced by the increasing number of PEGylated drugs in the clinic, such as
PEGASYS (interferon alfa-2b, Roche) and Exubera (inhaled insulin, Pfizer). Strategies
for utilizing PEGylation for future protein-based therapies are dependent on the
development of PEGylation chemistries which confer the advantages of PEG without significant losses in protein functionality. This goal is being achieved with therapeutic monoclonal antibodies. Random PEGylation of protein amine groups using
PEG coupled with NHS esters was implicated in the loss of monoclonal antibodybinding affinities in several cases. However, attachment of PEG to Fab′ fragments
using cysteine cross-linking in the hinge region (i.e., through thiol-maleimide
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Jayagopal and Shastri
chemistry) reduced or eliminated most affinity problems while conferring enhanced
circulation half-life (24). To expand this idea, conjugation of branched and linear
bifunctional PEGs to single-chain antibodies (scFvs) with an engineered unpaired
cysteine to create multivalent conjugates was also carried out with success (25,26).
Although the problem is more difficult for proteins other than antibodies which
have the unique benefit of unpaired cysteines not needed for biological activity, sitespecific PEGylation has also been reported using enzymatic or site-directed mutagenetic approaches (27,28). The technology for site-specific PEGylate therapeutic
proteins is an area of active research and that is likely to expand the library of
available protein-based therapies.
TARGETED THERAPIES
Several approaches for the specific targeting of therapeutics to disease sites have
been developed, for the purposes of maximizing the effective dose delivered to the
site while lowering the total dose needed, and sparing healthy tissue from potential
adverse drug effects via confinement of therapeutic activity. Generally, these
strategies involve optimization of geometry of the nanocarrier to optimize tissue
uptake of DDSs, a generally passive route, and the bioconjugation of ligands such
as peptides to the nanocarrier surface to promote interactions with diseased tissue,
an active targeting mechanism.
Strategies for Passive Tissue Targeting
The geometry and mechanical properties of a carrier has a profound influence on its
therapeutic efficacy in that the ability for DDSs to target tissue can be dependent on
them. For example, spherical and inflexible carriers have been suggested to have
reduced transport in tumor interstitial models compared to flexible and extended
structures, and due to tumor vessel pore size restrictions, extravasation of nanospheres
in cancerous tissue may be a highly size-dependent event (29). Gastrointestinal
mucosa has size-dependent uptake phenomena affecting oral dosage of biodegradable nanoparticles (30,31). It follows that the physical characteristics of DDSs must be
optimized for the intrinsic properties of the specific targeting site. The leaky vasculature
of tumors has been exploited by the well-known enhanced permeability and retention
effect, whereby long-circulating PEGylated “stealth” liposomes are capable of accumulating in the tumor space due to a lack of rapid clearance conferred by PEG, combined with the accessibility of the tumor pores to passive liposomal uptake.
Strategies for Actively Targeted Therapies
The bioconjugation of ligands, such as monoclonal antibodies, proteins, or peptides
to the nanocarrier surface, has been exploited on many nanoparticulate DDSs for the
purpose of concentrating therapeutic action to specific sites. Nonspecific peptidebased internalization systems, such as those based on cell-penetrating peptides,
have been shown to be efficient in preliminary in vitro studies, and may internalize
in an energy-independent mechanism depending on the cargo involved (32,33), but
their use for in vivo targeting may not be clinically relevant due to the ability of these
peptides to nonspecifically target many cell types irrespective of pathology. Cancerous
cells have been targeted by nanoparticles toward unique surface antigens inherent to
the tumor type (e.g., HER2 in breast cancer) (34) or radiation area (35). Other targets
exploited by nanoparticulate systems with promising therapeutic outcomes include
Nanoengineering of Drug Delivery Systems
107
folate receptors (36) and cell-adhesion molecules (37–39), both prominent examples
of ligand-mediated internalization processes that are clinically relevant in a number
of pathologies.
Strategies for Drug Delivery System Transport Across Tight
Endothelial Junctions
Tight endothelial barriers, such as the blood–brain barrier (BBB) and blood–retinal
barrier, pose significant challenges to the transport of therapies, most of which
traverse such barriers only under extreme pathological conditions. Methods to
traverse these tight junctions would enable the clinical usage of multiple therapies
and imaging modalities using contrast agents. Nanoparticulate delivery across the
BBB for the transport has been achieved with numerous approaches. Transport of
ligand-coated PEGylated polylactide–PLGA nanoparticles across the BBB has been
demonstrated (40). A novel strategy was recently reported for PLGA nanoparticle
transport across the BBB with the surface engineering of peptides similar to synthetic
opioids (41). Our laboratory is currently investigating the potential of functionalized
solid–lipid nanoparticles to transport magnetic resonance contrast agents, proteins,
and fluorescent probes across the BBB. Thus, one can retain the desirable properties
of nanocarriers with the added feature of BBB penetration, for the delivery of therapeutic or diagnostic agents which are not lipid-soluble enough or do not meet the
physical criteria to normally cross the BBB.
FUTURE DIRECTIONS
The potential of nanoengineering strategies in the development of controlled- and
triggered-release DDSs, preservation of drug activity, and disease-specific targeting
has been presented. Many of these studies are preliminary in nature, and various
systems require further toxicity and efficacy data to facilitate their transition to clinical trials. However, the synthetic and natural basic units which form the basis of
these strategies, such as the therapeutics, PLGA, and solid lipids, have entered these
phases. The success of these initial nanodelivery systems will usher in the clinical
application of combinatorial, integrative strategies which constitute “smart” delivery
systems, such as the multimodal “nanocell,” a PLGA, PEG, and phospholipid-based
nanoparticle which delivers both antiangiogenic and chemotherapeutic agents with
temporally controlled kinetics (42). Other considerations, such as the ability to
scale-up laboratory techniques for mass production of DDSs and the ability to
upgrade existing, approved systems with stepwise functional additions, will also
expedite their clinical usage. Nanoengineering offers the unique ability to manipulate the smallest interactions at the most fundamental scales, which is certain to
provide a series of major advancements in drug delivery technology.
ACKNOWLEDGMENTS
This chapter was made possible by generous funding from the Vanderbilt University
Institute for Integrative Biosystems Research and Education (VIIBRE), and the
Vanderbilt Vision Research Center Training Grant (NEI T32 EYO7135).
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8
Aerosol Flow Reactor Method for the
Synthesis of Multicomponent Drug
Nano- and Microparticles
Janne Raula
NanoMaterials Group, Laboratory of Physics and Center for New
Materials, Helsinki University of Technology, Helsinki, Finland
Hannele Eerikäinen
Pharmaceutical Product Development, Orion Corporation
Orion Pharma, Espoo, Finland
Anna Lähde
NanoMaterials Group, Laboratory of Physics and Center for New Materials,
Helsinki University of Technology, Helsinki, Finland
Esko I. Kauppinen
NanoMaterials Group, Laboratory of Physics and Center for New
Materials, Helsinki University of Technology, and VTT Biotechnology,
Helsinki, Finland
INTRODUCTION
Nano- and microparticle drug carriers have potential applications for administration of therapeutic molecules (1,2). Targeted drug delivery can be achieved by small
particles due to their tendency to accumulate in targeted areas of the body. Moreover,
the solubility of material from the nanoscale objects is notably enhanced due to the
increased surface-to-volume ratio of small particles. As well, the systemic side effects
in drug targeting, for example, into a cancerous tumor, can be minimized by the
decrease a particle size (3). Sub-micron, solid-state drug particles that tend to be
unstable for many drug molecules can be stabilized by specific polymers. In these
composite particles, the polymer may act, in addition to stabilizer, as functional material that controls the release and diffusion of a drug depending on the environmental
conditions such as pH, temperature, ionic strength, humidity, and so on (4,5).
Several methods have been studied for the preparation of the drug nanoand microparticles. A common liquid route method to prepare the drug nano- and
microparticles is the use of an emulsion. Probably, the most used method is an oilin-water emulsion consisting of two immiscible solvents such as chloroform and
water (6). As the drug particles show a large tendency towards agglomeration and
growth, surfactants have to be added to stabilize the droplets and particles. Other
related methods include salting-out (7), nanoprecipitation (interfacial precipitation)
(8), phase separation (9), and evaporative precipitation into aqueous solution (10).
In general, the production of particles including two or more drug molecules is
difficult.
Size-reduction techniques, such as wet milling and high-pressure homogenization, have also been used to prepare nano- and microparticles (11). High-shear
111
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Raula et al.
forces and thus high energies used in milling processes can create uncontrollable
changes in the product, such as chemical degradation, changes in surface energetics,
and damage in crystal structure. Furthermore, long processing times increase the
risk of microbiological contamination. Surfactants are needed to reduce particle
agglomeration and sintering during wet milling. Multicomponent particles with
controlled morphology and surface characteristics (composition and morphology)
cannot be produced.
Small particles have also been prepared using supercritical fluids either
as solvents or antisolvents for the drug and the polymer (12). The preparation of
uniform multicomponent particles consisting of a drug and a polymer has been
shown to be difficult due to different crystallization and precipitation kinetics of
the drug and the polymer molecules and due to partitioning of the drug into supercritical fluid.
Spray-drying has been widely used for the production of micrometer-sized
particles (13). Spray-drying involves the conversion of a solution droplet into a dry
particle by evaporation of the solvent in a one-step process. Both water-soluble and
-insoluble compounds can be prepared. Thus, the recovery of the drug in the particles is almost quantitative. Also temperature-labile compounds such as proteins
and enzymes have been successfully spray-dried. The solvent properties and the
spray-drying variables can control the particle properties, especially morphology.
Multicomponent particles with spherical morphology can be produced. However,
particles are typically amorphous, and surface structure as well composition cannot
be controlled in detail.
This chapter presents the novel aerosol flow reactor method for the synthesis
of single- and multicomponent nanoparticles as well as nanostructured, micronsized particles (14–21). We demonstrate the production of powders made from
different drug and/or excipient materials. Besides the drug molecules, the polymeric drug nanoparticles contained methacrylic polymers, which are pharmaceutically acceptable (22,23). It was observed that the solubility of a drug within the
polymer matrix depends not only on the amount of the drug, but also on the
polymer itself. The drug dissolution form of the nanoparticles, interactions between
the drug and the polymer, and drug crystallinity within the polymer particles will
be discussed. This chapter also discusses the formation of the polymeric nanoparticles from several solvents. The studied subjects were polymer solubility in a
solvent medium and the influence of solvent vapor pressure.
Chemical and physical stability of dry powders in storage is crucial. Amorphous
materials tend to crystallize in time, thus building bridges between individual
particles. This in turn affects the powder flowability and handling. A part of this
work presents the attempts to crystallize the drug within the aerosol reactor in
different conditions. This chapter discusses the reactor conditions affecting the drug
crystallization.
Adding leucine derivatives to drug particles has been shown to reduce the
adhesion between particles (24,25). Furthermore, a α-amino acid l-leucine is a surface-active compound in aqueous solutions. It is also important that l-leucine is
generally regarded as safe material for the human body. This chapter discusses the
preparation of the composite powders using l-leucine to increase the stability of
powders. Main materials were sodium chloride, which is a representative inorganic
material that crystallizes with ease, and lactose that forms amorphous particles. The
incorporation of l-leucine modifies particle surface structure and changes, for
instance, spherical particles to wrinkled ones depending on the l-leucine content.
113
Aerosol Flow Reactor Method
Moreover, l-leucine has been shown to reinforce the structure of particles where
material is in a rubbery state.
AEROSOL FLOW REACTOR METHOD
In the aerosol flow reactor method, the solution containing solute(s) is atomized to
produce droplets that are transferred with the aid of a carrier gas to a heated flow
reactor (part I in Fig. 1). The inert carrier gas is either dry or saturated with a solvent.
In the latter, the droplets remain wet in the reactor until the aerosol is heated (part II
III. POLYMERIC DRUG
NANOPARTICLES
A) and C) matrix-type
B) and D) core-shell
dru g mole cule s
dru g crystals
polymer chains
COLLECTION
T(ambient)
HEATING ZONE
HEATING ZONE
DILUTION
II. PARTICLE FORMATION
- drying
- self-assembling
(core-shell)
- molecular arrangement
(crystal growth)
- morphology transformation
Temperature
increases
Dry or solvent saturated
carrier gas mixed with droplets
Starting solution:
drug, polymer,
and/or other
excipients
dissolved in
solvent
I. DROPLET GENERATION
- jet nebulizer
- ultrasonic nebulizer
- electrospray
FIGURE 1 Schematical presentation of the novel aerosol flow reactor method for the production
of nano and micronsized drug powders.
114
Raula et al.
in Fig. 1). The fast drying promotes the formation of amorphous particles. Allowing
time for droplet drying provides also time for molecular arrangements such as crystal growth to occur. Varying flow rate and the residence time, drying rate of particles
and accordingly particle morphology can be manipulated in the process. Downstream
from the heated section, the dry aerosol is diluted with dry inert gas to avoid solvent
condensation onto dry particles. Polymeric drug nanoparticles of several types, that
is, matrix and core shell as shown in part III in Figure 1 have been produced.
Moreover, this aerosol method produces dry partocles directly without a need for
further purification, and no additives are required. In the following, we describe the
production of nano- and micron-sized dry particles containing the drugs, polymers,
and some model materials.
EXPERIMENTAL
Materials
Drug materials beclomethasone dipropionate (BDP; Sicor S.p.A., Italy), ketoprofen
(2-(3-benzoylphenyl) propionic acid) (Sigma, U.S.A.), naproxen [(S)-2-(6-methoxy-2naphthyl) propionic acid] (Sigma, U.S.A.), and acetyl salicylic acid (ASA;
Sigma-Aldrich, Germany) were used as received. Reference materials lactose monohydrate (provided by Orion Pharma, Finland), sodium chloride, NaCl ( J.T. Baker,
Holland), and excipient material l-leucine (Fluka, Switzerland) were used as received.
Pharmaceutically accepted methacrylic polymers Eudragit® L100 (Röhm Pharma,
Germany), Eudragit® E100 (Röhm Pharma, Germany), and Eudragit® RS (Röhm
Pharma, Germany) were used as received. Solvents ethanol (99.6%, Alko Oyj, Finland),
tetrahydrofuran (THF; J.T. Bakers, U.S.A.), and toluene ( J.T. Bakers, U.S.A.) were used
as received. Water was purified by ion-exchange (Millipore), and was measured to
have pH 6.
Precursor Solutions
For pure polymer solutions, the polymer was dissolved in either THF or ethanol. In
a case of solvent mixtures, water or THF was added into polymer solution while
stirring the mixture. The volume ratios of the solvents were 0.1, 1.0, and 9.0. The
total concentration of the polymer varied between 0.2 and 1.5 g/L. The equilibrium
concentrations for Eudragit® L100 in toluene (0.058 g/L) and in water (0.029 g/L)
were determined after filtration and solvent evaporation.
The drug–polymer solutions were prepared by separately dissolving the polymer and the drug in ethanol using a magnetic stirrer and then mixing the solutions
in respective amounts. The total concentration of a drug and an excipient varied
between 0.25 and 2.0 g/L.
BDP solutions 5 and 25 g/L were prepared by dissolving the drug in ethanol
at room temperature. A saturated BDP solution was prepared by stirring the
solution for 1 hour until no further dissolution of BDP was observed.
Aqueous l-leucine, lactose, and NaCl precursor solutions were prepared by
separately dissolving in water and stirring, and then the solutions were combined
in respective amounts. ASA was dissolved in ethanol, and the solution was combined with the aqueous l-leucine solution. The total solution concentration varied
from 0.25 to 30.0 g/L.
Particle Production by the Aerosol Flow Reactor Method
The set-up of the aerosol flow reactor is presented in Figure 2. All the experimental
conditions for the preparation of nano- and microparticles are given in Table 1. The
115
Aerosol Flow Reactor Method
Vac. 25 L/min
Charger
Exhaust
ELPI
impactor
p
Vac. 9.6 L/min
BLPI
impactor
Computer
Vac. 0.3 L/min
N2, 3.0 L/min
Electrostatic
SEM/TEM sampler
Optional
heater
Dilution, N2
T
p
Porous
tube
aerosol
diluter
Preimpactor
DMA
Kr-85
p
Filter
CPC
Heater
Thermal insulation
and/or heating tape
Critical orifice
p
Pressure meter
T
Temperature meter
Vac. 3.0 L/min
Vac. 0.3 L/min
Computer
Reactor tube:
Stainless steel
Inner diameter 30 mm
Heated length 800 mm
Valve
Pyrosol
Collison
N2
p
N2
p
Fresh
Excess
starting solution
solution
Solution Precursor
feeding
solution
container
FIGURE 2 Experimental set-up used in the preparation of nano and microparticles. N2 = clean, dry
pressurized nitrogen. Abbreviations: Vac., vacuum; L /min, liters per minute; Kr-85, aerosol neutralizer
using 85Kr β-source; BLPI, berner-type low pressure impactor; CPC, condensation particle counter;
DMA, differential mobility analyzer; ELPI, electrical low-pressure impactor; SEM, scanning electron
microscope; TEM, transmission electron microscope.
solutions to produce nanoparticles were atomized using a Collison-type air jet
atomizer (TSI 3076, TSI, Inc., Particle Instruments, St. Paul, U.S.A.). The atomizer
produces a lognormal droplet size distribution with a geometric number mean
droplet size of approximately 300 nm and a geometric standard deviation between
1.6 and 2.0. The microparticle droplets were generated with an ultrasonic nebulizer
(RBI Pyrosol 7901, Meylan, France). The droplets were carried in dry nitrogen gas
into a heated reactor where aerosol flow was laminar. Depending on the flow rate,
the residence time for droplet/particles in the reactor can be modified. The temperature of the stainless steel reactor was controlled with four separate heaters.
Downstream from the heated section of the reactor, the aerosol was diluted by dry
nitrogen in a porous diluter to avoid solvent condensation onto particle surface as
well as particle deposition to reactor walls via thermophoresis. After fully mixing
the reactor and dilution gas flows, a sample of particles was collected by
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Raula et al.
TABLE 1 Summary of the Conditions for the Production of NanO- and Micronsized
Dry Powders
Polymeric drug
nanoparticles
Type of powder
Droplet generation
Materials
Solvents
Precursor solution
concentration (g/L)
Carrier gas flow rate (L/min)
Precursor solution
consumption (mL/min)
Reactor wall temperature (°C)
Residence time in reactor (s)
Flow rate in dilution (L/min)
Dilution gas temperature (°C)
{
BDP
Composite L-leucine
microparticles
microparticles
Collisona
BDP, ketoprofen,
naproxen
Eudragit® L100, E100 and
RS100
Water, ethanol, THF,
toluene
0.2–1.5
Pyrosolb
BDP
Pyrosolb
NaCl, lactose, acetyl
salicylic acid,
L-leucine
Ethanol
Water
5–50
1.0–27.5
1.5–3.5
0.2–0.4
1.5–2.0
0.09–1.6
1.4
0.28–0.45
40–200
6–21
25
20
50–150
16–21
10–75
50–100
20–100
19–24
76
20
aCollison-type
air jet atomizer.
nebulizer.
Abbreviations: BDP, beclomethasone dipropionate; THF, tetrahydrofuran.
bUltrasonic
a point-to-plane electrostatic precipitator (InTox Products, Albuquerque, U.S.A.)
onto either a plain or carbon-coated copper transmission electron microscope (TEM)
grid (Agar Scientific Ltd., Essex, U.K.). Large-quantity collections were conducted
by a Berner-type low-pressure impactor (BLPI) (26) or a small-scale cyclone (27).
Particle size and size distribution were measured directly from the gas phase downstream from the dilution and mixing processes. Computer fluid dynamics calculations of the reactor tube were performed for different flow rates ≤ 3.5 L/min with
temperatures covering the temperature range used in the experiments. The calculations showed that in the heated zone a fully developed laminar flow was achieved,
and the wall temperature was reached for all the particles (28).
Instrumentation and Characterization
The particle size (geometric number mean diameter, GNMD) and polydispersity
(geometric standard deviation, GSD) of the nanoparticles in the gas phase, that is, at
the reactor outlet, were determined with a TSI scanning mobility particle sizer
equipped with a differential mobility analyzer (DMA, model 3081) (TSI, Inc., Particle
Instruments, St. Paul, U.S.A.) and a condensation particle counter (CPC, model
3027) (TSI, Inc., Particle Instruments, St. Paul, U.S.A.). The average values of GNMD
and GSD were determined from three to six measurements. The maximum standard
errors were 5% and 0.5% for GNMD and GSD, respectively.
The GNMD and GSD of the produced microparticles were determined with
an electrical low-pressure impactor (ELPI; Dekati Ltd., Tampere, Finland) (29,30).
Oiled porous collection stages (Dekati Ltd., Tampere, Finland) were used to avoid
the bouncing of the particles from upper stages to lower ones. The GNMD and GSD
of the particles were calculated using equations GNMD = exp(Σ (ni ln Di)/N) and
GSD = exp((Σni (ln Di – ln GNMD)2)/(N – 1))1/2, respectively. Here ni is the number of
Aerosol Flow Reactor Method
117
particles in the ith group, Di the aerodynamic diameter of the ith group, and N the
total number of particles, that is, Σni.
Mass median aerodynamic diameter (MMAD) and GSD of the dispersed particles were determined with the BLPI and using equations MMAD = exp(Σ (mi ln
Di)/M and GSD = exp((Σ (mi D3i (lnDi – lnMMAD)2))/(Σ (miD3i) –1))1/2, respectively. Here mi is the mass fraction of the particles on the collection stage and M the
sum of mass fractions and is unity. In all the experiments, the density of the particle
is assumed to be 1 g/cm3.
The morphology of the particles was analyzed with a field emission scanning
electron microscope (SEM; Leo DSM982 Gemini, LEO Electron Microscopy, Inc.,
Oberkochen, Germany). Internal structure of the particles was studied using a TEM
(TEM; Philips CM200 FEG, FEI Company, Eindhoven, The Netherlands). Some of
the samples were coated with platinum sputtering in order to stabilize the particle
under electron beam and to enhance image contrast.
The crystallinity of nanoparticles was studied using X-ray diffraction (XRD;
Philips PW 1710, Eindhoven and Almelo, The Netherlands) with Cu Kα radiation.
Diffraction angles (2θ) (goniometer PW 1820) used in the recording of the XRD
patterns were 3° to 40°. The crystallinity of individual microparticles was investigated with TEM as well. The thermal properties of the nanoparticles were studied
using a differential scanning calorimeter (DSC; Mettler Toledo DSC 822e, Mettler
Toledo AG, Greifensee, Switzerland) where the samples were heated from 25°C to
300°C or from −50°C to 200°C using a heating rate of 10°C/min. A nitrogen purge
of 50 mL/min was used in the oven.
RESULTS AND DISCUSSION
Polymer Nanoparticles
This section discusses the control of the polymer Eudragit® L100 nanoparticle
morphology when produced from different solvent media (21). The morphology of
nanoparticles depends on both the droplet drying rate and the polymer solubility in
solvent medium. Vapor pressure of the solvent is essential because it mainly determines the solvent evaporation rate. Vapor pressures for THF, ethanol, toluene, and
water at 25°C are 21.6, 7.87, 3.79, and 3.17 kPa, respectively. Also, a concept of solvent
quality for a polymer is noteworthy (31). The solubility of Eudragit® L100 in used
solvents was experimentally observed, and found to decrease in the following order:
ethanol > THF > toluene > water. The first two solvents are good solvents for Eudragit®
L100. The size of the dry Eudragit® L100 particles from ethanol and THF solutions
increased, respectively, from 75 to 130 and 65 to 95 nm within the concentration range
from 0.2 to 1.5 g/L. Accordingly, the GSD of the particles, that is, the spread of the size
distribution of the particles from THF increased from 1.9 to 2.2 but that from ethanol
remained the same, ~1.9. With added water, the solvent quality for the polymer worsens, that is, the polymer coil shrinks in the solution prior to atomization. As a result, the
size of dry nanoparticles prepared from ethanol/water = 1:1 solution decreased
approximately to two-thirds of the particles from pure ethanol. Moreover, the sizes
from THF/water = 1:1 solutions at 0.2 and 1.0 g/L decreased, respectively, to fourfifths and two-thirds of the particles from pure THF. Polymer particles from toluene
and water were small with the particle size around 35 nm and the GSD less than 1.9.
Figure 3 shows exemplary SEM images of the particles prepared from different
solvent media. The figure summarizes the factors influencing the particle formation. Three main routes are discussed: (A) Non-hollow solid particles can be obtained
118
Raula et al.
drug
and/or
polymer
dissolved
in droplet
Toluene, sat/filt
g
C
Cp*
a
A
Ethanol, 0.2 g/L
B
100 nm
b
200 nm
Ethanol, 1.0 g/L
THF:H2O=1:9, 0.2 g/L
f
d
e
200 nm
c
THF 1.0 g/L
THF, 0.5 g/L
THF:H2O=9:1, 0.2 g/L
500 nm
200 nm
500 nm
500 nm
FIGURE 3 Scheme describes the influence of the experimental parameters on the particle formation
and resulting particle morphology. A droplet containing drug and/or polymer dries in the reactor
mainly in three ways: A, solidifying, B, film formation, and C, early precipitation. Along the solid arrow
is the solvent vapor pressure, solvent evaporation rate , and polymer concentration increase. Along the
dashed arrow is the solvent quality for the polymer (Eudragit® L100) decreases.
when solvent(s) evaporates slowly. The particles are often spherical. (B) If solvent
evaporation is fast, the solute(s) forms a crust or film on the droplet surface. If the
crust is impermeable to solvent, the particle will form a hollow interior due to pressure build-up on further solvent evaporation, which expands the particle (32–36).
Hollow spherical particles can also collapse to raisin- or lens-shaped particles (SEM
images c and d in Fig. 3) (34,37–39). (C) Polymer being close to its solubility limit in
solvent prior to atomization precipitates at very early stages of solvent evaporation.
For clarification, the solvent evaporation rate and polymer concentration increase
along a curved solid arrow. When following the curved dashed arrow, the solvent
quality for the polymer worsens. At some point of droplet drying, the polymer
reaches its solubility limit (Cp*) and precipitates (image g in Fig. 3). These particles
have often very irregular shape. In between these two extremes, the particles may
have very unusual morphologies, such as blistered (image f in Fig. 3) and wrinkled
(image e in Fig. 3) structures (21).
Polymeric Drug Nanoparticles
This section discusses the production of polymeric drug nanoparticles where the
drug and the polymer form a solid solution (18,19). The morphological stability of
the nanoparticles, that is, whether the state of matter of the drug will change during
storage, depends on the polymer used and will be discussed. Table 2 lists the experimental conditions as well as the physical properties of some of the polymeric drug
nanoparticles. The solvent was ethanol in every solution. Nanoparticles prepared
119
Aerosol Flow Reactor Method
TABLE 2 Summary of the Conditions and Physical Characteristics of Some of the Drug
Nanoparticles Produced
Drug material
Eudragit®
Ketoprofen 0%
Ketoprofen 10%
Ketoprofen 25%
Ketoprofen 33%
Ketoprofen 50%
Ketoprofen 67%
Ketoprofen 33%
Ketoprofen 33%
Naproxen 50%
Naproxen 50%
Naproxen 50%
Naproxen 50%
Naproxen 50%
L100 100%
L100 90%
L100 75%
L100 67%
L100 50%
L100 33%
E100 67%
RS 67%
L100 50%
L100 50%
L100 50%
L100 50%
L100 50%
Cdrug (g/L)
2
2
2
2
2
2
2
2
1
2.5
5
10
25
Treactor (°C)
GNMD (nm)
GSD
80
80
80
80
80
80
80
80
80
80
80
80
80
120
120
116
112
108
107
106
96
94
124
136
160
201
1,8
1,8
1,8
1,9
1,9
1,9
1,7
1,7
1,8
1,8
1,8
1,7
1,6
Drug concentration in precursor solution (CDRUG) and reactor wall temperature (TREACTOR) when producing drugpolymer composite nanoparticles are shown. Also geometric number mean diameter and geometric standard
deviation of the particle size distribution as determined with the differential electrical mobility method are shown.
Abbreviations: GNMD, geometric number mean diameter; GSD, geometric standard deviation.
with various proportions of ketoprofen and Eudragit® L100 at the reactor wall
temperature of 80°C showed a decreasing particle size as a function of the amount
of ketoprofen. With the same amount, 33% (w/w) of ketoprofen but a different
polymer, the particle size decreased in order Eudragit® L100 > Eudragit® E100 >
Eudragit® RS. The effect of solution concentration on particle size is also shown
in Table 2.
The amorphous form of the drug has a tendency to spontaneously convert to
a crystalline form (40,41). The structural stabilization, however, can be achieved by
incorporating the drug into the polymer matrix. Polymer-drug nanoparticles were
prepared from the drug materials BDP (18), naproxen (19), and ketoprofen (19). For
naproxen and ketoprofen, it was observed that there is a limit to how much drug
can be incorporated in an amorphous form into nanoparticles. When the amount of
the drug in the nanoparticles was for ketoprofen ≤ 33% (w/w) and for naproxen ≤
10%, the nanospheres were amorphous, proved by XRD, DSC, and TEM. No crystallinity or grain boundaries were found; instead, the nanoparticles were smooth with
a uniform interior. The formation of an amorphous, polymer–drug structure has
previously been observed for spray-dried particles containing methacrylic polymers and ketoprofen drug (42,43). When the amount of the drug in the nanoparticles
was increased to 50% (w/w) ketoprofen and 25% naproxen, an endothermic transition corresponding to melting of the drug crystals was detected with DSC (Fig. 4).
Moreover, in the XRD analyses, the broad background diffraction pattern of the
amorphous structure became overlapped by peaks corresponding to the diffraction
from the drug crystal lattice. However, the drug molecules on the particle surface
crystallized, thus forming large crystalline bridges between neighboring nanoparticles,
that is, phase-transition-induced sintering of the nanoparticles occurred.
Molecular size was assumed to affect drug crystallization during particle formation. BDP (Mw 521 g/mol) is a larger molecule than ketoprofen (254 g/mol) or
naproxen (230 g/mol) molecules. The molecular mobility of BDP was expected to be
slow due to its large size and high glass transition temperature (Tg ∼ 66°C) (44,45).
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Raula et al.
FIGURE 4 DSC thermograms of the Eudragit® L100 nanoparticles with drugs beclomethasone
dipropionate (BDP), ketoprofen, and naproxen (marked in the figures). The content of BDP (a) 100%
(w/w), (b) 80%, (c) 60%, (d) 50%, (e) 40%, (f) 20%, and (g) 0%. The content of ketoprofen (a) 100%,
(b) 50%, (c) 33%, (d) 25%, (e) 10%, (f) 0%. The content of naproxen (a) 100%, (b) 50%, (c) 33%,
(d) 25%, (e) 10%, (f) 0%. Abbreviations: cr, crystallization; m, melting; gt, glass transition.
The glass transition temperature, Tg, was calculated using the relation Tg ~ 0.7 Tm
where temperatures were in Kelvin (46,47), a melting temperature, Tm, measured by
DSC (18). Thus, the amorphous state of pure BDP nanoparticles, that is, no added
polymer, is kinetically preserved even though it is not thermodynamically stable.
The nanoparticles containing BDP both with and without polymer showed single,
separate nanoparticles. Crystallization of BDP in the nanoparticles could be induced
by heating, which can be seen as an exotherm in DSC measurements (Fig. 4).
In the amorphous structure, the drug is solubilized by the polymer, and the two
components form an amorphous solid solution (48). The interactions between the
drug and the polymer molecules determine the solubility of these materials in each
other (49). When the drug–polymer molecular interactions are comparable to the
drug–drug and the polymer–polymer interactions, the drug is well solubilized by
the polymer and large amounts of drug can be incorporated in the polymer matrix
without drug crystallization (50). However, when the drug–polymer interactions
are weaker than the drug–drug or the polymer–polymer interactions, the drug and
the polymer have a preference to interact with the molecules of their own kind,
leading to a potential for drug crystallization. The amount of drug below the solubility limit of the drug in the polymer can be solubilized by the polymer matrix
(51,52). On increasing the amount of the drug over this limit, the drug is no longer
soluble in the amorphous polymer, but can form separate crystallites (53). An analogy to the solubility of the drug in the polymer can be found from plasticization,
where flexible small molecules are mixed with the polymer to lower the Tg (53–55).
Also in plasticization, it is essential that the plasticizing material forms a uniform
mixture with the polymer, and is solubilized by the polymer. Consequently, from
this point of view, ketoprofen was a more effective plasticizing agent for the
Eudragit® L100 polymer than naproxen (53). Thus, the choice of the polymer is
important for controlling the amorphous–crystalline transition. The nanoparticles
were stored at three different conditions to study the influence of temperature and
relative humidity on the morphology of particles. The nanoparticles containing 33%
121
Aerosol Flow Reactor Method
(w/w) of ketoprofen and 67% of Eudragit® L100 polymer were stored for three
months in a refrigerator, at 25°C and 0% of relative humidity, and at 25°C and 75%
of relative humidity (sat. NaCl). After storage for three months, no changes in the
morphology of the nanoparticles were observed, that is, still spheres were observed.
As well, no crystallization of the drug within the nanoparticle occurred. To conclude, these nanoparticles were physically stable in all the conditions; however,
chemical stability was not studied during this period of time.
Nanostructured Drug Microparticles
Drug Crystallization in the Reactor
Amorphous materials tend to crystallize over storage causing problems in instability but also in powder handling. Therefore, the crystalline form of the drug is
preferred. The experimental conditions to crystallize beclomethasone dipropionate
particles while producing them will be discussed in this section (56). Two solutions,
5 and 25 g/L, were atomized in the heated reactor at different temperatures. Table 3
summarizes the physical characteristics of the particles. The saturated precursor
solutions prior to atomization included crystal seeds that were expected to induce
the particle crystallization. The sizes of the powders were in a respirable size range
(1–5 µm) and the size distributions were relatively narrow (i.e., GSD <2.1).
The main issue of this work was to explore the reactor conditions affecting the
particle crystallization while producing them. Table 3 also lists the state of matter
(crystalline and/or amorphous) in the final powders. Low solution concentration as
well as low reactor temperature resulted in smooth-surfaced spheres that were
amorphous. However, the increase in reactor temperature seemed to promote the
crystallization process in the reactor. Appearance of roughness on the particle
surface was indicative of crystals (Fig. 5A), and it was, in fact, proved by the XRD
analysis. For comparison, fully amorphous, smooth-surfaced BDP powder is shown
in Figure 5B. The evaporation of ethanol from the droplets is kinetically faster than
the organization of the drug molecules to crystals. Therefore, the BDP powders were
often amorphous. However, the diffusion of the BDP molecules in the amorphous
solid state depends on Tg and the overall rate of crystallization. The latter is at its
maximum at the mid-temperature between Tg and Tm (41). As the Tg and Tm were,
TABLE 3 Summary of the Production Conditions and Physical Characteristics of
the Beclomethasone Dipropionate Powders Prepared from Ethanolic Solutions
CBDP (g/L)
Treactor (°C)
MMAD (μm)
GSD
5a
25a
25b
25a
Saturatedb
150
50
100
150
100
1.4
2.5
1.7
2.2
2.1
1.8
2.0
2.0
2.0
2.1
State of matterc
Amorphous
Amorphous
Some crystallites
Crystalline
Crystalline core
The concentration of BDP in precursor solution (CBDP) and reactor wall temperature (Treactor) when producing drug microparticles are shown. Also mass median aerodynamic diameter and geometric standard deviation of the particle size distribution as determined with the Berner-type low pressure impactor
are shown.
aFlow rate 1.5 L/min through pyrosol.
bFlow rate 10 L/min through pyrosol, then divided into 5 tubes where 2 L/min.
cDirectly from the production.
Abbreviations: BDP, beclomethasone dipropionate; GSD, geometric standard deviation; MMAD, mass
median aerodynamic diameter.
122
Raula et al.
FIGURE 5 SEM images of the beclomethasone dipropionate powders in different state of matter: (A)
rough and crystalline and (B) smooth and amorphous.
respectively, ∼66°C (see the relation above) and ∼212°C, the maximum crystallization rate was expected to be around 139°C. This explains the significance of reactor
temperature for the particle crystallization. Thermal treatment accelerated crystal
growth in the sites of crystal nucleation. The slurry solution with crystal nuclei was
expected to promote heterogeneous nucleation and directing the crystal growth
(57). The particles produced at 100°C were spheres with smooth particle surface that
is a strong indication of amorphous state. However, the XRD analysis showed that
the powder was crystalline. The particles contained crystalline core of the original
crystal seeds. The crystals were covered by amorphous drug material.
Postannealing of the collected amorphous particles at various temperatures
initiated and/or accelerated crystallization. The crystallization unfortunately extended
between neighboring particles forming crystalline bridges between them. As
powders tend to agglomerate and form clusters, the crystal bridge formation was
inevitable. However, the agglomerates can be broken easily to individual particles
by mild milling.
Coating of Particles by a Peptide
The morphology as well as stability of the particles can be modified with a surfaceactive peptide such as l-leucine (58). This is analog to the stabilization of drug
nanoparticles by polymers as discussed earlier. The materials chosen for the preparation of the composite l-leucine powders have different physical and chemical
properties. NaCl crystallizes with ease, whereas large lactose molecules are not able
to organize during the droplet drying and the powders are amorphous in the collection. Moreover, using the relation Tg ∼ 0.7 Tm, the Tg of ASA was estimated to be
rather low, around 13°C. Accordingly, pure ASA microparticles should have poor
physical stability when collected.
Leucine derivatives are surface-active materials in aqueous solutions. Besides
the surface tension measurements, the surface activity of a compound can be determined by surface accumulation parameters that reflect the degree of hydrophobicity of the compound. This is calculated by the equation −(Δσ/Δc), where a change
in surface tension Δσ is measured as a function of solution concentration Δc (59). On
the flat surface, the −(Δσ/Δc) values for glycine, S-, l-, di-, and tri-leucine are −0.9
(σ values from Ref. 59), 40 (σ values from Ref. 59), 20 (σ values from Ref. 60), 870 (σ
values from Ref. 24), and 1414 (σ values from Ref. 24), respectively. The accumulation of the surface-active molecules on droplet surface is most probably enhanced
123
Aerosol Flow Reactor Method
FIGURE 6 SEM image (left ) of the pure L-leucine particles prepared from aqueous 15 g/L solution.
The scheme on the right presents the self-assembling of L-leucine molecules on the surface of the
droplet at the interface of water and air.
due to a high surface-to-volume ratio of the droplet. Pure l-leucine particles had
spherical but strongly wrinkled structure (see an SEM image in Fig. 6). l-Leucine
molecules gathered at the air–water interface on the droplet thus form a film. This
film increasingly prevents the evaporative removal of water as it thickens. A drying
particle expands while water molecules penetrate through the l-leucine film and
after complete removal of water the film collapses to give wrinkled morphology
(see the scheme in Fig. 6). During the surface organization, the l-leucine molecules
crystallize to some extent giving partly polycrystalline structure (see diffraction
patterns in the inset of Fig. 6).
Table 4 summarizes the physical characteristics of the produced composite
powders. The size of saline composite particles increased slightly with the added
l-leucine and the spread of size distribution did not vary, with GSD being around 2.5.
With the added l-leucine from 0 to 7.5 g/L, however, the size of lactose particles
decreased notably from 0.9 to below 0.5 µm, whereas the size distribution broadened
and GSD increased from 1.9 to 3.8.
TABLE 4 Summary of the Production Conditions and Physical Characteristics of the L-leucine
Composite Powders
Material
L-leucine
NaCl 20 g/L
NaCl 20 g/L
NaCl 20 g/L
Lactose 20 g/L
Lactose 20 g/L
Lactose 20 g/L
ASA 15 g/L
CL-leucine (g/L)
GNMD (μm)
GSD
State of matter
15
0
2.5
7.5
0
2.5
7.5
7.5
0.67
0.54
0.64
0.66
0.91
0.52
0.51
0.43
2.1
2.7
2.6
2.5
1.9
3.5
3.8
1.9
Partly polycrystallinea
Crystallinea,b
Crystallineb
Crystallineb
Amorphousa,b
Amorphousb
Amorphousb
Amorphousb
The concentrations of L-leucine in aqueous precursor solutions (CL-leucine) when producing drug microparticles at
the reactor wall temperature of 100 °C are shown. Also geometric number mean diameter and geometric standard
deviation of the particle size distribution as determined with the electrical low pressure impactor (ELPI) are
shown.
aTEM.
bXRD.
Abbreviations: GNMD, geometric number mean diameter; GSD, geometric standard deviation.
124
Raula et al.
FIGURE 7 SEM images of (A) NaCl, (B) NaCl/L-leucine, (C) lactose, and (D) lactose/L-leucine particles. In the precursor solutions the concentration of NaCl and lactose was 20 g/L, and that of L-leucine was 7.5 g/L.
The morphology of the composite particles was controlled by the amount of
l-leucine. The surface of lactose particles changed from smooth to wrinkled, whereas
cubic crystalline salt particles changed from spheroidal to leafy ones with increasing
l-leucine content (Fig. 7A–D). The XRD studies showed that the produced powders
were amorphous except saline composite particles that have the strong diffraction
peaks of NaCl at 31.7° and 45.5° (2θ). The structure of the NaCl composite particles
was studied by the evaporative removal of l-leucine in the oven at 200°C for
10 minutes (Tsub of l-leucine is ~145–148°C) (61). After the thermal treatment,
the remaining particle was a salt cube with holes on the surface. This showed
that the l-leucine molecules formed the shell around the salt core, partly mixing
with salt core. The structure of lactose composite powder could not studied because
the Tg of lactose monohydrate (~101°C) was lower than the temperature for
l-leucine removal.
The materials that have Tg below the collection temperature and do not crystallize in the reactor cannot be collected as compact individual particles. Above Tg,
the material is in rubbery state, and is soft and even fluidly. Pure ASA particles
produced from an ethanolic solution (15 g/L) were found flat and crystalline—a
rough particle surface is indicative of crystallinity—on the collector surface. They
flatten during collection with the formation of rubbery film that subsequently crystallizes after collection. Thus, l-leucine has been used to stabilize the structure of the
ASA particles. The aqueous l-leucine solution was mixed with the ethanolic ASA
solution and then droplets were produced. At low l-leucine concentration (2.5 g/L),
some flat particles were still observed (Fig. 8A) but at the concentration of 7.5 g/L
all the particles were spherical and compact with porous structure (Fig. 8B).
Aerosol Flow Reactor Method
125
FIGURE 8 SEM images of acetyl salicylic acid composite particles prepared from aqueous ethanol
solution. The concentration of ASA was 15 g/L and that of L-leucine was (A) 2.5 g/L and (B) 7.5 g/L
in ethanol/water solution. Flat and crystalline particles where L-leucine content is low (2.5 g/L) are
circled in the image.
The physical stability, that is, morphological changes of the powders were
examined in two conditions: at 25°C and 0% relative humidity, and at 25°C and 44%
relative humidity (sat. K2CO3) for three months. After three months, pure l-leucine
as well as all saline particles showed no changes in morphology. However, pure
lactose powder showed neck formation between particles even at 0% relative
humidity but this did not occur with l-leucine-containing lactose particles. The
surface of lactose particles was hardened by an l-leucine layer and such bridges
between neighboring particles were not observed.
CONCLUSIONS
The aerosol flow reactor method is a novel method to produce drug particles from
atomized solutions. The reactor can be modified to produce particles of different
sizes, that is, nano- or microparticles. This method has been used to prepare
spherical and amorphous matrix-type drug–polymer nanoparticles as well as
micrometer-sized dry powders for dry powder inhalation. Experimental conditions influenced drastically the particle formation and morphology. Solvent vapor
pressure as well as the crust formation on the droplet surface mainly determined
the resulting morphology of the polymer nanoparticles. Moreover, the solute
solubility in the starting solution also affected the particle shape. The state of drug
in the polymer nanoparticles depends on the drug molecule itself, that is, size and
chemical structure. With increasing size of the molecule, the diffusion within the
polymer matrix decreased. As it was shown, large drug molecule BDP was amorphous in all particles, whereas smaller molecules ketoprofen and naproxen were
amorphous to some extent. As the number of these latter molecules increased, they
crystallized and the nanoparticles coalesced. The difference in the interaction
between the drug and the polymer was observed as the change of the glass
transition temperature of the particles. The drug crystallization depended on the
thermal history experienced in the reactor. The molecular mobility increased with
increasing reactor temperature that in some cases produced fully or partly crystalline drug particles.
Particle morphology and surface properties of the microparticles can be
modified by the peptide l-leucine. Owing to the surface activity of the peptide in
126
Raula et al.
aqueous solutions, it is also possible to stabilize the structure of the drug particles
that are otherwise structurally unstable.
ACKNOWLEDGMENT
Financial support from the Academy of Finland is gratefully acknowledged.
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9
Supercooled Smectic Nanoparticles
Heike Bunjes and Judith Kuntsche
Department of Pharmaceutical Technology, Institute of Pharmacy,
Friedrich Schiller University Jena, Jena, Germany
INTRODUCTION
Among the different types of nanoparticulate drug carrier systems, lipid nanoparticles are particularly promising with regard to physiological compatibility as they
can be prepared predominantly or even exclusively on the basis of physiological
compounds. Colloidal lipid emulsions and solid lipid nanoparticles are the classical
examples of matrix-type lipid nanoparticles, which are mainly under investigation as
carriers for poorly water-soluble and lipophilic drugs. These two types of
nanodispersions each have their own advantages and disadvantages with regard to
their use as drug carrier systems. For example, colloidal lipid emulsions of natural
and semisynthetic oils have a solid basis of technological know-how and a welldocumented record of physiological compatibility arising from their decade-long use
in parenteral nutrition (1). Also their interactions with drugs have been characterized,
and some drug-loaded formulations are commercially available (1–3). The drug incorporation capacity of the liquid droplets is comparatively high in comparison with
solid lipid nanoparticles and can be modified by variation of the oil compound. On
the other hand, the liquid nature of the emulsion droplets may lead to disadvantages.
In cases of insufficient stabilization, the droplets may grow by coalescence; incorporated drug molecules have a high mobility and can easily migrate into the surfactant
layer or leak into the aqueous phase. Solid lipid nanoparticles with their rigid core
were developed to overcome these limitations (4,5). It has, however, turned out that
the usually crystalline nature of the solid particle core counteracts incorporation of
larger amounts of drugs so that the drug carrier capacity of these particles is frequently rather low (6). Moreover, the physical behavior of the systems is often quite
complex (e.g., with regard to crystallization, polymorphic transitions, and interparticle interactions) and certain proposed advantages of solid lipid nanoparticles, such as
the ability to provide better control of drug release, still remain to be proven.
In search of an alternative lipid carrier system, which could combine the
advantages of lipid emulsions and solid lipid nanoparticles, we were aiming to prepare carrier particles with a less ordered state than solid lipid nanoparticles in order
to achieve a higher incorporation capacity for foreign substances.The particles should,
however, still reduce the mobility of incorporated drugs and surface active agents
as much as possible. We focused our interest on liquid crystalline phases (also called
mesophases), which combine a certain mobility on the molecular level with an often
rather high macroscopic viscosity. Formation of liquid crystalline phases can be
induced by addition of solvents (lyotropic mesophases) or by temperature (thermotropic mesophases). Whereas lyotropic mesomorphism is frequently encountered
and also utilized in the area of pharmaceutics (7,8), thermotropic mesophases have
not yet received much attention in this field, although thermotropic mesomorphism
has been described for some drug substances (9). The solvent-free character of
lipophilic thermotropic mesophases makes them similar to the dispersed phase of
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colloidal lipid emulsions and suspensions, and thus promising to be explored as
potential matrix materials in colloidal drug delivery systems.
THERMOTROPIC MESOPHASES
While lyotropic mesophases are formed by amphiphilic molecules in the presence
of a suitable solvent, thermotropic mesomorphism is a specific property of certain
substances with strongly anisometric molecular shape that does not require any
additives and occurs in dependence on temperature (10). Thermotropic mesophases
are classified into two main groups: calamitic (formed by molecules with rod-like
shape, Fig. 1) and discotic (molecules with disc-like shape) mesophases. Two main
calamitic mesophases, the smectic and the nematic phases, can be distinguished.
When a substance forms a smectic as well as a nematic phase, the smectic phase
always exists at lower temperatures. In the smectic phase, the molecules are aligned
side by side forming a layered structure. Different modifications of the smectic phase
have been described, for example, the molecules can be arranged perpendicular
(smectic A phase) or tilted (smectic C phase) with respect to the smectic layers.
Smectic A phases are typical for saturated cholesterol esters such as cholesteryl
myristate (CM) and palmitate (CP); for cholesteryl oleate (CO), a smectic C phase is
assumed (12). In the nematic phase, the molecules are arranged nearly in parallel but
not in specific layers. The cholesteric phase, which is characteristic of cholesterol
esters, can be regarded as a twisted nematic phase. Individual nematic molecular
layers are twisted against each other in a certain angle forming a helical structure.
The distance between the molecular layers with the same orientation (pitch) depends
on temperature and is often in the range of the wavelength of visible light leading to
characteristic color effects.
Thermotropic mesophases are broadly applied in the technical field, for example, for liquid crystal displays, and, consequently, most known thermotropic mesogens have been developed for technical applications (13). Cholesterol fatty acid
esters as a class of physiological substances are, however, also capable of forming
thermotropic mesophases (12). As their smectic phase has a high viscosity, it
appeared promising to provide the desired characteristics for the development of
the novel type of lipid nanoparticles.
CHOLESTERYL MYRISTATE NANOPARTICLES: GENERAL
PROPERTIES AND PHASE BEHAVIOR
For first investigations on the preparation of smectic nanoparticles, the physiological ester CM was used as model matrix lipid due to its fully reversible and
FIGURE 1 Structures of thermotropic calamitic mesophases.
Source: Adapted from Ref. 11.
Supercooled Smectic Nanoparticles
131
FIGURE 2 (Top) Differential scanning calorimetry heating and cooling curves (5°C/min) of
CM in bulk and in colloidal dispersion (5% CM,
3.2% S100, 0.8% SGC). The samples were
heated to the isotropic melt (curve not shown
for the bulk material), cooled down below the
crystallization temperature, and heated again
(sm-ch: smectic–cholesteric, ch-i: cholesteric–
isotropic phase transition). (Bottom) Small- and
wide-angle X-ray diffraction patterns of CM in
bulk and in colloidal dispersion (5% CM, 3.2%
S100, 0.8% SGC). Bulk: crystalline powder
(A) at 20°C and smectic mesophase (B) at
60°C (formed upon cooling the isotropic melt).
Dispersions: stored at 4°C (C) and at 23°C (D)
measured at 20°C. The graphs of the bulk
material and the dispersions are not on the
same linear intensity scale (s = 1/d = 2 sin Θ/
λ where 2Θ is the scattering angle and λ is the
wavelength, equal to 0.15 nm). Abbreviations:
CM, cholesteryl myristate; SGC, sodium glycocholate; S100, purified soy bean lecithin Lipoid
S100 (Lipoid KG, D-Ludwigshafen).
well-characterized phase behavior (11,14). Crystalline CM melts around 72°C into a
smectic phase, transforms into the cholesteric phase around 79°C, and finally melts
into an isotropic liquid around 84°C (Fig. 2). Upon cooling the melt, the liquid
crystalline phase transitions are observed at about the same temperatures as upon
heating. Crystallization occurs between about 40°C and 30°C. Thus, the smectic
phase of bulk CM does not exist under pharmaceutically relevant conditions such
as body and room temperature. For colloidal particles of CM, the same phase transitions are, in principle, observed as for the bulk material except for the distinctly
higher supercooling of the smectic phase upon cooling (Fig. 2). The fact that CM
crystallization starts only below 20°C in the colloidal state is the basis for the preparation of smectic nanoparticles. When the hot dispersions are only cooled to room
temperature and stored under this condition, their smectic state can be retained for
many months (11,14). Upon heating of a stable smectic CM dispersion, only the
liquid crystalline phase transitions are observed. These transitions are very small
and, for the colloidal dispersions, broad and not well separated, making their quantification difficult. As an additional feature to characterize the smectic state of the
nanoparticles, the very sharp small-angle X-ray reflection arising from the layered
structure of the smectic phase can be used (Fig. 2). This smectic reflection is broader
and less intensive for the nanoparticles, but its position is comparable with that of
the bulk material when measured under comparable conditions. As the repeating
unit (spacing) of the smectic phase increases with decreasing temperature (15),
a larger spacing of the strongly supercooled smectic nanoparticles is measured at
20°C (11). In the wide-angle range, only diffuse X-ray scattering is observed because
of the lack of molecular order. For the crystalline material (the nanoparticles can be
crystallized by storage at, e.g., 4°C), characteristic wide-angle reflections are visible
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Bunjes and Kuntsche
beside small-angle reflections of first and higher orders. For the crystalline nanoparticles, the reflections are less sharp, but their positions are comparable with those of
the bulk lipid.
PREPARATION METHODS FOR SMECTIC NANOPARTICLES
Aqueous dispersions of smectic cholesterol ester nanoparticles can be prepared by
high-pressure homogenization of a hot pre-emulsion (obtained, e.g., by ultra-turrax
vortexing) at temperatures above the melting point of the respective matrix lipid in the
presence of emulsifiers. The process, known as high-pressure melt-homogenization,
is also commonly used for the preparation of solid lipid nanoparticles (4,5) and, in
principle, similar to the preparation of colloidal fat emulsions (3,16). After homogenization (using a microfluidizer or piston-gap homogenizer), the hot nanoparticle
dispersions are usually filtered (0.2 or 5 µm). High-pressure melt-homogenization
yields dispersions of smectic nanoparticles with mean particle sizes between 100
and 200 nm, avoiding the use of organic solvents (11,14).
The high temperature during the process of melt-homogenization may lead to
a partial degradation of thermally sensitive compounds and drugs. To avoid this
problem and for the preparation of dispersions with smaller particles, a modification of the so-called emulsification solvent-evaporation method (17) can be used.
The lipid and lipophilic stabilizers are dissolved in an organic solvent, which is not
miscible with water, for example, cyclohexane. A crude pre-emulsion with the
aqueous phase (containing hydrophilic surfactants) is prepared by ultra-turrax
vortexing and the dispersion is high pressure homogenized, for example, using a
microfluidizer. The whole process is carried out at room or slightly lower temperature.
The organic solvent is removed under reduced pressure and at slightly elevated
temperature. The dispersions are filtered and heated above the melting temperature
of the matrix lipid for about 10 minutes to ensure the smectic state of the nanoparticles. With this method, dispersions with mean diameters distinctly below 100 nm
can be prepared (11). Residual organic solvent in the dispersions may, however, be
a disadvantage of this procedure (18,19).
Particle size is a crucial point in the development of dispersions that are stable
upon storage with regard to the metastable smectic state of the matrix material, as
has been shown for CM dispersions stabilized on the basis of phospholipids.
Dispersions of smaller particles are more stable against recrystallization upon storage than dispersions with larger particle sizes (14). Moreover, the storage temperature has to be adjusted to the dispersion properties: If the dispersions are stored too
close to the recrystallization temperature of the nanoparticles, the matrix lipid will
crystallize during storage.
For administration of nanoparticles via the parenteral or ocular route, sterility
of the dispersions is an important issue. First studies indicate that dispersions of
supercooled smectic nanoparticles can be sterilized by autoclaving at 121°C without
increase in particle size (stabilization with phospholipid/bile salt mixtures) or with
only a small increase in particle size [stabilization with poloxamer 188 or poloxamine (Tetronic 908, BASF, D-Ludwigshafen)]. Small-sized dispersions can also be filtered through 0.2-μm filters as outlined above.
INFLUENCE OF THE STABILIZER SYSTEM
Colloidal dispersions are thermodynamically unstable systems because of their high
interfacial energy. For the preparation and stabilization of colloidal lipid particles,
Supercooled Smectic Nanoparticles
133
surface active agents, which accumulate in interfaces and reduce the interfacial
energy, are required. Depending on the properties of the surface active agent, stabilization occurs by electrostatic (charged surfactants) or steric (e.g., amphiphilic
polymers) stabilization or a combination of both. Stabilizers may also influence the
phase behavior of dispersed lipids, like their crystallization and polymorphic
behavior (20–22). Furthermore, the stabilizer system determines the surface properties influencing the fate of the nanoparticles in vivo (23,24). For the stabilization of
smectic CM nanoparticles, purified phospholipids (soybean and egg-yolk phospholipids) alone or in combination with the bile salt sodium glycocholate, sodium
glycocholate alone, sodium oleate, different polymers [poloxamer, poloxamine,
partially acetylated polyvinyl alcohol (PVA)], Tween 80, a sucrose ester mainly
containing sucrose monolaurate, sodium caseinate, and gelatin polysuccinate have
been tested so far. With exception of the sugar ester, all stabilizers led to stable
colloidal dispersions after melt-homogenization with respect to macroscopic appearance and particle size upon long-term storage. The smectic state of these nanoparticles could clearly be identified by the characteristic small-angle X-ray reflection
without influence of the stabilizer system on the position of the reflection.
The stabilizer system does, however, strongly influence the phase behavior of
CM nanoparticles, particularly the crystallization process (25). On the basis of the
crystallization pattern and the recrystallization tendency upon storage, two groups
of stabilizers can be distinguished:
■
■
Stabilizers inducing a multiple crystallization event depending on the thermal
history of the sample and a high recrystallization tendency upon storage
(phospholipids, sodium oleate, sucrose monolaurate).
Stabilizers inducing a monomodal crystallization event mostly independent of
the thermal history of the sample and a low recrystallization tendency upon
storage (sodium glycocholate, synthetic polymers, gelatin polysuccinate, sodium
caseinate).
The common feature of stabilizers resulting in smectic nanoparticles with a
multiple crystallization pattern, and a comparatively high recrystallization tendency
upon storage is the presence of an acyl chain in the molecule. All corresponding
dispersions show a very complex crystallization behavior (Fig. 3), which is probably
caused by particles of different shapes (spherical and cylindrical) present in the
dispersions as observed in electron microscopy (see below) (14). In contrast, smectic
nanoparticles stabilized with polymers and sodium glycocholate alone have a distinctly lower crystallization tendency upon storage despite their relatively high
crystallization temperature (Fig. 3). Electron microscopic investigations of these
dispersions indicate a more homogeneous particle structure (see below). The dispersion stabilized with Tween 80 (a stabilizer which also contains an acyl chain)
cannot be clearly classified into one of the groups mentioned above. Although stored
smectic nanoparticles stabilized with Tween 80 display only one crystallization
event upon cooling in differential scanning calorimetry (DSC) and the recrystallization tendency of these smectic nanoparticles is low, freshly melted nanoparticles
give a bimodal crystallization pattern upon cooling.
INFLUENCE OF THE MATRIX COMPOSITION
CM nanoparticles are excellent model systems for studying the influence of parameters such as particle size and stabilizer system on storage stability and phase
behavior. Their comparatively high crystallization temperature, which is observed
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Bunjes and Kuntsche
FIGURE 3 Crystallization onset temperatures
(differential scanning calorimetry, cooling of
thermally untreated samples, 5°C/min) (top)
and amount of recrystallized matrix lipid after
storage (bottom) for nine months (eight months
for samples marked with an asterisk) in dependence on the stabilizer system. All dispersions
contain 5% CM and, with exception of SGC (2%),
4% stabilizers. Abbreviations: CM, cholesteryl
myristate; E80, purified egg yolk lecithin Lipoid
E80; PVA, polyvinyl alcohol; S100, purified soybean lecithin Lipoid S100; SGC, sodium glycocholate; sucrose ML, sucrose monolaurate.
even in optimized systems, precludes, however, their development into a robust
drug delivery system for practical use. Such a system should display a high stability
against recrystallization upon storage, preferably also at temperatures down to
about 0°C which can occur during transport and may be required during storage to
protect sensitive drugs or excipients, for example, phospholipids, from degradation. Therefore, cholesteryl nonanoate (CN), an ester with nine carbon atoms in the
acyl chain, and CO, containing an unsaturated C18 chain, were tested as alternative
matrix materials to increase the crystallization stability of smectic nanoparticles
(26). Both esters do not crystallize in the colloidal state, neither during cooling to
−10°C nor upon storage (even not at refrigerator temperatures). Pure CO is, however,
unsuitable as matrix lipid, because its liquid crystalline to isotropic phase transition
is very close to body temperature so that CO nanoparticles would lose their smectic
state upon administration. The use of CN seems to be very promising from the
physicochemical point of view, but its nonphysiological nature requires some safety
issues to be resolved prior to its use as drug carrier.
Both esters can also be used as additives to longer-chained cholesterol esters
such as CM or CP, which has an even higher crystallization tendency than CM, for
the preparation of smectic nanoparticles being more stable against recrystallization.
The admixture of already relatively small amounts of these esters to CM (10%, 20%,
40% CN, 20% CO related to the whole matrix lipid) or CP (40%, 50% CN, 50% CO
related to the whole matrix lipid) leads to smectic nanoparticles which are stable
against recrystallization over at least 18 months at 23°C although the crystallization
temperature is only slightly decreased particularly for the CM-based dispersions
Supercooled Smectic Nanoparticles
135
FIGURE 4 Particle size [photon correlation spectroscopy (PCS)] and crystallization onset temperatures of freshly molten samples (differential scanning calorimetry, 5°C/min) in dependence on the
matrix composition (left) and stabilizer system (right ). For the dispersion with a CM/CN-matrix
stabilized with polyvinyl alcohol, only the very beginning of the crystallization event was recorded so
that the onset temperature (<0°C) could not be exactly determined. Abbreviations: CM, cholesteryl
myristate; CN, cholesteryl nonanoate; CO, cholesteryl oleate; CP, cholesteryl palmitate; PVA, polyvinyl alcohol; SGC, sodium glycocholate; S100, purified soy bean lecithin Lipoid S100.
(Fig. 4). In contrast, the amount of recrystallized matrix lipid was 8% and 94% for
the dispersions with a pure CM or CP matrix, respectively. Admixture of higher
amounts of CN (60% of the matrix) to CM smectic nanoparticles resulted in dispersions which could be stored at 4°C without nanoparticle recrystallization. The crystallization temperature-suppressing effect of CN admixture can be further increased
by the use of polymeric stabilizers such as PVA and poloxamer in dispersions with
a mixed cholesterol ester matrix (Fig. 4).
INCORPORATION OF MODEL DRUGS
Ibuprofen, miconazole, etomidate, and progesterone were used as poorly watersoluble model drugs for the preparation of drug-loaded smectic nanoparticles (11).
The dispersions containing 5% CM as matrix lipid and a phospholipid/bile salt mixture (3.2/0.8%) as stabilizers were prepared by high-pressure melt-homogenization.
The lower melting drugs ibuprofen, miconazole, and etomidate could be incorporated in an amount of 10% relative to the lipid matrix by dissolving the drug in the
cholesterol ester melt. Higher drug loads have not been investigated yet but a drug
load of 10% is already higher than achieved for solid lipid nanoparticles with
miconazole and ibuprofen (27). Progesterone, which melts above 100°C, could be
dissolved only at a concentration of 1% in the CM melt within an appropriate time
(<30 minutes).
Drug incorporation into the dispersions did not influence the occurrence and
position of the small-angle X-ray reflection characteristic of the smectic nanoparticle
state. Particularly the liquid crystalline phase transitions were, however, shifted to
lower temperatures in DSC, indicating an interaction of the drug molecules with the
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Bunjes and Kuntsche
FIGURE 5 Crystallization onset temperatures (differential scanning calorimetry, 5°C/min, freshly molten samples,
main crystallization event) of drug-loaded
dispersions in comparison to those of a
corresponding drug-free dispersion with
comparable particle size in dependence
on storage time.
smectic phase. The decrease of the liquid crystalline phase transition temperatures
was different and least pronounced in the dispersion with 1% progesterone. Except
for the dispersion loaded with progesterone, the crystallization temperature was
also decreased in the presence of drug (Fig. 5), indicating that drug incorporation
may have a beneficial effect on the recrystallization tendency.
The drug-loaded dispersions were stable upon storage for at least 12 months
with respect to particle size, macroscopic appearance, and recrystallization of matrix
material. Only in the progesterone-loaded dispersion was a small amount of recrystallized matrix lipid (4%) detected after 12 months of storage.
First preliminary studies point to a rapid release of at least a major fraction of
ibuprofen and etomidate from the smectic nanoparticles into the aqueous phase.
Such behavior would lead to carrier-independent drug distribution after IV administration as is also often observed for lipid emulsions (3). The use of more lipophilic
prodrugs (e.g., esters) or ion-pairs is expected to enhance drug retention after
administration and might also allow loading with comparatively hydrophilic
compounds (28–31).
ULTRASTRUCTURE OF CHOLESTERYL MYRISTATE NANOPARTICLES
Different electron microscopic techniques can be used for the investigation of the
ultrastructure of colloidal lipid particles. In cryoelectron microscopy, a thin film of
rapidly frozen dispersion is viewed directly, and a good impression of the shape of
the projected particles is obtained. In contrast, freeze-fracturing, where the frozen
sample is fractured, and the fracture plane shadowed with a thin heavy metal layer
leads to images that yield information also about the inner structure of the particles.
Electron microscopic images indicate a mostly nonspherical, nearly cylindrical shape of the CM nanoparticles. The particle shape is strongly influenced by the
stabilizer system ranging from nearly exactly cylindrical particles to “paving stonelike” ones (Fig. 6). In dispersions stabilized on the basis of phospholipids and with
sodium oleate, an additional particle fraction was observed in cryo-preparations
which was very unstable in the electron beam. An onion-like internal structure
detected for some particles in freeze-fractured specimen of phospholipid-stabilized
dispersions points to a spherical shape of these particles. Consequently, these dispersions seem to contain two different types of smectic nanoparticles—cylindrical
and spherical ones. Owing to the layered structure of the smectic phase, a cylindrical particle shape should be energetically more favorable than a spherical one and,
Supercooled Smectic Nanoparticles
137
FIGURE 6 Cryoelectron micrographs of selected dispersions containing 5% CM. (A and B)
Dispersion stabilized with the phospholipid/bile salt system (3.2/0.8%). (B) The fraction of instable,
“bubbling” particles is clearly visible. (C) Dispersion stabilized with 4% sodium oleate. (D) Dispersion
stabilized with 4% polyvinyl alcohol. (E) Dispersion stabilized with 4% poloxamer 188. (F) Dispersion stabilized with 4% Tween 80. Magnifications illustrate the dependence of the particle shape on
the stabilizer system. The freeze-fracture micrograph shows an image of a CM particle with an
onion-like inner structure (3.2% S100, bar = 140 nm). A schematic model representation of the
structure of cylindrical and spherical CM nanoparticles is also given. Abbreviations: CM, cholesteryl
myristate; S100, purified soy bean lecithin Lipoid S100; SGC, sodium glycocholate.
therefore, the increased sensitivity of the spherical particles toward the electron
beam might be caused by a higher energy level of these particles. A cylindrical shape
has also been observed in cryoelectron microscopic investigations of native lowdensity lipoprotein (LDL) particles (32,33) which contain a core of cholesterol esters
that should be in the smectic phase under the conditions used for sample preparation
at room temperature (34,35).
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Besides smectic nanoparticles, other colloidal structures such as micelles,
mixed micelles, and vesicles can be present in the dispersions depending on the selfassociation characteristics of the stabilizers used. Particularly in the dispersion stabilized with sodium oleate, a variety of different structures was observed in agreement
with literature data on cholesterol ester-free samples (36) (Fig. 6). The presence of
such other colloidal structures is generally not desirable as they may also solubilize
lipophilic drugs and could induce redistribution processes of incorporated drugs
upon storage. Moreover, they may lead to additional physiological effects. For
example, liposomes formed by an excess of phospholipids used for the stabilization
of colloidal fat emulsions for parenteral nutrition (37) influence fat metabolism and
can lead to a decreased degradation of the fat emulsion particles (38).
SUMMARY AND CONCLUSION
Supercooled smectic cholesterol ester nanoparticles are a promising new carrier
system for the delivery of lipophilic drugs. They can be prepared by different methods, for example, with the established high-pressure melt-homogenization technique. Their stability against recrystallization depends on the matrix composition,
emulsifier system, and on the particle size. If properly designed with respect to
composition, preparative, and storage parameters, they are stable on storage in
liquid dispersion with regard to particle size and liquid crystalline state of the matrix
material for pharmaceutically relevant periods of time. The high viscosity of the
liquid crystalline matrix is expected to counteract coalescence processes. Lipophilic
drugs can be loaded into the nanoparticles as shown above for ibuprofen, etomidate, and miconazole which can be incorporated at considerable concentration.
Although basic knowledge has already been collected on the characteristics of this
new type of particles, they are still in an early stage of development as drug delivery
systems. Future activities will focus on the further optimization of their composition, in particular with respect to stability against recrystallization, drug incorporation, and retention as well as their surface characteristics.
The lipidic nature and small particle size of supercooled smectic nanoparticles
offer several interesting possibilities with virtually all ways of administration (e.g.,
parenteral, ocular, dermal, or peroral). In future, corresponding interactions with
biological systems will have to be studied in detail. First cytotoxicity studies have
already revealed promising results. Concerning parenteral, in particular intravenous
administration, the influence of the new type of matrix structure on the pharmacokinetics and biodistribution requires intensive investigation. As supercooled smectic nanoparticles show similarities in composition and structure with the protein-free
core of LDLs (32,33), they may have interesting potential as easily accessible artificial
constructs with similar properties, for example, for drug targeting to certain tumors
or the brain via the LDL receptor (39–41).
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10
Biological and Engineering Considerations
for Developing Tumor-Targeting Metallic
Nanoparticle Drug-Delivery Systems
Giulio F. Paciotti and Lawrence Tamarkin
CytImmune Sciences, Inc., Rockville, Maryland, U.S.A.
INTRODUCTION
Currently, first-line treatment of resectable solid tumors most commonly involves
surgery followed by a regimen of chemotherapy and/or radiation. Unfortunately,
this strategy often fails because of recurrent or metastatic disease. To change this
paradigm, new cancer therapies must deliver multifunctional therapeutics capable of
destroying the heterogeneous population of tumor cells present within solid tumors.
Targeting cancer therapeutics to solid tumors is facilitated by particle-delivery
systems capable of escaping phagocytic clearance by the reticuloendothelial system
(RES) (1–3). Under ideal conditions, such delivery systems preferentially extravasate
the tumor vasculature and accumulate within the tumor microenvironment (4,5).
Additionally, these nanotherapeutics may be engineered to contain tumor-targeting
ligands that bind to specific cells within solid tumors to anchor the nanoparticle
within the solid tumor. By design, particle-delivery systems capable of sequestering
cancer drugs solely within a tumor may also reduce the accumulation of the drugs
in healthy organs (1–5). Consequently, these delivery systems may increase the
relative efficacy or safety of cancer therapies, and thus serve to increase their
therapeutic index.
In recent years, the field of nanoparticle-based drug delivery has been reinvigorated by a convergence of nanotechnology and medicine (6–9). In essence, the blending of these fields is leading to the generation of innovative synthetic vectors with the
potential of achieving the long sought after goal of tumor-targeted drug delivery: getting the active agent(s) solely where they are needed, the solid tumor. Furthermore,
the versatility of these nanoparticle systems may for the first time lead to the development of multifunctional nanotherapeutics (Fig. 1) that detect and attack the heterogeneous population of tumor cells present in a solid tumor. As shown in Figure 1, these
putative vectors may consist of an immune-avoidance moiety (a “stealth” moiety), a
tumor-targeting motif, multiple therapeutics, and a diagnostic/sensing component,
all of which are delivered on a single nanoparticle no larger than 50 nm.
Over the past 20 years, the field of nanoparticle-based drug delivery focused
on two chemically distinct colloidal particles: liposomes and biodegradable polymers (10–14). Both delivery systems encapsulate/entrap the active drug within their
structures and release the active agent as the particle lyses, in the case of liposomes,
or disintegrates as described for biodegradable polymers. A recent newcomer to this
field is the metallic nanoparticle. Like their organic counterparts, the metallic particles have a long history of use in biology; however, only recently have they been
formulated into tumor-targeting vectors.
This chapter is divided into three parts. The “Biological Considerations” section
focuses on the biological barriers facing metallic-nanoparticle-based, tumor-targeted
141
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Therapeutic 1
Stealth Moiety
Therapeutic 2
Diagnostic/
Sensing
Component
Therapeutic 3
Tumor
Targeting
Motif
FIGURE 1 Schematic representation of a multifunctional nanoparticle
therapeutic.
drug delivery, and will center on the delivery of protein and small-molecule-based
therapies. Some of these barriers are naturally present in healthy portions of the
body, whereas others are established during tumor growth and progression. The
section “Metallic Nanoparticle Drug-Delivery Systems” will describe three examples of metallic-nanoparticle-based, tumor-targeted drug delivery and the specific
means used to overcome these barriers. Specifically, an overview of magnetic-iron
and colloidal gold-based nanoparticle drug-delivery systems will be discussed. The
final section will focus on the design and manufacturing of multifunctional metallic
nanoparticles that target the delivery of multiple cancer therapies with a single
metallic nanoparticle.
BIOLOGICAL CONSIDERATIONS
Clearance by the Reticuloendothelial System
The painstaking lessons learned from the liposomal- and polymer-based drugdelivery systems must be considered in the formulation of the metallic nanoparticles. There is ample precedent suggesting that immediately after their exposure to
the circulation, unprotected metallic nanoparticles will most likely undergo opsinization in the blood and clearance by the RES. Opsinization is the process by which
blood-borne proteins, known as opsins, bind to the metallic nanoparticle surface.
Opsonic proteins include immunoglobulins, complement proteins C1 and C3,
apolipoproteins, von Willebrand factor, thrombospondin, fibronectin, and mannosebinding protein (15–19).
Once bound to the surface of the particles, opsins act as ligands that are
recognized, bound, and internalized by a class of macrophages known as Kupffer
cells. Kupffer cells are specialized macrophages that reside in the liver and represent
the primary cellular component of the scavenging system of the body, known as the
RES. Under physiologic conditions, the RES rapidly and very efficiently clears
opsinized particulates, such as bacteria and colloidal particles, from the body. For
example, in mice intravenously injected with unprotected colloidal gold nanoparticles (i.e., colloidal gold nanoparticles that contain only drug bound to their surface),
we observed that 90% to 95% of these gold nanoparticles are cleared from the circulation within 5 to 10 minutes after injection. These observations are consistent with
the historical biodistribution data of unprotected liposomes.
Early work in the field of particle-based drug delivery demonstrated that RES
clearance was saturable, because pretreating animals with high doses of drug or
placebo nanoparticle vectors (i.e., vectors lacking the active product ingredient)
Biological and Engineering Considerations
143
exceeded the phagocytic capacity of the RES (20–24). Consequently, nanoparticles
not trapped in the liver or spleen were available to deliver drugs to solid tumors.
A more practical approach to address RES uptake and clearance is the modification
of the nanoparticle surface to include hydrophilic blockers such as polyethylene
glycol (PEG) as well as block copolymers of the tetronic and pluronic families of
surfactants. Grafting/binding these hydrophilic moieties onto the surface of the
colloidal nanoparticles made them “invisible” to the RES. Typically, such sterically
stabilized nanoparticles exhibit prolonged circulatory half-life and as discussed
below such nanoparticle formulations may passively accumulate in solid tumors
through extravasation of the leaky vasculature that feed them (25–28).
The nature and conformation of potential stealth molecules influence the
ability of the immune system to detect and clear nanoparticles. As an example,
Moghimi (29) demonstrated that the concentration and conformation of one such
polymer, poloxamer 407, significantly altered the biodistribution of polystyrene
nanoparticles after subcutaneous injection. When the ethylene oxide (ETOX) tails of
the poloxamer polymer were grafted at relatively low concentrations, the molecule
assumed a flat or mushroom-like configuration on the surface of the particles,
whereas at higher polymer concentrations the closely packed ETOX tails assumed a
brush-like conformation on the surface of the particles. The particles with the
mushroom-like configuration were readily picked up by the macrophages present
in the primary or secondary lymph nodes, whereas those carrying the brush-like
configuration evaded the lymph nodes, traveled through the lymphatic system and
reentered the circulation.
In our laboratory, we also observed that the manner by which a stealth
molecule binds to drug-coated colloidal gold nanoparticles influences their biodistribution (30). As described above, 95% of various formulations of drug-stabilized,
monodispersed colloidal gold nanoparticles [nanoparticles bound with a saturating
amount of the protein therapeutic, tumor necrosis factor (TNF) alpha], were cleared
by the RES within 5 to 10 minutes of intravenous injection. The fate of these
monodispersed nanoparticles was evident upon animal necropsy as both the livers
and spleens were black with aggregated colloidal gold precipitates. Several block
copolymers including pluronic, tetronic, carbowax, and monothiolated forms of
PEG (PEG-THIOL) were tested for their ability to block RES uptake and clearance.
Although all four classes of polymers bound and stabilized the colloidal gold
nanoparticles in a test tube, only PEG-THIOL prevented RES uptake and clearance
when injected intravenously. Unlike the poloxamer-stabilized nanoparticles, we
believe that the pluronic-, tetronic-, and carbowax-based polymers were unable to
bind to the surface of the colloidal gold nanoparticle. However, the presence of the
single sulfhydryl group on the PEG molecule (PEG-THIOL) allowed it to bind directly
to the surface of the particle, interspersed between the protein therapeutic. In this
case, we believe that the molecules of PEG-THIOL assumed a brush-like configuration on the surface of the nanoparticles, and the biological response (i.e., avoiding
immune recognition) was similar to that seen in the example above.
Surface characteristics including size, surface contour/facets, and charge
may also regulate which opsins bind and the extent to which they adhere to the
surface of the particles. Particle size can affect particle clearance by two mechanisms
(31–34). First, larger particles more efficiently activate the complement pathway for
clearance by the RES. Second, in the red pulp of the spleen, the interendothelial slits
present on the endothelial cells of the venous sinusoids may filter particles, merely
based on size (larger than 200 nm). The latter effect seems more problematic for the
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earlier formulations of liposomes rather than metallic nanoparticles as metal-based
nanoparticles range in size from 5 to 40 nm. Also, as described below, particle size
may influence the ability of nanoparticles to effectively deliver drugs to the tumor
interstitium.
Surface charge must also be considered in formulating metallic nanoparticles
because the overall surface charge may determine the degree to which particles are
cleared by the RES, as well as their ability to interact with endothelial cells that form
the tumor neovasculature. For example, the RES readily clears anionic lipid particles,
whereas neutral lipid particles are not readily cleared (18,33–35). Cationic liposomes
containing an RES-avoidance system (i.e., PEG) preferentially bind to the surface of
tumor endothelia (36,37).
The process of particle opsinization and RES clearance is an active process by
which particle-bound opsins are recognized, bound, and internalized by the macrophages present in the liver and spleen. Overcoming this biological barrier is just the
first step in tumor-targeted drug delivery. So, even when RES clearance is adequately
addressed, the physical barriers established during solid tumor formation and progression must next be addressed for the development of a successful tumor-targeted
metallic drug delivery vector.
Tumor Angiogenesis and Vascularization
In 1971, Folkman (38) laid out his hypothesis describing the angiogenic model of
tumor growth and metastases. A central component of this hypothesis is that solid
tumors, in response to physiologic stresses such as hypoxia, secrete factors that cause
new blood vessels to sprout and grow from existing blood vessels toward the tumor.
Eventually these angiogenic blood vessels become inexorably entwined in the solid
tumor mass. Such angiogenesis is required to ensure continued tumor growth and
progression, and this process is currently the focus of many novel cancer therapies.
Angiogenesis, the process by which new blood vessels sprout from existing
normal blood vessels, occurs in a variety of disease states including cancer and
inflammation as well as during normal physiological events such as the thickening
of the uterine wall, follicle development, and wound healing. During wound healing,
for example, angiogenesis and clot formation are highly interconnected processes
that are regulated by platelets and soluble blood proteins (39).
Upon injury, tissue factor is released and induces clot formation through the
well-characterized cascade of clotting factors that ultimately lead to the formation
of a clot by the thrombin-mediated conversion of fibrinogen to fibrin. Thrombin also
activates platelets to release a variety of potent pro-angiogenic factors including
various forms of vascular endothelial growth factors A, B, and C (VEGF A, B, and C,
respectively), fibroblast growth factor-2, and angiopoietin-2. Finally, thrombin
stimulates endothelial cells to secrete a variety of enzymes responsible for the
degradation of the basement membrane causing the release of VEGF from these
compartments (39,40).
Platelets may also be recruited and activated to secrete these pro-angiogenic
factors by interaction with collagen or indirectly through interaction with von
Willebrand factor. These pro-angiogenic factors signal the surrounding vascular
endothelial cells to grow and migrate through the fibrin clot (41). To keep the
angiogenic process in check, platelets also secrete antiangiogenic factors such as
thrombospondin-1, plasminogen, and TGFβ-1, which serve to block angiogenesis,
and induce the expression of enzymes that help break down the clot. Tight control
Biological and Engineering Considerations
145
between the pro- and antiangiogenic signals ensures that as the process of tissue
repair and remodeling is completed, angiogenesis is terminated.
Within solid tumors, however, the homeostatic mechanism for controlling new
blood vessel formation is aberrant, often favoring angiogenesis and continued tumor
growth and metastasis. For example, the constitutive expression of tissue factor by
tumor and supporting stromal cells may induce the clotting mechanism by circulating coagulating factors, such as factor VIIa, to ultimately cause platelet-mediated
coagulation and release of VEGF (42,43). Indeed there is a reasonable precedent in
the literature showing a potential link between the state of coagulation and tumor
progression (44,45). Gasic et al. (46) demonstrated this potential link by showing
that antibodies which blocked platelet function inhibited the formation of pulmonary metastases. More recently, Cramerer et al. (47) highlighted a model in which
circulating tumor cells use secreted tissue factor as a means of inducing the clotting
mechanism to arrest their migration from the normal vasculature, to traverse the
endothelial cell barrier, and ultimately to form secondary metastases.
A newly identified and more problematic cell type may provide solid tumors
with not only angiogenic support, but also with strong protection against the ability
of the immune system to induce an antitumor response. Tumor-associated macrophages (TAMs) are monocyte-derived macrophages that migrate from the vasculature to regions of hypoxia in solid tumors. Unlike wild-type macrophages that are
present in normal tissues and in well-oxygenated areas of tumors, TAMs exhibit
altered phenotypes. For example, TAMs secrete immunosuppressive cytokines,
such as IL-10, TGFβ, and prostaglandins (48–50). Finally, TAMs residing in the
hypoxic centers of tumors are not effective at antigen presentation (51). Thus, TAMs
may indirectly support tumor growth by blunting the ability of the immune system
to mount an effective antitumor response.
In fact, once trapped in the hypoxic regions of tumors, TAMs may actually
begin to support tumor growth. In response to the low oxygen tension present in the
hypoxic regions of solid tumors, TAMs upregulate the expression of hypoxiainducible factors one and two (HIF-1 and HIF-2, respectively) (52–54). Tumor cells
also express these factors in response to hypoxia (55). The HIF family of proteins are
transcription factors that translocate to the nucleus, bind specific DNA regions
known as hypoxia-response elements to stimulate gene expression, and ultimately
lead to VEGF secretion by the macrophages and tumor cells (52,55,56).
Vascular Endothelial Growth Factor
VEGF was identified in the late 1970s for its ability to induce vascular permeability
of venules and small veins present in the postcapillary plexus. VEGF-induced
permeability is mediated by the formation of vesicular vacuolar organelles (VVOs)
(57–59). VVOs are a series of interconnected vesicles and vacuoles that span the
entire distance of endothelial cells from the lumenal to the ablumenal side of the
endothelial cells. In the absence of VEGF, access of macromolecules to the VVOs is
limited by the presence of a membrane-covered stomata which is connected to the
plasma membrane of the endothelial cell. In response to VEGF stimulation, the
stomata open to allow passage of tracer elements, plasma, and plasma proteins from
the lumenal side of the endothelial cell, through the VVOs, to the extravascular
space of tissues.
Additional structures such as intercellular openings and intracellular holes
may also increase the permeability of tumor blood vessels. For example, Hashizume
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et al. (60) described round or oval openings between the endothelial blood vessel
cells supporting the growth of MCa-IV tumors. These openings are much larger
than the fenestrations (50–80 nm) (60) reported for tumor blood vessel endothelium
and ranged from 0.3 to 4.7 µm with an average of 1.7 µm. In addition, transcellular
holes ranging from 0.2 to 0.9 µm were also reported. Although additional study is
needed, the presence of such extra-fenestral openings suggests a potential mechanism by which nanoparticle-delivery systems passively accumulate in solid tumors
rather than other healthy filtering organs such as the liver, an organ in which fenestrae ranging from 50 to 300 nm have been described.
The increase in vascular permeability and the deposition of a fibrin clot are
among the responses attributed to VEGF-mediated angiogenesis. Within the fibrin
clot, VEGF-mediated angiogenesis starts with the generation of a mother vessel, a
vessel characterized by poor coverage with both a basement membrane and pericytes. Subsequently, the formation of new blood vessels may occur by one of three
mechanisms. First, the mother vessel may undergo sprouting, endothelial cell
migration, and proliferation. Second, the mother vessel may undergo bridging in
which translumenal bridges, formed between the endothelial cells, serve to divide
the lumen of the mother vessels into smaller channels. Ultimately, these multichannel vessels give rise to new daughter vessels. Finally, during intussusceptive vessel
formation, the mother vessel invaginates to give rise to two vessels.
Characterization of the Angiogenic Tumor Blood Vessels
During normal physiologic conditions, the neonatal vasculature is under the tight
control of two processes: vasculogenesis and angiogenesis. Vasculogenesis, the de
novo generation of blood vessels, is initiated from a precursor stem cell or angioblast, whereas during angiogenesis these newly formed vessels are instructed to
undergo branching and sprouting by signals including VEGF. These fragile, newly
formed blood vessels are reinforced by a cast of supporting cellular (mural cells)
and structural elements including pericytes, smooth muscle cells, and the extracellular matrix, a process which is regulated and coordinated by extracellular signals
including PDGF, EDG1, and S1P1 (61). Finally, during embryonic development,
the vasculature is tailored and remodeled to fit the needs of the organ it supplies by
the localized expression of chemoattractants and repellents such as ephrins and
neuropilins (62–64). Overall, the process results in the formation of highly ordered
vascular beds comprised of arteries, arterioles, capillaries, venules, and veins.
In tumors, however, the organization of angiogenic blood vessels resembles
haphazard arrays of sprouting blood vessels, tumor cells, and pericyte-like cells which
collectively form a vessel that oftentimes lacks an organized extracellular matrix.
The hyperpermeability of these blood vessels, a process which is in part mediated
by VEGF, allows the extravasation of colloidal particles such as colloidal carbon and
colloidal gold, as well as tracer molecules including radioisotopes, fluorescently
labeled dextrans, and albumin (58,65–67). In certain instances, erythrocytes cross
the vascular bed and pool to form vascular lakes of blood cells (60).
Recently, the tumor cells themselves were reported to form the so-called
“tumor blood vessels.” Chang et al. (68) described the presence of “mosaic” blood
vessels in a murine colon tumor model. These mosaic blood vessels consisted of
both tumor and endothelial cells which were distinguished by fluorescence. In these
studies, Chang implanted LS174T murine colon carcinoma tumor cells, which
were stably transfected to express green fluorescence protein (GFP), into SCID mice.
Biological and Engineering Considerations
147
With tumor formation, the intracellular expression of GFP was used to identify the
tumor cells, whereas Cy5 and rhodamine-conjugated anti-CD31/105 systems were
used to identify the endothelial cells. They report that nearly 15% of the vessels
examined did not show any endothelial cell staining, but did show the green fluorescence of the implanted tumor cell line, creating the lumen of the “tumor blood
vessel.” In contrast to the data reported by Monsky et al. (69) (see later), who
reported that the permeability of angiogenic blood vessels was influenced by the
location of tumor cell implantation, the location of tumor cell implantation in the
studies conducted by Chang et al. did not affect the presence or percentage of the
mosaic cells. These mosaic blood vessels are thought to represent one mechanism by
which shedding tumor cells enter the circulation in formation of metastases.
Furthermore, given that mosaic cells, similar to those described above, were also
identified in biopsies of colon cancer patients (68) has significant implications for
angiogenic-targeted cancer therapies, because the main target for these therapies,
the newly formed blood vessels, may not be present in all cases.
Tumor Angiogenesis and Interstitial Fluid Pressure
For some time now the inherent leakiness of the tumor vasculature has been a target
for developing nanoparticle-based, tumor-targeted delivery vectors. Because of
their size, these particle-based therapeutics may passively extravasate the leaky
tumor vasculature to accumulate within solid tumors. Tumor–host interactions are
also shown to control the degree of endothelial cell leakiness. Monsky et al. (69)
demonstrated that the degree of tumor blood vessel porosity is dependent not only
on the tumor type, but also on its location. For example, the inherent vascular permeability of human breast cancer tumors was dependent on the site of implantation.
Although tumors implanted in the cranial wall were highly angiogenic, when
compared with the same tumors implanted in mammary fat pad, they were also less
permeable to fluorescently labeled indicators.
Nevertheless, the formation of an intratumor clot, in the presence of continued angiogenesis, and a late-forming lymphatic system combine to increase the
interstitial fluid pressure (IFP) of solid tumors. This hypertensive state represents a
major obstacle for the penetration of macromolecular- or nanoparticle-based therapies
to solid tumors. To effectively overcome this barrier, nanoparticle-based delivery
systems must contain elements that break down this interstitial pressure gradient.
Described below are two classes of therapeutic compounds shown to disrupt the
IFP gradient.
Antiangiogenic Therapies
One clear choice for reducing the IFP in solid tumors is to block VEGF signaling.
Preclinical and clinical studies with monoclonal antibodies designed to block the
interaction of VEGF with its receptor show that the antibody not only inhibits the
formation of new blood vessels, but also causes the regression of already established vessels, an observation which is consistent with the role that VEGF serves as
a survival factor for newly formed blood vessels (70,71).
More interestingly, treatments with such monoclonal antibodies cause an
apparent reorganization and restructuring of the blood vessels within solid tumors.
This process, termed normalization (70,72), involves a pruning of existing blood
vessels, improved covering of the normalizing blood vessels with pericytes, and
reestablishment of basement membrane. Overall, normalization results in improved
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blood flow in solid tumors, decreases the IFP of tumors (an observation which has
also been demonstrated in colon cancer patients), and establishes a hydrostatic pressure gradient across the vascular wall. The combination of these events improves
the accumulation and efficacy of secondary therapies, such as chemotherapy.
There is, however, one caveat associated with the anti-VEGF-induced normalization of the tumor vasculature and its proposed synergy with chemotherapy.
Apparently, there is a period of time, known as the “normalization window” (71), in
which synergy of anti-VEGF therapy with radiation or chemical therapies is optimal.
This period of time is temporally flanked (i.e., prior to and immediately following
the normalization window) by two time periods in which the combination is less
effective. Not surprisingly, prior to normalization, when solid tumors are supported
by their torturous/leaky circulation vasculature, the synergy of anti-VEGF treatment and chemotherapy proves ineffective. During normalization, the drop in IFP
coupled with the reestablishment of a vascular pressure gradient improves the
efficacy of secondary therapies. Finally, as the anti-VEGF therapy causes the death
of the tumor neovasculature, the efficacy of anti-VEGF and radio or chemotherapies
once again decreases. These data, however, suggest that in order to maximize the
synergy between anti-VEGF and chemotherapy, the normalization window in
cancer patients must be clearly identified to ensure maximal synergy between the
two therapies.
Cytokine-Mediated Reduction of Interstitial Fluid Pressure
TNF was initially described as a serum factor isolated from the blood of endotoxintreated mice, which was capable of inducing hemorrhagic necrosis of solid tumors.
Nevertheless, in the 1980s hopes of using TNF as a cancer therapy were nearly
dashed by the life-threatening toxicities it induced (72), a fate shared by many cytokines showing promise in preclinical tumor models. The major toxicity observed
during the systemic administration of TNF is severe hypotension.
Yet the pioneering works of Lienard and Lejuene (73,74) demonstrated vast
improvements in the therapeutic index of TNF by limiting its delivery to solid
tumors. For TNF, the isolated limb protocol (ILP) demonstrated that the combination of surgically localized delivery of the cytokine with regionally administered
chemotherapies induced sustainable antitumor responses in patients failing traditional standards of care (72–74). By localizing the delivery of TNF to solid tumors,
Lienard and Lejuene harnessed TNF to induce at least one of the possible mechanisms by which the cytokine induces an antitumor response. For example, TNF may
directly cause apoptosis of various tumor cell types (75,76), activate and drive
immune-based antitumor responses (77), and induce vascular leak in tumor blood
vessels (78,79), leading to a destruction of the IFP gradient in solid tumors (80,81).
The clinical observations made on the timing/order of TNF and chemotherapy treatment is consistent with a reduction in the IFP. In order to achieve maximal
antitumor responses, TNF treatment either preceded or was given simultaneously
with chemotherapy. Reversing the order, that is, giving the chemotherapy first
followed by TNF treatment, produced less than optimal results. In effect, ILP with
TNF served to sensitize the tumors to chemotherapy, and in preclinical models, this
observation correlated with improved uptake of chemotherapeutic agents by the
solid tumors (82). In the following section, we will describe our recent efforts to
simulate the ILP using colloidal gold-based nanotherapeutics that sequester not
only TNF, but also a chemotherapeutic (i.e., paclitaxel) in murine tumors.
Biological and Engineering Considerations
149
Barriers Within the Tumor Interstitium: Intratumor Barriers
In order to achieve effective delivery of antineoplastic agents to solid tumors,
nanoparticle delivery systems must effectively evade immune detection by the RES,
find and extravasate the leaky tumor vasculature, and simultaneously possess
methods for reducing the IFP. Yet, even when these prerequisites are met, the intratumor microenvironment presents one additional obstacle that may thwart effective
drug delivery to the tumor interstitium. These nanotherapeutics must overcome the
spatial barrier established by the extracellular matrix of solid tumors. This viscous
matrix consists of a fibrous mesh of proteoglycan-assembled collagen fibers in which
stromal and tumor cells are suspended.
The orientation of collagen type I fibers present in the extracellular matrix may
retard the trafficking and penetration of large macromolecules such as IgG, IgM,
and high-molecular-weight dextran and nanoparticle-delivery systems through the
tumor interstitium. The main obstacle posed by collagen is the spatial orientation of
the individual collagen fibrils and collagen bundles. Depending on the tumor
model, the small space between individual collagen fibers may hinder the movement of particles larger than 40 nm, whereas the collagen bundles may retard the
trafficking of particles that are 75 to 130 nm in diameter. Interestingly, in murine
tumor models, the relative contribution of the collagen mesh appears to be controlled by tumor–host interactions because the density of collagen and the complexity of the mesh are dependent on the site of tumor implantation (83).
Active vs. Passive Tumor Targeting
Passive extravasation through the leaky tumor vasculature is the primary mechanism by which nanoparticle vectors gain access to the tumor interstitium. Upon
extravasation, these nanotherapeutics rely on diffusion, rather than convection, as
the primary mechanism for delivering drugs to cancer cells. In effect, the passive
accumulation of the nanotherapeutics is designed to limit the biodistribution of the
therapeutics.
To further improve the delivery of the therapeutic payload, the nanoparticles
may be engineered to contain tumor-targeting ligands that guide the delivery of
these vectors to tumor, endothelial, and potentially stromal cells. Such targeting
ligands also assist in the uptake of the nanotherapeutics by tumor cells by receptormediated mechanisms including receptor-mediated endocytosis and potolysis.
Examples of tumor-targeting ligands include oligosaccharides, folic acid, EGF, and
TNF, as well as antibody-directed cell surface receptors (84–89).
Tumor Models
One final challenge in developing nanoparticle therapeutics lies in the extrapolation
of preclinical findings, often conducted in rodent tumor models, to the clinical setting. Typically, after showing promise in such preclinical studies, candidate drug
therapies are tested in FDA-guided GLP toxicology studies, and indeed these studies are requisite to ensure patient safety. Yet, testing such therapeutics in patients
with naturally occurring cancer is the critical test. In advance of formal clinical trials,
treating companion animals (i.e., pets) with inoperable/unresponsive tumors on a
compassionate use basis may provide additional insight to data obtained from
animal tumor models. It is estimated that in the U.S.A. alone there are 130 million
cats and dogs, of which 10% to 25% present with a variety of naturally occurring
solid tumors (90). We believe that such terminally ill animals may serve as sentinels
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for the human condition and may provide valuable information regarding drug
performance in human cancer patients.
METALLIC NANOPARTICLE DRUG-DELIVERY SYSTEMS
Although metallic nanoparticles have a long history of use in biology, their application to tumor-targeted drug delivery has only recently been described. For example,
colloidal gold nanoparticles have been used as diagnostic indicators and therapeutics since the early 1950s. Similarly, although the magnetic properties of iron
nanoparticles have made them a valuable tool for magnetic resonance imaging
(MRI), their use in magnetic-targeted drug delivery to solid tumors has only recently
been described. In the following sections, we describe the characterization of these
nanoparticulates, toxicologic considerations, and formulation into tumor-targeted
drug-delivery vectors.
Primer on the Use of Colloidal Gold and Iron Nanoparticles
in Biotechnology
Gold Nanoparticles
In 1857, Michael Faraday (91) when generating multicolored solutions by reacting gold
chloride with sodium citrate, he could not have appreciated the fact that he was
laying the foundation for the field of nanotechnology. Unbeknownst to him what he
actually described was the synthesis of colloidal gold nanoparticles that ranged
from 12 to 100 nm in diameter. Since that time gold nanoparticles have been used to
meet a variety of needs in science and medicine. In the 1950s, the discovery that the
particles bound protein biologics without altering their activity paved the way for
their use in hand-held immunodiagnostics and in histopathology (92). Of particular
importance to the current discussion was the long-standing use of colloidal gold in
the treatment of rheumatoid arthritis as well as the use of radioactive gold nanoparticles to treat liver cancer (see later). More recently, gold nanoparticles have been
assembled into scaffolds for use in DNA diagnostics and biosensors (93).
Iron Nanoparticles
Iron nanoparticles have been used in both diagnostics and therapeutics agents, with
specific applications as contrasting agents for MRI and magnetically targeted drug
delivery. Two types of iron oxide nanoparticles have been used as imaging agents:
superparamagnetic iron oxide (SPIO) and ultra-small superparamagnetic iron oxide
(USPIO) nanoparticles (94). Typically, these nanoparticles are coated with a variety of
stabilizing agents including dextran, albumin, starch, or silicones. The major difference between SIPOs and USPIOs relates to their size and circulatory half-life. Both
particles may be used as contrasting agents to image the gastrointestinal tract, liver,
spleen, and lymph nodes, although the USPIOs may be used to demonstrate blood
pooling in diseases such as brain and myocardial ischemia.
Safety of Metallic Nanoparticle Administration in Man
Unlike biodegradable particles such as liposomes and polymeric-based nanoparticles, metallic nanoparticles are relative newcomers to the field and thus the available toxicology data for each nanoparticle system are limited. From a historical
perspective, the question of colloidal gold toxicity may be gleaned from the use of
radioactive colloidal gold nanoparticles as a cancer therapy in the late 1950s.
Biological and Engineering Considerations
151
Specifically, radioactive colloidal gold nanoparticles were made from Au198 and
used for the treatment of liver cancer and sarcoma (95,96). Safety data revealed that
the observed toxicities of these intravenously administered radioactive nanoparticles were due to radiation exposure. No demonstrable toxicities were noted from
the particles themselves. In other reports, gold was also reported to be inert and
biologically compatible (97). Data from our GLP toxicology study on the safety of
drug stabilized colloidal gold nanoparticles agree with the historical data (Paciotti
et al., unpublished observations). Like gold, iron-based nanoparticles are also biologically inert drug carriers Alexiou (98) and Lubhe (99) showed that dextran-coated
magnetite particles exhibit no discernable toxicity (i.e., LD50) in mice and rats (98,99).
Furthermore, clinical studies on the use of epirubicin demonstrated that ferrofluids
were well tolerated, with toxicity being limited to the active agent, rather than the
iron nanoparticles themselves (99,100).
Metallic Nanoparticle-Binding Chemistry
Inherent in the development of metallic nanoparticle drug vectors is an understanding of the binding mechanisms involved in adsorbing proteins and other drugs to
the surface of metallic nanoparticles. For colloidal gold, proteins and other drugs
bind to the surface of the colloidal gold particle by one of three mechanisms. Two of
these mechanisms, ionic and hydrophobic binding, are relatively weak interactions
that often result in the generation of poor-quality vectors. The third method involves
the formation of a dative/coordinate covalent bond between free sulfhydryl/thiols
of the biomolecule and the gold atoms present on the surface of the particles. Dative
bonds are very stable, possessing the energy equivalence of a covalent bond, and
are only disrupted by strong reducing agents such as dithiothreitol or mercaptoethanol (30,101). Additionally, we recently generated functionalized colloidal gold
nanoparticles containing various surface groups that are useful for covalent drug
immobilization. These functional groups, which include NH2, COOH, SH, and OH,
may be used to link drugs through well-described cross-linking chemistries such as
NHS-based ester formation and EDC-based coupling (Silin et al., unpublished
observations).
For iron-based particles, drug binding primarily occurs through indirect coupling of the drug to the polymer coat surrounding the particle. For example, Allport
and Weissleder (102) attached the HIV tat peptide to the surface of dextran-coated
iron nanoparticles. Alternatively, Jain et al. (103) used a double coating of oleic acid
and pluronic F127 around the central iron nanoparticle to allow doxorubicin to partition in this hydrophobic layer while relying upon the pluronic coating to increase
the hydrophilic nature of the particle (see earlier). One unique aspect of magnetic
particles is that the particles themselves represent a novel therapeutic. For example,
through magnetic excitation, the particles become heated and have been shown to
be useful in thermal ablation of cancer cells (see later).
Preclinical and Clinical Studies on Gold and Iron Nanoparticle
Drug Delivery
Gold Nanoparticles
CYT-6091 is a multivalent drug that is assembled on 26-nm particles of colloidal
gold and designed to actively sequester recombinant human TNF alpha within solid
tumors (30). The drug is manufactured by covalently linking molecules of TNF and
thiol-derivatized polyethylene glycol (PEG-THIOL) onto the surface of the colloidal
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gold nanoparticles (30). Following intravenous administration, CYT-6091 rapidly
accumulates in MC-38 colon carcinoma tumors and shows little to no accumulation
in the livers, spleens (i.e., the RES), or other healthy organs of the animals. The
tumor accumulation was evidenced by a marked change in the color of the tumor as
it acquired the bright red/purple color of the colloidal gold nanoparticles, and was
coincident with the active and tumor-specific sequestration of TNF. Finally, CYT6091 was less toxic and more effective in reducing tumor burden than native TNF
because maximal antitumor responses were achieved at lower doses of the drug.
Our experimental data suggest that the inherent leakiness of the tumor neovasculature facilitates the passive extravasation of CYT-6091 into the solid tumor, limiting
the biodistribution of TNF and avoiding healthy organs and tissues. Once inside the
tumor, each molecule of TNF bound to the surface of PEGylated colloidal gold
nanoparticles may serve one of the two functions. First, as expected from TNF’s
known biological action, CYT-6091 serves as the anticancer therapeutic; second and
more importantly, TNF serves as a tumor-targeting ligand. The above studies were
conducted in the TNF-sensitive colon carcinoma model, MC-38, and thus the
observed antitumor effect may have been due to a direct action/binding of TNF on
these tumor cells. However, in human TNF-insensitive B16/F10 melanoma tumors,
CYT-6091 exhibited differential pharmacodynamics. Although CYT-6091 caused
similar accumulation of TNF in B16/F10 tumors, only transient inhibition of tumor
growth was observed. These data combined with the known cellular heterogeneity
of solid tumors support the development of multifunctional metallic nanotherapeutics that attack solid tumors on multiple levels. An example of such a vector is
presented below.
Iron Nanoparticles
The delivery of the epirubicin-conjugated iron particles was done by intravenous
injection of the nanoparticles into a vein, which was located contralateral to the
tumor. At the same time, a magnetic field, ranging from 0.5 to 0.8 T, was established
and maintained for 45 minutes around the site of the tumor. Preclinical data in
rodent models demonstrated that not only could this strategy concentrate the
nanoparticle drug vectors within the solid tumor, but also significantly improve the
antitumor efficacy of epirubicin treatment. In clinical trials, these studies show
similar findings: the magnetic particles were detected in 50% of the patients treated,
which correlated with the tumor areas showing the coloring of the iron nanoparticles. These data were further confirmed by histological examination of tumor tissue.
Data from preclinical studies revealed that the particles were cleared by the RES
after removal of the magnetic field. Other healthy filtering organs such as the lung
and the kidney did not show the presence of the nanoparticles (98–101).
Metallic Nanoparticle for the Thermal Ablation of Tumors
In 1957, Gilchrist et al. (104) proposed the use of thermal therapy as a treatment for
solid tumors. Data supporting such an approach was later provided by Jordan et al.
(105) and Neilsen et al. (106), showing that tumor cells were sensitive to temperatures above 41°C. The means of treating solid tumors with heat has advanced from
the use of ultrasonic and microwaves to the use of magnetic nanoparticles that heat
in response to alternating magnetic fields. The degree of heating and in turn the
strength of treatment is determined by the intensity of the magnetic field and the
size of the object being magnetized. Babincova et al. (107) demonstrated that the
use of this approach favored the destruction of neoplastic cells over healthy cells.
Biological and Engineering Considerations
153
Two approaches have been tried to improve the efficacy of thermal ablation therapy.
The first modification involves coupling of the thermal therapy with more conventional therapies such as chemo- or radiotherapy. The second method increases the
maximal treatment temperature from 45 to 47 to 55°C. Nevertheless, at these high
temperatures normal cells are affected, a side effect which may be overcome by
direct intratumor injection of the particles (see Ref. 93 for review).
Hirsch et al. (108) recently described a second nanoparticle approach for the
thermal therapy of solid tumors. They describe the synthesis of a gold/silica
nanoshell comprised of an aminated silica core particle which was studded with
ultra-small (1–3 nm) gold particles. These small gold nanoparticles are bound to the
surface of the silica core through their interaction with the amine group. The ultrasmall gold particles were subsequently used as nucleation sites upon which additional gold was reduced to form the encased gold nanoshell. This process allowed
for the synthesis of gold-covered particles with a tunable plasmon band in the nearinfrared region. Exposing these particles to a light source (i.e., diode laser) with a
wavelength of 820 nm causes the electrons present on the gold surface (i.e., plasmons) to become excited, resulting in particle heating. Interstitial injection of these
particles near the tumor followed by laser light excitation caused a significant reduction of transmissible venerea tumors (TVT) tumors growing in SCID mice.
Multifunctional Nanotherapeutics
Our previous discussion has centered on the use of metallic nanoparticles to target
delivery of singular therapies, whether chemical, biological, or physical (e.g., heat),
to solid tumors. Nevertheless, in recent years, it has become increasingly clear that
solid tumors may be viewed as unique organs that are constantly adapting to their
changing environment by altering their phenotype. As previously discussed, solid
tumors represent a collage of cells, which for the most part work in concert to ensure
continued tumor growth and metastasis. It seems unlikely that any given single
agent, regardless of its ability to perform in experimental tumor models, will yield
the exacting results in cancer patients.
Thus, one of the final challenges that metallic nanoparticle drug-delivery systems face is to attack the heterogeneous population of tumor cells present within a
solid tumor by delivering multiple therapeutics on a single nanoparticle.
Combinational cancer treatments have been a hallmark of recent clinical strategies
to treat cancer. One example, previously discussed, the isolated limb perfusion,
demonstrates dramatic efficacy of the combination of TNF and chemotherapy to
treat refractory cancer patients (72–74). More recently, a similar combination, Avastin
and 5-FU, has demonstrated significant benefit in colorectal cancer patients (70,71).
Our own experience with CYT-6091 also supports the development of multifunctional metallic nanotherapies. We observed that although CYT-6091 sequestered TNF
within B16/F10 tumors, it produced only marginal antitumor effects in this model.
Metallic nanoparticles may deliver these combination therapies either spatially, on the same particle, or temporally, over the course of the disease. Our initial
efforts to develop such a therapeutic centered on developing a colloidal gold
nanoparticle vector that simulates the ILP on a single nanoparticle. Recall that ILP
involves the surgical isolation of the blood supply of a solid tumor to allow for the
regional delivery of TNF and a chemotherapeutic, avoiding systemic toxicities.
To accomplish this goal required the binding of a tumor-targeting ligand, a
chemotherapeutic, and an RES-avoiding molecule onto the same 25-nm particle of
colloidal gold. Specifically, this nanoparticle, termed CYT-21001, uses TNF as the
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tumor-targeting ligand, a thiolated paclitaxel analog as the therapeutic, and PEGTHIOL as an RES-avoidance molecule. To date, we have observed that CYT-21001
delivers significantly more TNF and paclitaxel to solid tumors growing in mice,
when compared to native TNF/paclitaxel. Furthermore, we observed that unlike
CYT-6091, which only caused transient tumor inhibition, CYT-21001 induced significant antitumor responses in this TNF-insensitive tumor model (Paciotti et al.,
unpublished observations).
SUMMARY
Depicted in Figure 1 is a simplified blueprint for developing multifunctional metallic
nanotherapeutics. On the basis of the current knowledge of the biology of tumor
neovascularization as briefly summarized above, there is a rational strategy for
using metallic particles to both target the tumor and to deliver well-known, potent,
but potentially toxic cancer therapies to a broad spectrum of solid tumors.
Additionally, as new therapies are discovered, the versatility of the metallic nanoparticle surface may allow for the rapid development and evaluation of new vectors in
both preclinical and clinical settings. With the development of the first generation of
metallic nanoparticle cancer therapies almost ready for clinical trials, the potential
promise of these new cancer therapeutics may be realized in years not decades.
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11
Biological Requirements for
Nanotherapeutic Applications
Joseph F. Chiang
Department of Chemistry and Biochemistry, State University of New York at
Oneonta, Oneonta, New York, U.S.A., and Department of Chemistry,
Tsinghua University, Beijing, China
INTRODUCTION
In order to discuss the biological requirements for nanotherapeutic application, a brief
discussion of cellular and tissue structure is necessary to provide some basic background for readers. Understanding the structure and the function of living cells and
their interactions with their environments is essential to understand the “biological
requirement” and the interactions of nanotherapeutic devices with living systems.
Cells are alive and are the smallest units that exhibit the property of life. They can
be cultured in vitro and could die by themselves. The structure of cells is a very
complex and well-organized system. Cells perform many functions in living things.
The following are functions of cells:
1.
2.
3.
4.
self-replications and regulations,
taking energy from the metabolites,
carrying out mechanical and translocation works, and
producing metabolic functions and chemical reactions using enzymes.
There are two classes of cells: prokaryotic and eukaryotic. They have different sizes
and different internal structures. The former, such as bacteria, is simpler, and the
latter, such as fungi, plants, and animals, is more complex in structure. The contents
of a cell are called “protoplasm.” They can be further divided into cytoplasm (all
the protoplasm except the contents of the nucleus) and nucleoplasm (all of the
materials, plasma and DNA, etc. within the nucleus).
STRUCTURE OF CELLS
Every cell has three major components: plasma (cell) membrane, cytoplasm, and
nucleus (Fig. 1).
Plasma (Cell) Membrane
Plasma membrane is a semipermeable outer boundary of cells. It is a bilayer chain
of molecules (phospholipids). The membrane has a hydrophobic tail pointing
inward, and a hydrophilic head which is in contact with the surrounding environments. Cholesterol is also a component of the membrane in the hydrophobic part of
the membrane. Plasma membrane controls the transport of nutrients, water, and
ions, such as Na+, K+, Ca2+, Mg2+, and so on.
Cytoplasm
All parts are within the cell membrane. The fluids contain salts, sugars, lipids,
vitamins, nucleotides, amino acids, RNA, and proteins. Cell metabolism, replication,
159
160
FIGURE 1
Chiang
(See color insert.) Structure of a typical cell.
and growth all occur within the cytoplasm. Cytosol is a part of cytoplasm, which
refers to a protein-rich environment. Cytoskeleton is the network fibers of proteins,
which maintain the shape of the cell. Ribosomes are the sites for protein synthesis
and consist of ribosomal RNA (rRNA) and structural proteins. Mitochondria,
lysosomes, peroxisomes, and chloroplast (in plants only) are also located in the
cytoplasm. Peroxisomes are membrane-bound vesicles containing enzymes to regulate hydrogen peroxide in the cells. Endoplasmic reticulum is the interconnected
membrane function for protein synthesis and transport. Rough endoplasmic reticulum (Rough ER) is connected to the nuclear envelop. The smooth ER is involved in
transport and other cell functions. The Golgi body functions as an intracellular tool
for sorting new proteins made from Rough ER. Lysosomes are used in phagocytosis
in which foreign materials are brought into the cell for breakdown of the materials.
Mitochondria, (organelles in globular shape) have two membranes. They serve as
sites of energy release and adenosine triphosphate (ATP) formation. Organisms in
storage areas use vacuoles, single-membrane organelles.
Nucleus
The nucleus contains chromatin, DNA–chromosomes, nucleoli, and nuclear proteins. The nucleolus is the site for rRNA gene transcriptions. The nuclear envelope
is used to separate cytoplasm and nucleus. It is a structural frame for the nucleus.
The nuclear membrane acts as a barrier and prevents free passage of molecules
between nucleus and cytoplasm. Nuclear pore complexes regulate the exchange
of molecules between nucleus and cytoplasm. Nuclear membranes are associated
with Rough ER. Ribosomes are bounded to Rough ER.
THE CHEMICAL COMPOSITIONS OF CELLS
Cells are made of water, organic and inorganic ions, and organic molecules. Water
plays an important role in cells. It is a polar molecule with a dipole moment of
1.94 Da. Water has the capability to form hydrogen bonding due to the polarity
Biological Requirements for Nanotherapeutic Applications
161
property, and also interacts with cations and anions. Many organic molecules are
nonpolar and are water-insoluble in an aqueous environment. The former is called
“hydrophilic,” whereas the latter is called “hydrophobic.” Inorganic ions in cells
containing sodium ion (Na+), potassium ion (K+), calcium ion (Ca2+), magnesium
ion (Mg2+), monohydrophosphate ion (HPO42), chloride ion (Cl−), and bicarbonate
ion (HCO3) play critical roles in cell functions. The organic molecules in cells can be
classified as small organic molecules and macromolecules. Lipids belong to the
small organic molecules category, whereas carbohydrates (the polysaccharides),
proteins, and nucleic acids belong to the macromolecules category.
Small Organic Molecules
Lipids
Lipids include triacylglycerol, phosphoglyceride, glycolipid, steroid, wax, terpene,
and prostaglandin and are nonpolar molecules which are soluble in organic solvents. The simplest lipids are fatty acids consisting of long hydrocarbon chains of
16 to 18 carbons with a carboxylic group (–COO). The other end contains nonpolar
C—H bonds which will not interact with water. There are two types of fatty acids:
saturated and unsaturated. Unsaturated fatty acids have one or more double C—
—C
bonds. Phospholipids are the principal components of cell membranes. Two fatty
acids and a phosphate are combined with glycerol to form phosphoglycerides.
Triacylglycerols, or fats, have three fatty acids bound to a glycerol molecule. Energy
sources are stored in fats. Phospholipids are amphipathic molecules, one end is
water-soluble and the other end is water-insoluble. Another type of phospholipids is sphingomelin, which is the only nonglycerol phospholipid in the cell
membrane.
Macromolecules
Carbohydrates
Simple sugars and polysaccharides are carbohydrates. Simple sugars, such as
glucose, serve as major nutrients for cells. The chemical reactions of carbohydrates
will provide energy for cells and also provide sources for the synthesis of other
cellular products. Polysaccharides are also used for protein transport to and association with other parts of the cellular system. Simple sugar or monosaccharide has
the experimental formula (CH2O)n. Glucose, one of the simple sugars in which
with n = 6, is the simplest sugar, (CH2O)6, or C6H12O6. Oligosaccharides are formed
by condensation of several simple sugars, whereas polysaccharides are formed with
a large numbers of simple sugars. Glycogen is a polysaccharide in animals.
Oligosaccharides and polysaccharides are also used for chemical interaction
between cells.
Nucleic Acids
Nucleic acids are genetic information materials in cells. DNA (deoxyribonucleic
acid) is the genetic material in the nucleus. RNA (ribonucleic acid) is responsible for
cellular activities in the cell. There are several types of RNAs. Messenger RNA
(mRNA) serves as a template for protein synthesis. It carries genetic information to
ribosomes from DNA.
Transfer RNA (tRNA) and ribosomal RNA are also used in protein synthesis.
Polymerization of nucleotides, a nucleic acid base attached to a sugar, and phosphate forms DNA and RNA. ATP, a type of nucleotide, is a chemical energy source
162
Chiang
in cells. A detailed description of DNA, RNA, and related subjects can be found in
any cell and molecular biology text. Nucleic acids act in many cellular processes.
They can produce their replications, direct protein synthesis, and carry out informational transfer.
Proteins
Proteins are biopolymers consisting of many different amino acids. Amino acids
have a general structural formulas of R-C(H)(NH3)+(COO). Condensation of two
amino acids between the carboxyl group of one and the amino group of the other
forms a peptide bond. The side chains are either straight carbon chains or carbon
rings. There are approximately 20 such amino acids. The characteristics of proteins
are based on the properties of the amino acids. Some amino acids have nonpolar
side chains, such as glycine, alanine, valine, leucine, isoleucine, proline, cysteine,
methionine, phenylalanine, and tryptophan. Thus, the side chain of these amino
acids will not interact with water or polar molecules. Amino acids serine, threonine,
tyrosine, asparagine, and glutamine have polar side chains of either OH group, or
amide group (O—
—C–NH2). They are hydrophilic and will form hydrogen bonding
with water. Lysine and arginine, with charged groups in the side chain, are
hydrophilic and will interact with water. Histidine, acting as a positive charged or
neutral amino acid based on the pH value, has the ability to produce H+ for enzymatic catalysis. Both aspartic acid and glutamic acid have a carboxyl group at the
end of the side chain. They are very hydrophilic. In forming proteins, this end is
located on the protein surface. Amino acid linked by peptide bonds is called protein. The properties of proteins are based on the sequence of amino acids. The properties of proteins also depend on the conformation of the proteins. Many proteins
perform enzyme catalytic reactions within the cells. A catalyst can either increase or
slow down the reaction rate without consuming the catalyst. Such a catalyst is called
an “enzyme” in biological reaction systems. The catalytic action of enzymes usually
increases the rate of reaction. A basic knowledge of chemical kinetics is essential to
understanding any enzymatic process.
DEVELOPMENT OF NANOTECHNOLOGY
Nanotechnology had a fast pace of development for the past decade. It has been
used to rearrange molecules so that every atom can be placed in the most efficient
way. Another term for such arrangement is “molecular nanotechnology” or molecular manufacturing. Thus, it can be used to construct shapes, machines, and products at the atomic level. Nanotechnology can also be defined as the application of
science that deals with materials in 100-nm size, but it is not a miniaturization. At
this range, materials produced will exhibit new properties we are looking for. In
this range, the surface area increases drastically, which will exhibit new chemical
and physical properties. Nanotechnology also provides applications in energy
storage, energy production, agriculture, air pollution, nanoelectronics, and healthcare. In the healthcare application, use of nanotechnology as tools to manipulate
biomolecules to regulate life and death, illness, and health will be the main goal to
achieve. One of the examples is the use of molecular diagnostics, which will enable
selection of the most efficient treatment for each individual. Nanoparticles are
thought to have potential as novel intravascular probes for diagnostics (e.g., imaging) and therapeutic purposes (e.g., drug delivery). The critical requirement is the
Biological Requirements for Nanotherapeutic Applications
163
ability to target specific tissues and cell types and escape from biological particulate filters. This is the so-called “reticuloendothelial system.” The advantage of
nano-technology over microsystems is that nanotherapeutics has higher intracellular uptakes, allowing drug release in different cellular compartments such as cytoplasm and nucleus. It can also be conjugated with a ligand to favor a targeted
therapeutic approach. Practically, drug load is a science based on size and structure
of device. Examples are the use of nanoparticles as controlled drug delivery for
cancer treatment. A combination of micro- and nanoparticles can also be used for
drug-delivery systems, such as micro/nanospheres, micro/nanocapsules, and
liposomes. Such combinations will differ in structure and biopharmaceutical properties for different therapeutic uses. Liposomes discovered by Baughman (1) are
the smallest artificial vesicles of spherical shape that can be produced from natural
untoxic phospholipids and cholesterols. Liposomes can be used as drug carriers
and loaded with a variety of molecules, such as small drug molecules, proteins,
nucleotides, plasmids, and carriers for lipophilic-antitumor drug N-octadecyl-1-βarabinofuranosyl cystine (2), azidothymidine, and dideoxycytidine. There are other
drug-delivery system approaches in addition to liposomes, such as polymer microcapsules, microspheres, polymer conjugates, and nanoparticles as mentioned previously. The most fundamental requirement for nanotherapeutic devices is to
deliver any drug at the right time and in the target where it is needed and at the
level that is required. Development of drug-delivery systems is also aiming at the
therapeutic and toxicological properties of existing chemotherapies. Nanoporous
membranes with 7 to 9 nm pores offer size-based exclusion and controlled diffusion of molecule drugs. Common technology for drug administration and devices
include: subcutaneous, implants, surgical, oral, intravenous, and possible pulmonary inhalation, whereas the characteristics of common routes of drug administration are shown in Table 1.
Research and technology development at nanoscale provides a fundamental understanding of materials at nanoscale in order to create devices and
systems in medicine–nanomedicine. Recently, a professional organization “American
Academy of Nanomedicine (AANM)” was established in Baltimore, Maryland, on
August 15, 2005. AANM is devoted to the study of nanomedicine and nanotherapeutic devices.
FABRICATION OF NANODEVICES
It is worthy to review and introduce the two approaches in fabrication of nanodevices and any nanoparticles at this point. The first one is the top-down approach
and the second one is called bottom-up. The top-down approach involves the reduction of materials from bulk size to micro- and to nanoscale. For a three-dimensional
case, if two dimensions are kept in macroscale, but the third one is reduced to
nanoscale, the structure is known as quantum well. If one dimension is kept in macroscale and the other two are reduced to nanoscale, this is called nanowire. If all
three dimensions are reduced to nanoscale, it is referred to as quantum dot. The
top-down approach begins with a reduction of macroscale substance to nanoscale
product. Lithographic technique as used in microchip fabrication is an example.
The standard process is listed below. An n-type silicon wafer doped with p-type
impurities can be used to create the drain and source of a transistor. A stepwise
illustration is shown as follows (Fig. 2).
164
Chiang
TABLE 1 Characteristics of Common Routes of Drug Administration
Route
Absorption pattern
Special utility
Intravenous
Absorption
circumvented
Potentially immediate effects
Valuable for emergency
use
Permits titration of dosage
Suitable for large volumes
and for irritating
substances, if diluted
Subcutaneous
Prompt, from
aqueous solution
Suitable for some
insoluble suspensions
and for implantation of
solid pellets
Intramuscular
Slow and sustained,
from repository
preparations
Prompt, from
aqueous solution
Variable; depends
upon many factors
Increased risk of adverse
effects
Must inject solutions
slowly, as a rule
Not suitable for oily
solutions or insoluble
substances
Not suitable for large
volumes
Possible slough from
irritating substances
Suitable for moderate
volumes oily vehicles,
and some irritating
substances
Slow and sustained,
from repository
preparations
Oral ingestion
Limitations and precautions
Most convenient, safe,
and economical
Precluded during anticoagulant medication
May interfere with interpretation of certain diagnostic tests (e.g., creatine
phosphokinase)
Requires patient cooperation. Absorption potentially erratic and
incomplete for drugs that
are poorly soluble and
absorbed slowly
1. A silicon wafer is cut from monocrystal silicon (1 1 1) direction.
2. A layer of silicon oxide is deposited on n-silicon.
3. A photosensitive emulsion is doped on SiO2. The etched area is removed by
chemical process.
4. A layer of SiO2 is formed by oxidation process on top of the silicon. A p-type of
impurities is used to cover the area to be metalized.
5. Oxidation is processed again.
6. Removal of the second oxide is performed.
7. Oxidation again on the newly grown areas.
8. Removal of the third layer again.
9. The final metallization will produce the final transistor.
There are many bottom-up processes in use for fabrication of nanomaterials. The
following are just a few and include sol–gel process, chemical vapor deposition
(CVD), laser pyrolysis, molecular condensation, hydrothermal process, and many
other newly developing techniques. A particular method developed by Xu (3) used
water-in-oil microemulsion and hydrothermal microemulsion to prepare CdS
nanocrystal. The group also prepared NiO nanoring and monodispersed monoclinic zirconia (m-ZrO2) via hydrothermal method in ethanol–water system (4). A
new tool: “Dip-pen nanolithography” was developed by Mirkin and coworkers (5).
This new method can be used for direct-write scanning probe-based lithography
165
Biological Requirements for Nanotherapeutic Applications
N-Type Silicon
N
(1) Substrate material - Silicon
(2) The first oxidation deposit thick SiO2
layer
P
P
N
N
(3) First oxide removal by etching
(4) Doping is performed by thermal
deposit
P
P
P
N
(5) Second thick oxidation
P
(6) Second oxide removal
P
P
N
P
N
(8) Third oxide removal – exposes
source and drain
(7) Gate oxidation
P
P
N
N
P
(9) Final metallization resulted in
finished transistor
P
N
P
(10) Tunnels with crossover
FIGURE 2 Standard method for fabrication of p-channel metal oxide semiconductor field effect
transistor (MOSFET).
which uses an atomic force microscopy (AFM) tip to deliver chemical reagents on a
target substrate.
AFM is a scanning technique to produce high resolution 3-D image of less
than 1 nN sample surface. It can be used to measure the force between an AFM tip
and the sample surface (conductor or insulator). The small forces are measured by
measuring the motion of a very flexible cantilever beam having an extra small mass.
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Chiang
In AFM, the forces between the tip and sample are detected rather than tunneling
current. It is a good tool to pattern many organic molecules (6 –11), organic (12–14)
and biological (15,16) polymers, and colloidal particles (17–19) to metal ions (20–22).
The sol–gel method is used to produce colloidal nanoparticles from liquid
phase, especially for oxide nanoparticles (23–25). It is a hydrolysis and condensation
of metal alkoxides process. In hydrolysis, precipitation from solution forms insoluble hydroxides. The hydroxides are then converted to oxides by dehydration. CVD
method has been used to produce single-walled carbon nanotubes either at raised
temperature (thermal CVD), or with plasma-enhanced chemical vapor deposition
(PECVD). Atomic or molecular condensation is used for producing metallic
nanoparticles. A raw material is heated in a vacuum to vaporize the material. Rapid
coating of the vapor-phase metal in noble gas results in the formation of metal
nanoparticles. Oxide nanoparticles will form by modifying this method using
oxygen instead of noble gas. Nanoparticles of many materials, such as organic,
biological molecules, metals, and inorganic oxides, can be produced by chemical
self-assembly technique. This technique can be used for attaching molecules to a
specific surface of a substrate. Many research groups use self-assembly as a method.
One of them is to fabricate alkylsiloxanes on silicone dioxide (26). Another example
is alkanethiolates on gold (27).
Nanotechnology can be used for therapeutics due to the compatible sizes of
proteins and nanoscale particles. For example, hemoglobin has a size of 4.5 × 7 nm2
dimension; lipoprotein, 20 nm2 in spherical shape; α-globulin, 4.3 × 26 nm2; and
fibrinogen, 4 × 76 nm2. These are in the nanometer range.
NANOTHERAPEUTIC DEVICES
The following is a list of systems in micro- and nanotherapeutic delivery devices:
1.
2.
3.
4.
5.
oral drug delivery
injection-based drug delivery
transdermal drug delivery
bone marrow infusion
organ-system-specific drug delivery
a. pulmonary drug delivery
b. nasal delivery to central nervous system
c. cardiovascular system (CV)
d. genito-urinary tract
e. gastrointestinal tract
f. ocular drug delivery
6. controlled release system
7. novel packaging and formulation
a. fast dissolving system
b. chewable tablets
c. solubility enhancement
8. target drug delivery
a. polymer and collagen system
b. particle-based system
i) therapeutic monoclonal antibodies
ii) liposomes
iii) microparticles
167
Biological Requirements for Nanotherapeutic Applications
c. modified blood cells
d. nanoparticles
e. viral-assisted intracellular gene delivery, and
f. nonviral intracellular gene delivery
9. implant drug-delivery system
Some of the above delivery systems are obvious from the title. Some do need certain explanations and examples. The targeted drug delivery is just one of the
cases. Poly(amino) amine dendrimers (28,29) are one of the polymer systems.
Block copolymer has potential to encapsulating large numbers of guest molecules
within the cavity for therapeutic applications. Nanoparticles engineered with
specific binding power can be suspended in body fluids or even injected into
the circulation system. Nanoparticles can be used both quantitatively and
qualitatively for in vivo detection of tumor cells. At the present time, most nanotherapeutic devices and drug-delivery systems are concentrated on oncology
for detection and curing tumor cells. In particular, nanotherapeutic systems
can reach tumor cells much easier than other cells. Tumor tissue is usually different from normal tissue. This phenomenon provides the so-called “enhanced
permeability and retention (EPR) effect” (30). Antitumor agent delivery with
EPR will be more effective. The application of nanoparticles to CVs has two
purposes: detection and targeted drug delivery. Recently, a report by Corie Lok
in the May 2005 Issue of the Technology Review described a new diagnostic tool
called “metabolomics” which can be used to detect diseases at an earlier stage
and can cure the diseases as well. The principles involved in metabolomics are
the analysis of a large number of small molecules, such as sugar and fats to reveal
abnormal behaviors. This will lead to metabolic fingerprints to give an earlier
and more accurate diagnosis for diseases. Implantable devices consisting of
sealed arrays of reservoirs can be used for in vivo and in vitro drug analysis and
delivery. These devices are filled with chemicals that can be released on demand
and can check the efficacy of drug released over a long period of time. Implantable
brain probes have been studied for neural activities. The results can be used in
the treatment of various mental and other brain disorders. Another example is
the nanorobots injected into the human body to work in patients’ bloodstreams.
In future, medical nanodevices can routinely be implanted or even injected into
the bloodstream to monitor wellness and to participate in the repair of the system
that deviates from the normal state. Implantable sensors could be used in the
detection of glucose for diabetes. Regulation of glucose metabolism in the human
body by insulin can be achieved by the design described by Gouch (31). The
device can be programmed to warn of hypoglycermia (low blood glucose) or
hyperglycermia (high blood glucose). Such indication can guide diabetics to
adjust insulin injection or even could be coupled to an implanted pump to deliver
insulin correctly. The principle of operation is based on glucose oxidase, which
catalyzes the reaction:
Glucose + O2 + Η2O
→ gluconic acid + H O
2
2
Oxygen is used for the reaction. The unreacted oxygen will be detected in the form
of electric current. Thus, the amount can be calibrated in relation to the current of a
system without the presence of the enzyme, the oxidase. Another biosensor developed recently is to measure the level of H2O2 using Pt electrode (32). One more
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Chiang
method mentioned here is the measurement of glucose concentration for diabetes
by direct electrochemistry of glucose oxidase absorbed on a colloidal gold-modified
carbon paste electrode (33). Castillo et al. (34) have measured glucose concentration
based on enzyme-based amperometric biosensor.
The search for nanotherapeutic applications is mainly due to the limitation of
current therapeutics, especially chemotherapies and radiotherapies. These therapeutics have a high degree of toxicity. They also produce many invasive side effects;
not only do they give unfavorable therapeutic effects, but they also cause damages
during administrations of such drugs or radiation treatments. There are many
advantages of nanotherapeutic systems over the traditional drug administration.
For any drug delivery, the penetrations of strong acids/bases depend on the permeability of cell membrane. Penetration of drug with nanodevices into mitochondria
follows the same principle as cell membrane. Also pH value between intracellular
and extracellular fluids is small (pH of 7.0–7.4). The gradient of drug across the
plasma membrane is small. In general, weak bases are slightly concentrated inside
cells, whereas weak acids are less concentrated inside cells. Lowering pH value, or
increasing the concentration of H+ ion of the extracellular fluids, will cause an
increase in the intracellular concentration.
REQUIREMENTS FOR NANOTHERAPEUTIC APPLICATIONS
1. Transit and penetration properties of nanoscale particles: Nanoparticles of size
less than 20 nm can transit through blood vessel wall. They can also penetrate
blood–brain barrier or stomach epithelium.
2. Interaction: the nanoscale size of nanoparticles can interact with biomolecules
on the cell surfaces, but will not change the biological properties of the molecules.
They also have the ability to interact with receptors, nucleic acids, and proteins
at the molecular scale.
3. Intracellular imaging: Nanoparticles, such as quantum dots, can be used for
intracellular imaging as they can accommodate large numbers of atoms, or
molecules of imaging agent, such as gadolinium, to increase the sensitivity for
detection.
4. Surface chemistry: modifying or altering the surface of nanoparticles to form
covalency with chemical and biological species results in the increase of the
solubility and biocompatibility. For example, attaching a hydrophilic polymer to
the nanoparticle surface will increase hydration. As absorption depends on
solubility, it eventually increases the absorption. They can also be used to
encapsule insoluble compounds.
5. Coating and uncoating: nanoparticles coated with hydrophilic polymers have
longer half-life; uncoated nanoparticles used in intravenous injection can be
cleared from bloodstream by reticuloendothelial system.
6. Ionic attachments: Attaching the surface with ionic cation or anion species can
influence the biocompatibility of nanoparticles and the ability to traverse
biological barriers. An example is dendrimers with amine, a cationic group on the
surface was more cytotoxic than carboxylic-terminated dendrimers.
7. Liposomes used as drug-delivery vehicles. They were developed in early 1960s.
Liposomes are vesicles enclosed by lipid bilayer from self-assembly process of
amphiphilic molecules. They have been used to study membranes, also as
vesicles for controlled drug delivery, catalysis, nonviral gene delivery, and
diagnostic devices. The disadvantage of liposomes application is due to the
169
Biological Requirements for Nanotherapeutic Applications
TABLE 2 Applications of Nanoparticles in Therapeutics
Nanoparticle
Descriptions
Nanospheres
Nanocapsules
Micelles
Drug is uniformly dispersed
Drug is enclosed by polymer membrane in a vesicular system
Amphiphilic blood-copolymer that can self-associate in aqueous
solution
Ceramic
Nanosphere of 17Y2O3–19Al2O3–64SiO2 (mol%) composition, 20–
30 µm diameter for targeted radiotherapy of liver cancer
nanoparticles
(nonradioactive 89Y can be activated by neutron bombardment to
90Y, a µ-emitter (t
1/2 = 64.1 hr)
Liposomes
Artificial spherical vesicles from natural phospholipids and
cholesterol
Dendrimers
Micromolecules consisting of a number of branches around the
inner core
TNT system
Developed by Triton BioSystem with the U.S. Army Lab attacking
cancer in three steps: (i ) patient receiving a single infusion
containing trillions of magnetic biosphere, bound to vantibodies;
(ii ) the biosphere will seek and attach to cancer cells in bloodstream; (iii ) switch on the magnetic field in the cancer region will
cause the biospheres to heat up to kill the cancer cells in minutes
References
(36,37)
(38)
(39,40)
(41)
(43)
(44)
(45)
Abbreviation: TNT, targeted nano therapeutics.
TABLE 3 Some Nanoparticles Used in Biological Research
Nanoparticle
Dendrimers
Application
Targeting of cancer cells, drug delivery, imaging,
boron neutron capture therapy
Ceramic nanoparticle
Passive targeting of cancer cells
Lipid-encapsulated perfluorocar- Passive targeting of cancer cells
bon nanoemulsions
Magnetic nanoparticles
Specific targeting of cancer cells, tissue imaging
LH–RH-targeted silica-coated
Specific targeting of cancer cells
lipid micelles
Thiamine-targeted nanoparticles Directed transfer across Caco-2 cells
Liposomes
Specific targeting of cancer cells, gene therapy,
drug delivery
Nanoparticle-aptamer
Targeting of prostate cancer cells
bioconjugate
Antiangiogenesis therapy
Anti-Flk antibody-coated 90Y
nanoparticles
Gold nanoshells
Tissue imaging, thermal ablative cancer therapy
Anti-HER2 antibody-targeted
Breast cancer therapy
gold/silicon nanoparticles
CLIO paramagnetic
Imaging of migrating cells
nanoparticles
Quantum dots
Tissue imaging
Silicon-based nanowires
Real-time detection and titration of antibodies,
virus detection, chip-based biosensors
Carbon nanotubes
Electronic biosensors
Transfersomes
Noninvasive vaccine delivery, drug delivery
Abbreviation: CLIO, cross linked iron oxide.
References
(44,46)
(47)
(48)
(49,50)
(49)
(51)
(52–54)
(55)
(56)
(57–59)
(59)
(60)
(61–63)
(64–66)
(67)
(68,69)
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Chiang
biological instability, so they have a short lifetime. This has been improved from
research on the synthesis of the polymeric nanocontainers.
8. Targeted drug delivery of nanoparticles can be achieved by binding monoclonal
antibodies on the surfaces of the targeted cells.
Various nanoparticles have been tried in imaging tumors in animal and human trials.
Samples are shown in Table 2 from McNeil (35). Studies of biological effects of chemical and medical devices on immunological, inflammatory, and proliferate effects of
nanomaterials, and the interactions of nanomaterials as applied in medical devices
and toxicological risks have been carried out at various laboratories (Table 3).
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12
Role of Nanobiotechnology in the
Development of Nanomedicine
K. K. Jain
Jain PharmaBiotech, Basel, Switzerland
INTRODUCTION
Given the inherent nanoscale functional components of living cells, it was inevitable
that nanotechnology would be applied in biotechnology giving rise to the term
“nanobiotechnology,” which will be used in this chapter and indicates biotechnology as the main field. A less recognized and less frequently used term, almost
synonymously, is “bionanotechnology,” which implies application in life sciences of
nanotechnology as the main discipline. An up-to-date description of nanobiotechnologies and their applications in healthcare are given in a special report on this
topic (1). The topic of this book is the nanoparticulate drug-delivery systems. This
chapter will provide an integrated overview of application of nanobiotechnologybased molecular diagnostics, drug discovery, and drug delivery in the development
of nanomedicine with the relationships as shown in Figure 1.
ROLE OF NANOBIOTECHNOLOGY IN MOLECULAR DIAGNOSTICS
Nanomolecular diagnostics is the use of nanobiotechnology in molecular diagnostics
and can be termed “nanodiagnostics” (2). In contrast to drug delivery which uses
mainly nanoparticles, nanodiagnostics uses both particulate and nonparticulate
technologies, which are described in detail in a book on this topic (3). Some examples
of the use of nanoparticles for molecular diagnosis are given here and described here
as they can be combined with drug delivery and therapeutics.
Nanoparticles for Molecular Diagnostics
Nanoparticles that are commonly used for diagnostics are
■
■
■
■
■
gold particles
magnetic and supramagnetic nanoparticles
quantum dot (QD) technology
nanoparticle probes
DNA–protein and nanoparticle conjugates
Gold Nanoparticles
Bits of DNA and Raman-active dyes can be attached to gold particles no larger than
13 nm in diameter. The gold nanoparticles assemble onto a sensor surface only in
the presence of a complementary target. If a patterned sensor surface of multiple
DNA strands is used, the technique can detect millions of different DNA sequences
simultaneously. Nanoparticle-based DNA detection systems are 10 times more
sensitive and 100,000 times more specific than current genomic detection systems.
ClearRead® (Nanosphere, Inc.), a nanoparticle technology, enables microarray-based
173
174
Jain
BIOTECHNOLOGY
NANOTECHNOLOGY
DRUG DISCOVERY
LIFE SCIENCES
NANOBIOTECHNOLOGY
DRUG DELIVERY
PHARMACEUTICAL
DEVELOPMENT
NANOARRAYS
MOLECULAR
DIAGNOSTICS
NANOMEDICINE
© Jain PharmaBiotech
FIGURE 1
Integration of nanobiotechnologies for the development of nanomedicine.
multiplex single nucleotide protein (SNP) genotyping of human genomic DNA
without the need for target amplification (4). This direct SNP genotyping method
requires no enzymes and relies on the high sensitivity of the gold nanoparticle
probes. A “spot-and-read” colorimetric detection method for identifying nucleic
acid sequences is based on the distance-dependent optical properties of gold nanoparticles without the need for conventional signal or target amplification (5).
Quantum Dots
There is considerable interest in the use of QDs as inorganic fluorophores, owing to
the fact that they offer significant advantages over conventionally used fluorescent
markers. For example, QDs have fairly broad excitation spectra − from ultraviolet to
red − that can be tuned depending on their size and composition. Potential applications of QDs in molecular diagnostics are
■
■
■
■
■
■
■
cancer
genotyping
whole blood assays
multiplexed diagnostics
DNA mapping
immunoassays and antibody tagging
detection of pathogenic microorganisms
Magnetic Nanoparticles
Magnetic nanoparticles are a powerful and versatile diagnostic tool in biology and
medicine. It is possible to incorporate sufficient amounts of superparamagnetic iron
oxide nanoparticles into cells, enabling their detection in vivo using magnetic resonance imaging (MRI) (6). Bound to a suitable antibody, magnetic nanoparticles are
used to label specific molecules, structures, or microorganisms.
Role of Nanobiotechnology in the Development of Nanomedicine
175
ROLE OF NANOBIOTECHNOLOGY IN DRUG DISCOVERY
The postgenomic era is revolutionizing the drug-discovery process. The new challenges in the identification of therapeutic targets require efficient and cost-effective
tools. Label-free detection systems use proteins or ligands coupled to materials the
physical properties of which are measurably modified following specific interactions.
Among the label-free systems currently available, the use of metal nanoparticles offers
enhanced throughput and flexibility for real-time monitoring of biomolecular recognition at a reasonable cost. Some nanotechnologies will accelerate target identification, whereas others will evolve into therapeutics. This is closely related to drug
delivery, another important pharmaceutical aspect of nanobiotechnology.
Nanoparticles for Drug Discovery
Nanocrystals (QDs) and other nanoparticles (gold colloids, magnetic nanoparticles,
nanobarcodes, nanobodies, dendrimers, fullerenes, and nanoshells) have received
considerable attention recently with their unique properties for potential use in drug
discovery.
Use of Gold Nanoparticles for Drug Discovery
Gold nanoparticles can emit light so strongly that it is readily possible to observe a
single nanoparticle at laser intensities lower than those commonly used for multiphoton absorption-induced luminescence (7). Moreover, gold nanoparticles do not
blink or burn out, even after hours of observation. These observations suggest that
metal nanoparticles are a viable alternative to fluorophores or semiconductor nanoparticles for biological labeling and imaging. Other advantages of the technique are
that the gold nanoparticles can be prepared easily, have very low toxicity, and can
readily be attached to molecules of biological interest. In addition, the laser light
used to visualize the particles is a wavelength that causes only minimal damage to
most biological tissues. This technology could enable tracking of a single molecule
of a drug in a cell or other biological samples.
Use of Quantum Dots for Drug Discovery
The use of QDs for drug discovery has been explored extensively. Both advantages
and drawbacks have been investigated (8). Advantages of the use of QDs for drug
discovery are as follows:
■
■
■
■
Enhanced optical properties as compared with organic dyes. QDs offer great
imaging results that could not be achieved by organic dyes.
Multiple leads can be tested on cell culture simultaneously. Similarly, the
absorption of several drug molecules can be studied simultaneously for a longer
period of time.
Using the surface functionalization properties of QDs, targeting capabilities can
be added as well.
Due to the inorganic nature of QDs, their interaction with their immediate
environment at in vivo states can be minimal compared with their organic
counterparts.
Drawbacks of QDs for drug discovery are as follows:
■
Size variation during the synthesis of single color dots is 2% to 4% and for
applications such as capillary electrophoresis or gel electrophoresis, it could
176
■
■
■
■
Jain
create false results. Therefore, QD synthesis techniques need to have improved
quality control with respect to size distribution before they can be seriously
utilized in drug-discovery research.
For absorption, distribution, metabolism and excretion (ADME) purposes, blue
QDs (diameter of 3.7 nm) are the smallest class of the QD family but they are
considerably larger than organic dyes. Hence, the use of QDs for this purpose
might not be desirable in special cases.
Similarly, the number of functional groups attached to an organic dye is usually 1,
or it can be controlled very precisely.
The transport of a large volume (due to multiple attachments of drug molecules
to a single QD) across the membrane will be more difficult than a single molecule
itself.
The “blinking” characteristics of QDs when they are excited with high-intensity
light could be a limiting factor for fast scan systems such as flow cytometry.
Other Nanotechnologies for Drug Discovery
None of the nanoparticles available is ideal for all requirements of drug discovery.
The choice may depend on the needs. Nanodevices such as nanobiosensors and
nanobiochips are being used to improve drug discovery and development. Nanoscale
assays can significantly contribute to cost-saving in screening campaigns. In addition, some nanosubstances may be potential drugs of the future. These include
dendrimers, fullerenes, and nanobodies (9).
Dendrimers are a novel class of three-dimensional nanoscale, core-shell structures that can be precisely synthesized for a wide range of applications. They are
most useful in drug delivery and can also be used for the development of new pharmaceuticals with novel activities. Polyvalent dendrimers interact simultaneously
with multiple drug targets. They can be developed into novel-targeted cancer therapeutics (10).
A key attribute of the fullerene molecules is their numerous points of attachment,
allowing for precise grafting of active chemical groups in three-dimensional orientations. This attribute, the hallmark of rational drug design, allows for positional control
in matching fullerene compounds to biological targets. In concert with other attributes,
namely the size of the fullerene molecules, their redox potential and their relative
inertness in biological systems, it is possible to tailor requisite pharmacokinetic characteristics to fullerene-based compounds and optimize their therapeutic effect (11).
Nanobodies are the smallest available intact antigen-binding fragments
harboring the full antigen-binding capacity of the naturally occurring heavy-chain
antibodies. Nanobodies have the potential of a new generation of antibody-based
therapeutics as well as diagnostics for diseases such as cancer (12).
ROLE OF NANOBIOTECHNOLOGY-BASED DRUG DELIVERY
IN DEVELOPMENT OF NANOMEDICINE
Several nanoparticle-based technologies for drug delivery are described in various
chapters of this book. This section will briefly describe the relevance of these technologies for practical human therapeutics.
An important clinical aspect of nanoparticle-based therapeutics is targeted
drug delivery. When nanoparticles are used in the treatment of cancer, their powerful targeting ability and potential for large cytotoxic payload dramatically enhance
the efficacy of conventional pharmaceuticals as well as novel therapeutic approaches,
such as gene therapy, radioimmunotherapy, and photodynamic therapy.
Role of Nanobiotechnology in the Development of Nanomedicine
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There are particular advantages of drug delivery for the treatment of various
diseases by nanoscale devices. There are several requirements for developing a
device small enough to efficiently leave the vasculature and enter cells to perform
multiple, smart tasks. However, the major requirement involves size. Vascular pores
limit egress of therapeutics to materials less than approximately 50 nm in diameter,
and cells will not internalize materials much greater than 100 nm. As a result, the
only currently available technology that fulfills these criteria consists of synthetic
nanodevices. These are designed, synthetic materials with structures less than 100 nm
in size. Unlike fictional mechanical nanomachines, based on devices that have been
“shrunken” to nanometer dimensions, several true nanomolecular structures have
now been synthesized and applied to drug delivery, gene transfer, antimicrobial
therapeutics, and immunodiagnostics.
Nanoparticles are important for delivering drugs intravenously so that they
can pass safely through the body’s smallest blood vessels, for increasing the surface
area of a drug so that it will dissolve more rapidly, and for delivering drugs via
inhalation. Porosity is important for entrapping gases in nanoparticles, for controlling the release rate of the drug, and for targeting drugs to specific regions. Owing
to their small size, lipid nanocapsules might be promising for an injectable as well
as for an oral drug-delivery system, providing sufficient drug solubility to avoid
embolization in blood after intravenous injection as well as a positive effect of drug
absorption after oral administration. A drug-delivery system for intravenous administration of ibuprofen has been developed which exhibits sustained-release properties by either oral or intravenous route and may be useful for the treatment of
postoperative pain.
NANOMEDICINE
Besides nanoparticles, various nanotechnologies and other nanomaterials that are
currently under investigation in medical research and diagnostics will soon find a
practical application in practice of medicine. Nanobiotechnologies are being used to
create and study models of human disease, particularly immune disorders.
Introduction of nanobiotechnologies in medicine will not create a separate branch
of medicine but simply implies improvement of diagnosis as well as therapy and
can be referred to as nanomedicine. This broad term covers various therapeutic
areas including treatments that may require surgical intervention. Applications of
nanobiotechnology in medicine are shown in Table 1.
Clinical Nanodiagnostics
Application of nanotechnology in molecular diagnostics will have a tremendous
impact on the practice of medicine. Biosensor systems based on nanotechnology
could detect emerging disease in the body at a stage that may be curable. This is
extremely important in management of infections and cancer. Some of the body
functions and responses to treatment will be monitored without cumbersome laboratory equipment. Some examples are a radiotransmitter small enough to put into a
cell and acoustical devices to measure and record the noise a heart makes.
Nanodiagnostics will also be integrated with nanotherapeutics.
Nanoendoscopy
Endoscopic microcapsules that can be ingested and precisely positioned are being
developed. A control system will enable the capsule to attach to the digestive tract
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TABLE 1 Nanomedicine in the Twenty-First Century
Nanodiagnostics
Molecular diagnostics
Nanoendoscopy
Nanoimaging
Nanotechnology-based drugs
Drugs with improved methods of delivery
Regenerative medicine
Tissue engineering with nanotechnology
Transplantation medicine
Exosomes from donor dendritic cells for drug-free organ transplants
Nanorobotic treatments
Vascular surgery by nanorobots introduced into the vascular system
Nanorobots for detection and destruction of cancer
Implants
Bioimplantable sensors that bridge the gap between electronic and neurological
circuitry
Durable rejection-resistant artificial tissues and organs
Implantations of nanocoated stents in coronary arteries to elute drugs and to prevent
reocclusion
Implantation of nanopumps for drug delivery
Minimally invasive surgery using catheters
Miniaturized nanosensors implanted in catheters to provide real-time data to surgeons
Nanosurgery by integration of nanoparticles and external energy
Source : From Ref. 1.
and move within it. Precisely positioned microcapsules would allow physicians to
view any part of the inside lining of the digestive tract in detail, resulting in more
efficient, accurate, and less invasive diagnoses. In addition, these capsules could be
modified to include treatment mechanisms as well, such as the release of a drug or
chemical near a diseased area.
PillCam™ capsule (Given Imaging Ltd., Yoqneam, Israel), an endoscope to
visualize small intestine abnormalities, was approved in 2001. Other companies are
now producing ingestible capsules for this purpose. The patient ingests the capsule,
which contains a tiny camera, and intestinal peristalsis propels the capsule for
approximately eight hours. During this time, the camera snaps the pictures and
images that are transmitted to a data recorder worn by the patient. The physicians
can review the images later on to make the diagnosis, but some abnormalities may
be missed as this method has only a 50% success rate in detection of diseases.
Controlling the positioning and movement on a nanoscale will greatly improve the
accuracy of this method. Similar nanorobots are under development for other parts
of the body.
Nanobiotechnology for Developing Stem-Cell-Based Therapies
Stem-cell-based therapies are one of the most promising areas of development in
human therapeutics. Nanobiotechnology can be applied to delivery of gene therapy
using geneti-cally modified stem cells and further applied in tracking stem cells
introduced into the human body.
Nanofibrous scaffolds are being developed for stem cells to mimic the nanometer-scale fibers normally found in that matrix (13). They are being used to grow
stem cells derived from adipose tissue. They can conceivably be used for tissue repair.
Role of Nanobiotechnology in the Development of Nanomedicine
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Application of Nanobiotechnology in Various Therapeutic Areas
Nanobiotechnology has been applied in almost every area of human healthcare.
Some examples are given of applications in important therapeutic areas: cancer,
neurological disorders, cardiovascular diseases, and infections.
Oncology
Nanoparticles can deliver chemotherapy drugs directly to tumor cells and then give
off a signal after the cells are destroyed. Drugs delivered this way are several times
more potent than standard therapies. Combination of gold nanoparticles followed
by X-ray treatment reduces the size of the tumors, or completely eradicates them in
mice (14). The technique works because gold, which strongly absorbs X rays, selectively accumulates in tumors. This increases the amount of energy that is deposited
in the tumor compared with nearby normal tissue. Gold nanoparticle radiotherapy
for patients is under consideration for clinical trials.
Nanoshells may be combined with targeting proteins and used to ablate target
cells. This procedure can result in the destruction of solid tumors or possibly metastases not otherwise observable by the oncologist. In addition, nanoshells can be
utilized to reduce angiogenesis present in cancer. Experiments in animals, in vitro
and in tissue demonstrate that specific cells (e.g., cancer cells) can be targeted and
destroyed by an amount of infrared light that is otherwise not harmful to surrounding tissue. This procedure may be performed using an external (outside the body)
infrared laser. Prior research has indicated the ability to deliver the appropriate
levels of infrared light at depths of up to 15 cm, depending upon the tissue.
Photothermal tumor ablation in mice has been achieved by using near-infraredabsorbing nanoparticles (15). Nanoshells enable a seamless integration of cancer
detection and therapy.
It is within the realm of possibility to use molecular tools to design a miniature
device that can be introduced in the body, locate and identify cancer cells, and finally
destroy them. The device would have a biosensor to identify cancer cells and a
supply of anticancer substance that could be released on encountering cancer cells.
A small computer could be incorporated to program and integrate the combination
of diagnosis and therapy and provide the possibility to monitor the in vivo activities
by an external device. As there is no universal anticancer agent, the computer program could match the type of cancer to the most appropriate agent. Such a device
could be implanted as a prophylactic measure in persons who do not have any obvious manifestations of cancer. It would circulate freely and could detect and treat
cancer at the earliest stage. Such a device could be reprogrammed through remote
control and enable change of strategy if the lesion encountered is other than cancer.
Disorders of the Central Nervous System
Applied nanobiotechnology aimed at the regeneration and neuroprotection of the
central nervous system (CNS) will significantly benefit from basic nanotechnology
research conducted in parallel with advances in cell biology, neurophysiology, and
neuropathology (16). QD technology is used to gather information about how the
CNS environment becomes inhospitable to neuronal regeneration following injury
or degenerative events by studying the process of reactive gliosis. Glial cells, housekeeping cells for neurons, have their own communication mechanisms that can be
triggered to become reactive following injury. QDs, with added bioactive molecules,
might spur growth of neurites in a way that provides a medium that will encourage
this growth in a directed way.
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Nanoparticles can be used as an aid to neurosurgery. Iron oxide nanoparticles
can outline not only brain tumors under MRI, but also other lesions in the brain that
may otherwise have gone unnoticed (17).
Cardiovascular Diseases
The diagnosis and treatment of unstable plaque is an area in which nanotechnology
could have an immediate impact. Nanoprobes can be targeted to plaque components for noninvasive detection of patients at risk. Targeted nanoparticles, multifunctional macromolecules, or nanotechnology-based devices could deliver therapy
to a specific site, localized drug release being achieved either passively (by proximity alone) or actively (through supply of energy as ultrasound, near-infrared, or
magnetic field). Targeted nanoparticles or devices could also stabilize vulnerable
plaque by removing material, for example, oxidized low-density lipoproteins.
Devices able to attach to unstable plaques and warn patients and emergency medical services of plaque rupture would facilitate timely medical intervention.
Restenosis after percutaneous coronary intervention continues to be a serious
problem in clinical cardiology. Recent advances in nanoparticle technology have
enabled the delivery of NK911, an antiproliferative drug, selectively to the ballooninjured artery for a longer time (18). NK911 is a core-shell nanoparticle of polyethylene
glycol-based block copolymer encapsulating doxorubicin. It accumulates in vascular
lesions with increased permeability. In a balloon injury model of the rat carotid artery,
intravenous administration of NK911 significantly inhibited the neointimal formation.
The effect of NK911 was due to inhibition of vascular smooth muscle proliferation but
not to enhancement of apoptosis or inhibition of inflammatory cell recruitment.
NK911 was well tolerated without any adverse systemic effects. These results suggest
that nanoparticle technology is a promising and safe approach to target vascular
lesions with increased permeability for the prevention of restenosis after balloon
injury. A novel approach to prevention of restenosis involves incorporation of nitric
oxide (NO)-eluting nanofibers into stents for antithrombogenic action. NO has vasodilating action as well, which may be beneficial in ischemic heart disease.
Infections
An important role of nanotechnology in the management of infections is use of formulations which improve the action of known bactericidal agents. The bactericidal
properties of some agents are manifest only in nanoparticulate form. Certain formulations of nanoscale powders possess antimicrobial properties. These formulations
are made of simple, nontoxic metal oxides such as magnesium oxide (MgO) and
calcium oxide (CaO, lime) in nanocrystalline form, carrying active forms of halogens, for example, MgO.Cl2 and MgO.Br2. When these ultrafine powders contact
vegetative cells of Escherichia coli, Bacillus cereus, or Bacillus globigii, over 90% are
killed within a few minutes. Likewise, spore forms of the Bacillus species are decontaminated within several hours. Dry contact with oflatoxins and contact with MS2
bacteriophage (surrogate of human enterovirus) in water also cause decontamination in minutes. A nanopowder of MgO can scour contaminated rooms of anthrax
spores (19). Unlike antibacterial gases and foams which are messy, corrosive, and
ruin electrical equipment, the powder can be sprayed into rooms and swept or
vacuumed up. The chemical specks attract oppositely charged spores. The particles
then cut open and chemically break down the spores’ tough outer shell.
Silver powder with particle size ranging from 50 to 100 nm has a homogeneous distribution of nanoparticles in the material and antiinfective properties.
Role of Nanobiotechnology in the Development of Nanomedicine
181
Silver nanoparticles have been incorporated in commercial preparations for wound
care to prevent infection.
A simple molecule from a hydrocarbon and an ammonium compound has
been used to produce a unique nanotube structure with antimicrobial capability
(20). The quaternary ammonium compound is known for its ability to disrupt cell
membranes and causes cell death, whereas the hydrocarbon diacetylene can change
colors when appropriately formulated; the resulting molecule would have the
desired properties of both a biosensor and a biocide.
Antimicrobial nanoemulsions, containing water and soybean oil with uniformly sized droplets in the 200 to 400 nm range, can destroy microbes effectively
without toxicity or harmful residual effects (21). The nanoparticles fuse with the
membrane of the microbe and the surfactant disrupts the membrane, killing the
microbe. The classes of microbes eradicated are viruses (e.g., HIV, herpes), bacteria
(e.g., E. coli, Salmonella), spores (e.g., anthrax), and fungi (e.g., Candida albicans,
Byssochlamys fulva).
ROLE OF NANOBIOTECHNOLOGY IN THE DEVELOPMENT
OF PERSONALIZED MEDICINE
Personalized medicine simply means the prescription of specific therapeutics best
suited for an individual. It is usually based on pharmacogenetic, pharmacogenomic,
and pharmacoproteomic information, but other individual variations in patients are
also taken into consideration (22,23). Personalized medicine is beginning to be
recognized and is expected to become a part of medical practice within the next
decade. Molecular diagnostics is an important component of personalized medicine. Improvement of diagnostics by nanotechnology has a positive impact on
personalized medicine. Nanotechnology has potential advantages in applications in
point-of-care diagnosis, for example, on patient’s bedside or the outpatient clinic,
self-diagnostics for use in the home, and integration of diagnostics with therapeutics. All of these will facilitate the development of personalized medicines. Cancer is
a good example of advantages of personalized management. In cases of cancer, the
variation in behavior of cancer of the same histological type from one patient to
another is also taken into consideration. Personalization of cancer therapies is based
on a better understanding of the disease at the molecular level, and nanotechnology
will play an important role in this area (24).
CONCLUDING REMARKS AND FUTURE PROSPECTS
Disease and other disturbances of function are caused largely by damage at the
molecular and cellular level, but current surgical tools are large and crude. Even a
fine scalpel is a weapon more suited to tear and injure than heal and cure. It would
make more sense to operate at the cell level to correct the cause of disease, rather
than chop off large lesions as a result of the disturbances at cell level.
Nanotechnology will enable construction of computer-controlled molecular
tools that are much smaller than a human cell and built with the accuracy and precision of drug molecules. Such tools will be used for interventions in a refined and
controlled manner at the cellular and molecular levels. They could remove obstructions in the circulatory system, kill cancer cells, or take over the function of subcellular organelles. Instead of transplanting artificial hearts, a surgeon of the future
would be transplanting artificial mitochondrion.
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Nanotechnology will also provide devices to examine tissue in minute detail.
Biosensors that are smaller than a cell would give us an inside look at cellular function. Tissues could be analyzed down to the molecular level, giving a completely
detailed “snapshot” of cellular, subcellular, and molecular activities. Such a detailed
diagnosis would guide the appropriate treatment.
An increasing use of nanobiotechnology by the pharmaceutical and biotechnology industries is anticipated. Nanotechnology will be applied at all stages of
drug development − from formulations for optimal delivery to diagnostic applications in clinical trials.
It is expected that within the next few years, we will have a better understanding of how to coat or chemically alter nanoparticles to reduce their toxicity to the
body, which will allow us to broaden their use for disease diagnosis and for drug
delivery. Biomedical applications are likely to be some of the earliest. The first
clinical trials are anticipated for cancer therapy.
REFERENCES
1. Jain KK. Nanobiotechnology: applications, markets, and companies. Basel: Jain Pharma
Biotech Publications, 2007.
2. Jain KK. Nanodiagnostics: application of nanotechnology in molecular diagnostics.
Expert Rev Mol Diagn 2003; 4:153–161.
3. Jain KK. Nanobiotechnology in Molecular Diagnostics. Norwich, UK, Norwich: Horizon
Scientific Press, 2006.
4. Bao YP, Huber M, Wei TF, et al. SNP identification in unamplified human genomic DNA
with gold nanoparticle probes. Nucleic Acids Res 2005; 33:e15.
5. Storhoff JJ, Lucas AD, Garimella V, Bao YP, Müller UR. Homogeneous detection of unamplified genomic DNA sequences based on colorimetric scatter of gold nanoparticle
probes. Nat Biotechnol 2004; 22: 883–887.
6. Bulte JW, Arbab AS, Douglas T, et al. Preparation of magnetically labeled cells for cell
tracking by magnetic resonance imaging. Methods Enzymol 2004; 386:275–299.
7. Farrer RA, Butterfield FL, Chen VW, Fourkas JT. Highly efficient multiphoton-absorptioninduced luminescence from gold nanoparticles. Nano Lett 2005; 5:1139–1142.
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10. Kukowska-Latallo JF, Candido KA, Cao Z, et al. Nanoparticle targeting of anticancer
drug improves therapeutic response in animal model of human epithelial cancer. Cancer
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11. Wilson SR. Nanomedicine: fullerene and carbon nanotube biology. In: Osawa E, ed.
Perspectives in Fullerene Nanotechnology. Kluwer Academic Publishers, 2002.
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13. Kang X, Xie Y, Kniss DA. Adipose tissue model using three-dimensional cultivation of
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14. Hainfeld J, Slatkin DN, Smilowitz HM. The use of gold nanoparticles to enhance radiotherapy in mice. Phys Med Biol 2004; 49:N309–N315.
15. O’Neal DP, Hirsch LR, Halas NJ, et al. Photo-thermal tumor ablation in mice using near
infrared-absorbing nanoparticles. Cancer Lett 2004; 209:171–176.
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by high-epitope density nanofibers. Science 2004; 303:1352–1355.
17. Neuwelt EA, Varallyay P, Bago AG, et al. Imaging of iron oxide nanoparticles by MR and
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18. Uwatoku T, Shimokawa H, Abe K, et al. Application of nanoparticle technology for the
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20. Lee SB, Koepsel R, Stolz DB, et al. Self-assembly of biocidal nanotubes from a singlechain diacetylene amine salt. J Am Chem Soc 2004; 126:13400–13405.
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TCRT 2005; 4:645–650.
13
Pharmaceutical Applications of
Nanoparticulate Drug-Delivery Systems
Yashwant Pathak
UCB Manufacturing, Inc., Rochester, New York, U.S.A.
Deepak Thassu
UCB Pharma, Inc., Rochester, New York, U.S.A.
Michel Deleers
Global Pharmaceutical Technology and Analytical Development (GPTAD),
UCB, Braine l’Alleud, Belgium
INTRODUCTION
The successful introduction of the first drug-delivery system (DDS) brought about
tremendous interest in the research and application of these systems, especially for
the entry of drugs in the systemic circulation of the body. The goal of any DDS is to
provide a therapeutic amount of drug to the proper site in the body while achieving
it rapidly and maintaining desired drug concentrations in the body circulation.
Most drugs are delivered to patients using a systemic approach, the belief being that
if you flood the body with enough active compounds, there will be a desired therapeutic effect on the body. The DDSs are supposed to deliver the drug at a rate dictated by the needs of the body over a specified period of the treatment. The idealized
objective for the DDS points to two major aspects, namely spatial placement and
temporal delivery of the drug. Spatial placement relates to targeting a drug to a
specific organ or tissue. Temporal delivery refers to controlling the rate of delivery
to the target tissues (1). Hundreds of drug-delivery products have been introduced
in the market and many are now in different stages of development.
Lately, in the field of pharmaceutical research, a plethora of drug-delivery
groups and companies have emerged, partly as a result of surging interest in generic
drug development and continued technological advances. The drug-delivery technology can enhance the therapeutic as well as the commercial value of the healthcare
products. It takes into consideration the carrier, the route, and the target, and
develops a strategy of processes or devices designed to increase the therapeutic
efficacy of the drug and in many cases reduce the side effects of the drug.
The drug-delivery sector has evolved from being simply a part of the pharmaceutical production process to a driving force for innovation and profits. Industry and
commercial interests in drug-delivery continue to build steadily. The benefits to the
patients can be seen in improved compliance and medical outcomes. The pharmaceutical industry is able to extend the patent protection through novel DDSs and bring
new therapies to the market. U.S. demands for DDSs will grow nearly 9% annually to
more than $82 billion by 2007. Growth opportunities extend into all therapeutic classes
of pharmaceuticals: respiratory, central nervous system (CNS), and cardiovascular
agents will remain the top three applications based on the special formulating needs
of medicines for conditions such as asthma, arthritis, and hypertension. The other
areas such as anticancer drugs, hormones, and vaccines will also follow the track.
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NANOPARTICULATE DRUG-DELIVERY SYSTEMS
Nanoparticulate drug-delivery systems (NPDDSs) are being explored for the
purpose of solving the challenges of drug delivery. Coming in many shapes and
sizes, most carriers are less than 100 nm in diameter. NPDDSs provide methods for
targeting and releasing therapeutic compounds in very defined regions. These vehicles have the potential to eliminate or at least ameliorate many problems associated
with drug distribution. As many drugs have a hydrophobic component, they often
suffer from problems of precipitation in high concentration, and there are many
examples of toxicity issues with excipients designed to prevent drug aggregation.
To combat these issues, many NPDDSs provide both hydrophobic and hydrophilic
environments, which facilitate drug solubility. Alternatively, many drugs suffer
from rapid breakdown and/or clearance in vivo. Encapsulating the drugs in a
protective environment, NPDDSs increase their bioavailability, thereby allowing
the clinicians to prescribe lower doses.
With recent advances in polymer and surface conjugation techniques as well as
microfabrication methods, perhaps the greatest focus in drug-delivery technology is
in the design and applications of NPDDSs. Ranging from simple metal–ceramic core
structure to complex lipid–polymer matrices, these submicron formulations (2) are
being functionalized in numerous ways to act as therapeutic vehicles for a variety of
conditions (Fig. 1).
NPDDSs can be defined as the DDSs where nanotechnology is used to deliver
the drug at nanoscale. Below 100 nm, materials exhibit different, more desirable
physical, chemical, and biological properties. Given the enormity and immediacy of
the unmet needs of therapeutic areas such as CNS disorders, this can lead to drugs
that can extend life and save untimely deaths (2).
Advantages of Nanoparticulate Drug-Delivery Systems
The nanoparticles (NPs) may offer some advantages such as protection of drugs
against degradation, targeting the drugs to specific sites of action, organ or tissues,
FIGURE 1 (See color insert.)
Different types of nanoparticulate
structures.
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
187
and delivery of biological molecules such as proteins, peptides, and oligonucleotides.
A number of different strategies have been proposed in order to modify the physicochemical characteristics of the NPs, and thus their interactions within the biological
systems. For example, it is possible to change the chemical nature of the polymeric
matrix of the NPs and thereby alter certain biological phenomena such as biorecognition, biodistribution, bioadhesion, biocompatibility, and/or biodegradation. Some
polymeric materials used for this purpose are gelatin, chitosan (CS), sodium alginate,
poly(alkyl)cynoacrylates, poly(lactic acid) (PLA), poly(lactic-co-glycolic acid)
(PLGA), poly(ethylene glycol-co-(lactic-glycolic acid), poly(caprolactone), and
polymethyl methacrylate (3–5). Another approach to modify the biological response
is based on the incorporation of suitable adjuvants in the NPs, like proteins such as
albumin, invasins, and lectins, and polymers such as poloxamers and poloxamines
(6). Different manufacturing methods can also enable modifications of the physicochemical characteristics of NPs such as size, shape, structure, morphology, texture,
and composition (3).
Manufacturing Techniques for Nanoparticulate Drug-Delivery Systems
Conventionally, two groups of manufacturing techniques have been reported for
producing NPs. The first involves polymerization of the monomers, whereas the
second one is based on dispersion of the performed polymers. The salting-out (7),
emulsification-diffusion (8), and nanoprecipitation (9) can be cited as typical examples of the second method. NP is a collective term used to describe the nanospheres
and nanocapsules (NCs). The difference between these forms lies in the morphology
and the architecture. NCs are composed of a liquid core (generally an oil) surrounded
by polymeric membrane, whereas nanospheres are formed by a dense polymeric
matrix (10). NCs are pharmaceutically attractive due to their oil-based central cavities, which allow a high encapsulation level for lipophilic substances, enabling
improved drug delivery. It is possible to avoid drug precipitation during preparation
and subsequent stability problems caused by the presence of the drug on the surface
of the NPs. Two techniques are widely used to prepare a biodegradable NC.
Interfacial Polymerization of Alkylcyanoacrylate Monomers
In this process, the cyanoacrylate monomer and the lipophilic drug are dissolved in
a mixture of oil and ethanol. This organic solution is then added slowly to water or a
buffer solution (pH 3–9) containing surfactants such as poloxamers or phospholipids. NCs are formed spontaneously by anionic polymerization of the cyanoacrylate
in the oily phase after contact with hydroxyl ions which act as initiators.
Interfacial Deposition of Performed Polymers
In this process, the lipophilic drug, oil polymer, and optionally phospholipids are
dissolved in a water-miscible solvent (e.g., acetone). This solution is then poured
while stirring into an aqueous solution containing a nonionic surfactant (e.g.,
poloxamer 188). NCs are instantly formed by the fast diffusion of solvent into water,
which provokes the spontaneous emulsification of the oily solution in the form of
nanodroplets where the dissolved polymer will form a film around the droplets that
contain the drug (11).
The method of interfacial polymerization is not ideal for three reasons: (i) the
probable presence of the residual, (ii) potentially toxic monomers or oligomers, and
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Pathak et al.
the possibility of cross-reaction with the drug, and (iii) the difficulty in predicting
the molecular weight of the resulting polymer (12). The principal drawback of this
method is the polymer aggregation that is frequently observed when working with
high polymer concentration or low organic solvent/water ratio. Guerrero et al. (13)
have described a new process based on an emulsification-diffusion technique, overcoming these drawbacks. This study demonstrated that the emulsification-diffusion
technique represents a viable alternative for preparing biodegradable NCs starting
from performed polymers. It is simple and versatile, and permits high efficiency of
entrapment of lipophilic drugs (13).
It is important to note that a deeper understanding of the physicochemical
phenomena involved during the NPs formation is also necessary. Specifically, the
relationship between physicochemical parameters and their quantitative effects on
NP features could be an invaluable tool in the controlled engineering of particles.
Knowledge of these fundamental relationships would allow NPs to be designed
with defined size and surface characteristics for delivery to specific cells or organs
without requiring exhaustive experimental procedures. Rodriguez et al. (3) studied
the influence of certain physicochemical properties of the aqueous and organic
phases used during NP preparation and the effects on the characteristics of NPs
produced by salting-out, emulsification-diffusion, and nanoprecipitation methods,
and concluded that the mean size of the NPs could be narrowed, using different
methods. For example, salting-out offered NP mean size range between 123 and
710 nm. Emulsification method gave 110 to 715 nm mean size, whereas nanoprecipitation gave a very narrow size range distribution of 147 to 245 nm (3).
The water–solvent interaction and diffusion motion of the solvent play an
important role in explaining the variation of the NP size during NP preparation by
the nanoprecipitation method. Common disadvantages of solid–lipid nanoparticles
(SLNs) include: particle growing, unpredictable gelation tendency, unexpected
dynamics of polymorphic transitions, and inherent low incorporation rates resulting
from the crystalline structure of the solid lipids (14).
NANOPARTICULATE DRUG-DELIVERY SYSTEM APPLICATIONS
NPDDSs have been utilized for their therapeutic applications over a wide range,
from cancer treatments to some over the counter (OTC) preparations. Many drugs
have been used as model drugs for specific spatial and temporal applications.
Table 1 enumerates various drugs which have been used in NPDDSs. In the
pharmaceutical formulations, NPDDSs can be used with advantages. Table 2 shows
applications of NPDDSs to address various formulations issues.
Nanoparticulate Drug-Delivery Systems for Proteins and Peptides
Large numbers of new therapeutic proteins and peptides are being discovered,
thus protein drug-delivery technologies are of ever increasing importance.
Traditionally, the protein is delivered parenterally via solutions that are injected
subcutaneously, intramuscularly, and intravenously. Although such injections benefit from high bioavailability, they fail to provide sustained plasma concentrations
and suffer from poor patient compliance due to the required frequency of injections. NPDDSs are designed to provide the drug release over an extended period
of time, thereby minimizing the need for frequent injections. These can be used for
systemic or oral delivery, and the biodegradable nature of the nanoparticulate
materials alleviates the need for surgical removal.
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Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
TABLE 1 Drugs Used for Nanoparticulate Drug Delivery Systems
Name of the drug
Aclacinomycin
Adriamycin
Antifungal drugs
Atovaquone
Betamethasone
Bifonazole
Brimonidine
Budesonide
Camptothecin
Cephalosporin
Cisplatin
Clotrimazole
Clobetasol
Clozapine
Curdlan derivative: anticancer drugs
B-Cyclodextrin
Cyclosporine
Cyclosporine
Cyclosporine
Cyclophosphamide
Diclofenac
Danazol
Darodipine
Delargin
Dexamethasone
Diminazine
Diminazenediaceturate
Gadolinium
5-Fluouracil
Flurbiprofen
Halofantrine
Heparin
Hydrocortisone
Idarubicin
Indomethacin
Isoniazid
Ketoprofen
Kytorphin
Loperamide
Methotrexate
Mitoxantrone
Nifedipine
Ontazolost
Paclitaxel
Phenothiazine
Pilocarpine
Praziquantel
Prednisolone
Carrier
PBCA NPs
PBCA NPs
Submicronized emulsion
SLN
CaCo3 NPs
B Cyclodextrin NPs
Polyacrylic NPs
Polylactic acid NPs nanosuspension
SLN
Nanoconjugates
Polymeric micelles
B Cyclodextrin NPs
SLN
SLN
SLN
Nanosphere
SLN
Stearic acid NPs
HPMCP
PBCA NCs
Inorganic microparticles
Polylactide
Polylactide
Caprolactone
Lipid-based emulsion
SLN
PBCA NPs
Supercritical carbon dioxide—PLGA
Lipid-based
Lipid drug conjugate
Lipid-based NPs
Colloidal NPs
Nanosuspension
Lipid-based emulsion
Methacrylate polymers
SLN
SLN
SLN
PL glycolide polymer
Polycaprolactone and Eudragit S 100
PBCA NPs
Polysorbate 80-coated PBCA NPs
Colloidal carriers
Magnetic NPs
SLN nanocrystals
Lipid-based delivery
SLN
Cetyl alcohol/polysorbate NPs
Gelatin NCs
PLGA
SLN
PLGA
PLGA NC
SLN
References
(15)
(16)
(17)
(18)
(19)
(20)
(21)
(22)
(23)
(24)
(25)
(20)
(26)
(27)
(28)
(29)
(30)
(31)
(32)
(33)
(34)
(35)
(36)
(37)
(35,38)
(36,39)
(37)
(38)
(39)
(40)
(41)
(42)
(43)
(44)
(45)
(46)
(47)
(48)
(49)
(50)
(51)
(52)
(42)
(53)
(54–56)
(57)
(58)
(59)
(60)
(61)
(54)
(62)
(63)
(64)
(Continued )
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Pathak et al.
TABLE 1 Drugs Used for Nanoparticulate Drug Delivery Systems (Continued)
Name of the drug
Porpofol
Progesterone
Protamine phosphorothioate
Pyrazinamide
Retinal
Rifabutine
Rifamycin
Tamoxifen
Tarazepide
Thiamine
Tobramycin
Tretinoin
Triclosan
Tubocurarine
Ubidecarone
UCB-35440-3
Vincristine
Vitamin A
Xanthone
Carrier
Lipid-free NC
SLN
NP complexes
PL glycolide
SLN
SLN
PL glycolide
Polycaprolactone NPs
Cyclodextrin
Lipid
SLN
SLN
Submicron emulsion and NCs
Polysorbate 80-coated PBCA NPs
SLN
Nanocrystals
Colloidal carriers
SLN
PLGA
References
(65)
(45)
(66)
(48)
(67)
(18)
(48)
(68)
(69)
(70)
(71)
(72)
(73)
(74)
(75)
(76)
(42)
(77)
(78)
Abbreviations: HPMCP, hydroxypropyl methyl cellulose phthalate; NC, nanocapsules; NPs, nanoparticles;
PBCA; polybutylcyno acrylate; PLGA, poly-lactic-co-glycolic acid; SLN, solid–lipid nanoparticles.
Biodegradable nanoparticulate delivery for protein must satisfy several technical requirements. Among these: the proteins should be encapsulated with a high
loading efficiency, and remain stable throughout the manufacturing process and the
course of their intended dosing period. NPs need to be less than 125 µm in diameter
and form a free flowing powder so that they can be resuspended in an injectable
vehicle and passed through a needle. The release profile of the drug needs to be
TABLE 2 Nanoparticulate Drug Delivery Systems: Formulation Applications
Addressing the drug-delivery problems
Solving the issues related to solubility
Overcoming the poor bioavailability of the drugs
Issues with fed/fasted variability
Pharmacokinetic variability
Finding solutions with nanoparticulate drugs
Technology advances
Reduction in particle size of the poorly water-soluble drugs
Increased active agent surface area
Benefits for faster dissolution
Greater bioavailability
Smaller drug doses
Diminished toxicity
Decreased dosing variability
Pharmacodynamic factors: applicable to peptides and other drugs
These can be formulated as receptor-specific
These can be more resistant to unspecific degradation
They can deliver the drug in encapsulated form to delay the degradation, set a depot form for
prolonged signaling, and increase the treatment efficacy as compared to substitution of the
natural form of peptide
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
191
reproducible and therapeutically effective and pose no pharmacological or
toxicological risks due to rapid early release or burst.
Macromolecules such as proteins and DNA play an increasingly important
role in our arsenal of therapeutic agents. Delivery of these molecules to their site
of action at the desired rate is a challenge because their transport through compartmental barriers, for example, endothelium and epithelium in the body, is
inefficient and/or they are readily metabolized. For controlled release or sitespecific delivery of such macromolecules, delivery systems are required which
need to be more sophisticated than our present day strategies. These systems
must be custom-made, taking into account both molecular size and specific
characteristics of these molecules. One has to build a platform of different delivery strategies that use input from technical, pharmaceutical, and biomedical
disciplines to meet these challenges.
The development of appropriate DDSs for new macromolecules coming out
of the biotech industry is a meaningful challenge to pharmaceutical scientists.
Proteins, peptides, oligonucleotides, and genes are very unstable compounds that
need to be protected from degradation in the biological environment. Their efficacy
is highly limited by their inability to cross biological barriers and to reach the target
sites. They are vulnerable to harsh conditions in the gastrointestinal tract, leading to
chemical and enzymatic degradation. The future of these molecules solely depends
on the delivery systems and appropriate carriers.
There are three possible pathways for protein and peptide drug absorption
through the GI tract. The first is via the M-cells of Payer’s patches, the second via a
transcellular route involving enterocytes, and the third via paracellular avenues
through tight junctions (6,79). The nanosystems are providing a viable alternative
for these drugs such as liposomes, polymeric micelles, and NPDDSs. One of the
crucial and pervasive troubles in human therapy is to achieve a balance between
toxicity and therapeutic effect of the drugs. Therefore, the site-specific delivery could
reduce such side effects at nontarget sites and increase the efficacy. Rodrigues et al.
(80) have reported an interesting work on lectin nanocarrier conjugate. They used
dextran/poly(e-caprolactone) polyester polymers and conjugated with three different proteins, lectins from leaves of Bauhinia monandra and Lens culinaris, and
bovine serum albumin (BSA). The NPs having a size around 200 nm could be used
for delivering proteins (80).
A polypeptide hormone consisting of 32 amino acids plays a crucial role in
both bone remodeling and calcium homeostasis. Yoo and Park (81) formulated
salmon calcitonin (sCT) into biodegradable PLGA NPs using sCT oleate complexes.
The sCT oleate complexes were prepared by hydrophobic ion pairing. SCT NPs
were readily taken up by Caco-2 cells and sCT was transported across the Caco-2
monolayer in vitro. In vivo experiments showed sCT was orally absorbed. Vranckx
et al. (82) also reported similar results where they used an NC formulation with
hydrophilic core for delivering salmon calcitonin in rats.
A study by Alphandary et al. (83) has shown the crossing of insulin through the
intestinal epithelial barrier to the blood compartment where it was absorbed by portions of the M-cell-free epithelium. The insulin was incorporated in biodegradable
poly (alkylcynoacrylate) NCs (83).
An excellent review discusses the strategies of enhancing the immunostimulatory effects of CpG oligonucleotides and outlines the latest development in the
application of liposomes and NPDDSs for the delivery of oligonucleotides with an
extensive literature survey (84). Leach et al. (85) demonstrated that excipient-free
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protein NPs prepared by spray-freezing into liquid technology (86) can be dispersed
into PLGA and PLA microparticles and the burst effect can be prevented. The
uniform encapsulation of the stable proteins at high loading was achieved with
minimal burst effect. NPs based on hydrogels are being developed for the delivery
of macromolecules.
An interesting mechanistic study reported by Mo and Lim (87) exhibited
uptake of wheat germ agglutinin-conjugated NPs by A549 cells. In this study, they
prepared the PLGA NPs by solvent diffusion method (88) and later surface modified
with wheat germ agglutinin through a two-step carbodiimide method. Cellular
uptake was studied using confluent A549 cells as an in vitro model of the type II
alveolar epithelial cells. Uptake of the WGA-conjugated PLGA NPs was compared to
that of NPs similarly modified with the bovine serum albumin (BSA) to demonstrate
the specificity of the surface WGA in enhancing the cellular uptake of the NPs. The
mechanism of uptake was studied by performing the uptake experiments under
several inhibiting conditions. They concluded that the grafting of WGA on PLGA
NPs has increased the uptake by five to eight times, hence this method can be
exploited for the intracellular delivery of therapeutic and diagnostic agents (87).
Targeted delivery of proteins and DNA requires a carrier system in submicron size or nanosize. This carrier needs to be target site (cell or tissue)-specific.
Often, the actual target site location is intracellular, and the delivery of the carrier
payload at this intracellular target site is a prerequisite for therapeutic success. For
example, plasmid DNA needs to be delivered inside the nucleus of the target cell
before the cell can express the desired therapeutic protein. A very good example of
this system is the immunoliposomes, where liposomes carrying the drug with
monoclonal antibodies or monoclonal antibody fragments are covalently attached
to the bilayer for targeting purposes. A selection of monoclonal antibodies, with
induced endocytic uptake, can lead to the entrance of immunoliposomes (Fig. 2)
into tumor cells (84,89).
Another application of liposome-dependent drug is diphtheria toxin A (DTA)
chain. Liposome dependent means that the drug as such cannot reach its target site
of action inside the cell without a carrier, as it cannot pass the cytoplasmic membrane
FIGURE 2 Immunoliposomes: the use of nanotechnology to build a carrier system for site-specific
delivery of a protein. The anti-EGFR antibody permits endocytosis helping the immunoliposomes
to enter the target cells after cell binding. The peptide di INF-7 induces the release of liposomeentrapped diphtheria toxin A chain from immunoliposomes. Abbreviation: EGFR, epidermal growth
factor receptor. Source: From Ref. 89.
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
193
without help (a carrier). Such a drug will show neither the desired nor the undesired pharmacological effects. Diphtheria toxin, a protein consisting of an A chain
coupled with a B chain, can readily enter cells through the transporter B chain.
Upon entering the cell, the A chain causes the cell kill with exceptional efficiency
by blocking ribosomal activity. Thus DTA (lacking B chain) alone needs a cellspecific transport system, that is, a system that transports it into the desired target
cells, for example, tumor cells (90).
The design of a custom-made carrier at the nanometer level is a targeted
delivery system for plasmid DNA to efficiently transfer only to target cells. The cationic polymer pDMAEMA (poly-(2)-(dimethylamino)ethyl methacrylate) condenses
plasmid DNA effectively into 100 nm NPs (polyplexes). In vitro transfection is very
efficient but in vivo was ineffective. Polyplexes were subsequently coated with lipids
yielding lipopolyplexes, which demonstrated target cell-specific binding characteristics, and were able to transfect the cells with hardly any cell toxicity (91).
Ocular Applications of Nanoparticulate Drug-Delivery System
Topical ophthalmic drugs have generally poor absorption in the eye due to the cornea’s low permeability to drugs and noncorneal factors such as rapid tear turnover,
nasolacrimal drainage, and systemic absorption. One of the major problems in
ocular delivery is providing and maintaining an adequate concentration of the therapeutic agent in the precorneal area. Topical drop administration of ophthalmic
drugs in aqueous solutions results in extensive drug loss due to tear fluid and eyelid
dynamics (21,92,93). Most noninvasive approaches for enhancing ocular drug
absorption involve the use of prodrugs, the use of viscosity agents designed to prolong the drug residence time, and colloidal systems (94,95). Polymeric NPs are
attractive colloidal systems because they demonstrate increased stability and have
a longer elimination half-life in tear fluid (up to 20 min), than do conventional drugs
applied topically to the eye, which have half-lives of just one to three minutes.
NPDDSs have been evaluated for ocular applications to enhance absorption
of thera-peutic drugs, improve bioavailability, reduce systemic side effects, and
sustain intraocular drug levels (96). NPDDSs have shown potential in the treatment
of external eye diseases (97). PLGA has been evaluated and proved to be a very
useful biodegradable polymer for NPDDS formulation due to its medical use,
biocompatibility, and safety (98). Qaddoumi et al. (95) have studied the characteristics and mechanism of uptake of PLGA-based NPDDSs for ophthalmic application.
They suggested that PLGA-based NPDDSs could be used for the enhancement of
drug absorption in the eye and the controlled release of proteins and drugs. Some
reports in ocular applications are very novel approaches involving periocular routes
for retinal drug delivery of Celecoxib and Aldose reductose inhibitors (99–101).
Salgueiro et al. (33) demonstrated ophthalmic application of cyclophosphamideloaded polybutylcyanoacrylate (PBCA) nanosphere as an immunosuppressive agent.
The morphometrical properties such as average particle size and polydisparity
index of these DDSs are adequate for ophthalmic application without induced
corneal or conjunctival irritation (33).
Nanoparticulate Drug-Delivery Systems for Pulmonary Treatment
Pulmonary drug delivery for both systemic and local treatments has many advantages over other delivery routes because the lungs have a large surface area
(43–102 m2), thin absorption barrier, and low enzymatic activity. In addition, the
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alveoli of the lungs have a slower mucociliary clearance than the airways, and the
lung epithelia are thinner and more permeable. There is a potential for possible systemic absorption of the peptides and proteins through the alveolar region of the
lungs. Several studies have exhibited the absorption of high-molecular-weight
drugs such as insulin, heparin, and GCSF (recombinant human granulocyte colonystimulating factor) through pulmonary DDSs (102–104,104a). As these peptides have
a short life, the development of delivery systems with sustained pharmacological
action would be very useful.
Innovative, noninvasive, inhalable DDSs for peptides are being explored for
lung disease therapy with the vasoactive intestinal peptide (VIP) being used in the
treatment of severe lung diseases. Owing to its known antiinflammatory and vasodilative properties, it has been demonstrated to possess a high therapeutic potential for
other lung diseases, which are common in industrialized countries. VIP unfortunately reveals a variety of bifunctions mediated by at least two different receptors on
the cell surface. The ways thought to deal with this situation are to develop a receptorspecific system to make it more specific in function. These need to be protected by
licensing, and allow for superior treatment compared to natural VIP. It can be
achieved by designing, engineering, and production of analog peptides, which will
be very useful. By modifying the peptide, VIP in the amino acid sequence become
more receptor-selective and more resistant to unspecific degradation. It can deliver
the new peptides in nanoencapsulated form (NPDDS) to delay the degradation, and
to set in depot form for prolonged signaling, increasing the treatment efficacy as
compared to substitution of the natural form of the peptide.
An enormous diversity of therapeutic agents is currently administered to the
patients via aerosol inhalation, and the number of potential drug candidates for
pulmonary application increase daily. The major areas of research and therapeutic
applications are asthma (105), cystic fibrosis (106), lung cancer (107), tuberculosis
(108–109), pulmonary hypertension (110), and diabetes (111). Nanostructured drug
delivery and targeting systems are tools to overcome the limitations of lung delivery by stabilizing and protecting the release in the bronchi and make lung therapy
through inhalation possible and effective. Some of the polymers such as PLGA,
protamine, thiomer, and lipid-based particles can be loaded by VIP or new designed
analogs. The parameters, which need to be tested, will be improved by in vitro
enzymatic stability, in vitro long-term drug release, and retarding properties and
bioavailability of the carrier.
Insulin-loaded PBCA NPs were studied by Zhang et al. (112); they demonstrated that the pulmonary administration of these NPs could significantly prolong
the hypoglycemic effect of insulin. It was reported that the bioavailability of insulin
NPs was relatively higher than that of solution when administered by pulmonary
route to normal rats, but when NPs were administered subcutaneously the bioavailability was comparatively lower compared to solution administered the same way
(112). Another study using PLGA NPs to deliver insulin by nebulization also showed
the usefulness of NPDDSs for insulin (113). An interesting study was reported by
Liu et al. (114,115) incorporating estradiol and colloidal gold NPs in PLGA NPs to
be used as a model for the pulmonary DDSs. They proposed that large porous NPs
can be used as delivery systems for the pulmonary tract.
Several issues complicating the development of aerosol formulation include:
compound loss during inhalation, dosing difficulties, enzymatic degradation within
the lungs, and the high cost of production. Nanoparticulate-controlled release DDS
has the potential to overcome many of these problems. Such formulation may be
incorporated in aerosol form, remaining stable against forces of degradation during
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
195
aerosolization. It can target a specific site or cell population in the lung, protect the
drug-aggressive elements in the pulmonary tract, and release the compound in a
predetermined manner concurrently. It can be inert to the surrounding tissues and
contains no irritant or toxic additives and degrades when applicable within an
acceptable period of time, producing no toxic byproducts (116). Polymeric
nanoparticulate systems show promise in fulfilling the stringent requirements of
the pulmonary DDSs.
An interesting study is reported by Dailey et al. (116) using short-chain
PLGA grafted onto an amine-substituted poly(vinyl alcohol) backbone (3-diethyl
amino-1-propylamine (8%)–poly(vinyl alcohol)-grafted poly(lactide-co-glycolide)
(DEAPA–PVAL-g-PLGA) polymer. This polymer has amphiphilic properties and
is highly suited for the pulmonary delivery system. It was also reported that by
adding varying amounts of polyanion such as carboxymethyl cellulose, dextran
sulfate, or even DNA to the polymer ring during NPs formation, NPs of variable
physicochemical properties could be generated enable. This can high loading of
various drugs along with greater stability. However, these polymer derivatives
were found to degrade from 24 hours to within a week. Some related studies have
described these aspects in detail (117,114).
NPs may be very effective DDSs for various pulmonary therapeutic schemes.
The study by Dailey et al. (118) investigated the effect of nebulization technology and
NP characteristics on the features of aerosol generation. They concluded that biodegradable NPs contained in the suspensions did not affect the aerosol droplet size in
a clinically relevant manner; however, both NP characteristics and the technique of
aerosolization influence NP aggregation occurring during the aerosolization (118).
Vila et al. (119) have shown that polyethylene glycol (PEG) coating of the PLA NPs
increased the absorption of drug in nasal mucosa. Pandey et al. (49) demonstrated
the application of NPDDSs for the treatment of experimental tuberculosis using
poly(d,l-lactide-co-glycolide) as a polymer. They used an inhalable system using the
NPs and three anti-TB drugs Rifampicin, Isoniazid, and Pyrazinamide (49).
Nanoparticulate Drug-Delivery Systems for Central Nervous System
The entry of a drug molecule into the brain is limited by one of the most challenging
barriers, the blood–brain barrier (BBB). The BBB consists of a continuous layer of
endothelial cells joined together by tight junctions (Zonulae occludens), which
severely restrict paracellular transport across the barrier. The BBB allows passive
diffusion of small lipid-soluble molecules, whereas hydrophilic substances or
molecules with high molecular weight have minimal passive permeation. The
mechanism of permeability regulation includes macrovascular endothelial tight
junctions, enzymatic regulation, and active brain efflux. Transport across BBB is
additionally regulated by a number of transporters including very effective efflux
transporters such as multidrug resistance-associated protein or p-glycoprotein.
Several strategies have been tried to cross the BBB; one alternative strategy is to use
drug-carrier systems such as liposomes, antibodies, and NPs (120). Numerous
studies have shown the applications of NPs for brain targeting (51,52,121,122).
Hexapeptide Dalargin, a Leu encephalin analogue with no BBB permeability
adsorbed to the surface of PBCA NPs, caused central analgesia after IV administration (123). Other drugs Tubocurarine (72), Doxorubicin (124), Kytorphin (51), and
Loperamide (52) were also used as model drugs for these purposes. Brain uptake of
NPs in these studies was suggested based on the fact that the drugs adsorbed to
PBCA NPs caused a resultant pharmacologic effect in the CNS (51,72,123). Brain
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Pathak et al.
distribution of drugs delivered on the surface of NPs was also confirmed by quantification of the drug in the brain tissue itself (124). Studies have also shown the intact
presence of NPs in brain cells in vivo (125). Koziara et al. (120) tried to quantify the
presence of the NPs in brain in situ and studied the impact on BBB parameters (126).
Another study from the same laboratories showed the effectiveness of using microemulsions as precursors to engineer NPs. The advantages with the microemulsions
were simplistic production of the NPs approximately 100 nm in diameter, possible
incorporation of hydrophobic drugs in oil droplets, and inclusion of site-specific
ligands. They also reported the kinetic modeling of brain uptake, and their data
suggested the probable mechanism of brain entry (70).
Nanoparticulate Drug-Delivery Systems for Enzymes
Many scientists attempted application of nanocarriers for delivery of therapeutic
enzymes (127,128). A catalase-delivery system seems to be an ideal testing model for
evaluating the key aspects of enzyme delivery by nanocarriers. There is a fascinating study describing the loading and protection of an active enzyme catalase into a
poly-nanocarrier composed of diblock PEG–PLGA copolymers (127). It showed
nanocarriers in the size range of 200 to 500 nm protecting at least 25% of the loaded
proteases. Figure 3 demonstrates the method of formulation of PLGA–PLA NPs
FIGURE 3 Formulation of PLGA–PEG nanoparticles: (A) scheme with the double emulsion synthesis procedure and (B) by sonication in DCM/acetone mixture. (C) The loading efficiency was
determined by tracing the amount of radiolabelled catalase contained inside either the micro particles
(i.e., pellet obtained after centrifuging for 15 minutes at 1000 X g) or nanoparticles fractions (2nd pellet
obtained by centrifuging for 30 minutes at 22,000 X g). Loading efficiency was greatly enhanced
(eightfold) when a freeze thaw cycle (grey bar) was included in the primary emulsion step. The data
in this figure is presented as M ± S.E.M. (n = 3). Source: From Ref. 127.
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
197
carrying the enzymes (127). Nanocapsulation helps to lengthen the therapeutic
window by designing biodegradable polymeric nanocarriers, which protect encapsulated catalase from lysosomal proteolysis, thus prolonging the duration of the
desired effects. They hypothesized that the poly-nanocarriers formation, sizing, and
loading should a potential basis for a more general framework for the formulation
of NPDDSs, especially for enzymes. Many enzymes using small substrates diffusing through polymer shells such as sugars, amino acids, and glutathione may be
amenable to loading into protecting poly-nanocarriers. The study recommended
testing of the delivery vehicles in cell culture and animal studies, as a new strategy
for a prolonged protection against vascular oxidative stress (127).
Weissenbock et al. (129) showed the application of wheat germ agglutinin to
enhance the absorption of PLGA NPs. The wheat germ agglutinin has cyto-adhesive
and cyto-invasive properties. This process of surface engineering the PLGA NPs
by wheat germ agglutinin promises high versatility of application in the search
for biorecognitive ligands enhancing the cyto-adhesion, cyto-invasion, and
probably transcellular transport of colloidal carriers after peroral administration
(129). Gref et al. (130) and Lochner et al. (131) reported novel surface engineering
of the NPDDS.
Proticles: Protamine Nanoparticles as Drug-Delivery Systems
Proticles are novel NPs composed of protamine, a peptide used in many pharmaceutical formulations with DNA, proteins such as albumin and other therapeutically active substances. Junghans et al. (132) reported the use of proticles as delivery
systems for oligonucleotides. The stability of the particles and the oligonucleotides
bound to the proticles was examined in fetal calf serum and cell culture medium.
Proticles significantly decreased cellular growth in a cell proliferation assay using
oligonucleotides against the c-myc proto-oncogene. Proticles can also be used for
diagnostics. NPs can bind large amounts of substances on their surface by adsorption, and so they can be used as an absorber. Proticles with ligand-specific as
amyloid-beta can bind the neurotoxic amyloid-beta protein. The neuro-protective
effects of this delivery system may find a novel therapy for Alzheimer’s disease.
Amyloid-binding peptide depot system is aimed at developing therapy for
Alzheimer’s disease. Amyloid-binding peptides are substances which can disintegrate amyloid plaques. These are hidden in the interior of proticles to cross the BBB
and for slow controlled release to achieve a high concentration of amyloid-binding
peptides in the brain over a long period. This can be further studied using the MRI
with gadolinium as a contrast substance. VIPs, which are used in the treatment of
severe lung diseases, can be packaged in proticles and can be used to create a
pulmonary depot for the treatment for 12 to 24 hours. Antigens are bound to the
surface of proticles for efficient composition of the proticles that might further
boost the immune response. Some recent reports show the application of proticles
as DDSs (133–135).
Mucoadhesive Nanoparticulate Drug-Delivery Systems
Mucosal surfaces are the most common and convenient routes for delivering drugs
to the body. However, macromolecular drugs such as peptides and proteins are
unable to overcome the mucosal barrier and are degraded before reaching the
bloodstream. NPDDSs show a promising strategy for delivering drugs through
mucosa. Polysaccharide Chitosan (CS) is mucoadhesive and CS NPs, CS-coated oil
nanodroplets (nanocapsules), and CS-coated lipid NPs have shown interesting
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Pathak et al.
possibilities for this purpose (136). CS-coated systems have exhibited an important
capacity to enhance the intestinal absorption of the peptide salmon calcitonin and in
vivo, a long-lasting decrease in the calcemia levels observed in rats (137).
Takeuchi et al. (138) have written a review on mucoadhesive NPDDSs for
peptide drugs. They discussed the preparation and methods for evaluation of mucoadhesive nanoparticulate systems. Mucoadhesive properties were conferred on the
systems by coating with mucoadhesive polymers such as CS and Carbapol. The feasibility of this surface adhesion was confirmed by measuring the zeta potential. They
suggested that these delivery systems could be used for delivery of peptides by oral
and pulmonary administration. Nanoprecipitation techniques using PLGA and PLA
polymers were found to be useful for nanoparticulate delivery of proteins and have
shown more versatility and flexibility in the formulation for protein delivery (139).
Takeuchi et al. (140) described mucoadhesive PLGA nanospheres prepared by surface
modification with CS for oral peptide delivery. CS-modified nanospheres were
applied to improve the pulmonary delivery of peptides by nebulization. The particle
average diameter of 650 nm of the aqueous dispersion of the nanospheres was an
important factor to enclose the particles in the aerosolized aqueous droplets produced
with the nebulizer. The elimination rate of CS-modified nanospheres from the lungs
was decreased significantly due to their mucoadhesive property after pulmonary
administration compared to that of the unmodified nanospheres, and as a result the
pharmacological action was significantly prolonged. It is also confirmed that the CS
on the surface of the nanospheres enhanced the absorption by opening the
intercellular tight junctions in the lung epithelium (104).
Nanoparticulate Drug Delivery in Cancer Treatment
Several NPDDSs are reported for the application in cancer therapy, transferring conjugated paclitaxel-loaded NPs (141), nanovaccines (142), adriamycin-loaded NPs for
hepatoma treatment (143), magnetic PBCA nanospheres with aclacinomycin A in
gastric cancer (15), near-infrared absorption nanospheres (144), polypropylenimine
dendrimer NPs for oligonucleotides (145), lytic-peptide-bound magnetite NPs for
breast cancer treatment (146), ceramic-based NPs entrapping water-insoluble photosensitizing anticancer drugs (147), and poly(epsilon-caprolactone) NPs for the
delivery of Tamoxifen for breast cancer treatment (148,149).
Yoo and Park (150) have reported a study of folate receptor targeted anticancer therapy using Doxorubicin-PEG folate nanoconjugates. Doxorubicin and folate
were, respectively, conjugated to alpha and omega terminal of the PEG chain. The
conjugates assisted formation of Doxorubicin nanoaggregates with average size of
200 nm in diameter when combined with an excess amount of depronated
Doxorubicin in an aqueous phase. In vivo studies have shown significant reduction
in the tumor volume in a human tumor xenograft nude mouse model. Controlled
release of paclitaxel through submicroemulsion with particle size of 45 to 270 nm
was evaluated in vitro and in vivo for their antitumor activity by Kang et al. (151).
They used PLGA polymer for formulating self-emulsifying DDSs and showed the
effectiveness of the system.
Nanoparticles in the Treatment of Vascular Thrombosis
The formation of blood clots in the circulatory system is associated with a range of
serious medical conditions, including heart attacks, pulmonary embolisms, strokes,
and deep vein thrombosis. The main component of the clot is the insoluble protein
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
199
fibrin. Treatment of vascular thrombosis involves the use of thrombolytic drugs that
break up the fibrin, allowing the clot to disperse. Biocompatible NPs are used to
develop such delivery systems, which can carry the thrombolytic drugs. Chellini
(152) explained that the thrombolytic drugs are powerful agents, with serious side
effects like causing hemorrhage if they are given systemically. However, orally they
are less efficient. If they can be incorporated in NPs, they can be delivered directly
to the specific site, using less drug materials, and the treatment will be cost-effective
with less side effects. The drug will be released from the NPs by diffusion, degradation, or erosion. A sustained release NP formulation may be more helpful (152).
Nanoparticulate Drug-Delivery Systems for Gene Therapy
Intracellular gene delivery involves changing the expression of genes in order to
prevent, cure, or treat a disorder or a disease. Therefore, this treatment method
alters the expression of a gene and corrects a defective gene that may be the cause of
a disease or a disorder. Nonviral vectors for gene therapy have inherent advantages
of safety and flexibility over viral vectors, although they are less efficient. Serious
issues of integration with the host genome to permanently alter its genetic structure,
self-replication capability, recombination potential, and the possibility of complement activation (immunogenicity) of the otherwise transfection efficient viral
vectors limit their use for gene delivery. In the last decade, the focus of development
is on nonviral gene-delivery systems. Specific characteristics that must be included
in nonviral vectors include small size and stability against aggregation in blood,
serum, and extracellular fluid, the ability to be efficiently internalized by the target
cells and the ability to disassemble and release the payload into the cell nucleus
once internalized.
Gene therapy has attracted considerable interest for the treatment of lifethreatening diseases. Several viral and nonviral vectors (transporting devices) are
under investigation (153,154). One of the common disadvantages of both types is
rapid clearance from the blood circulation to first-pass organs such as liver and
lungs. A frequently applied strategy to circumvent this rapid elimination is to coat
the outer surface of the complexes with hydrophilic uncharged polymers by which
the positive charge of these lipo/polyplexes is shielded (155–157). Several polymers
such as poly-l-lysine, CS (158) polyamidoamine dendrimers (159), polyethylenimine
(PEI) (160), poly(4 aminobutyl-l-glycolic acid) (161), poly-l-glutamic acid (162),
poly(β-amino esters) (163), poly(2(dimethylamino)ethyl methacrylate), plasmidlipid particles have been used for this purpose with PEG coating (161,162,164).
Funhoff et al. (165) have reported the use of PEG-shielded double-layered micelles
for gene delivery. Zhang et al. (166) reported galactosylated ternary DNA/polyphosphoramidate NPs as hepatocyte-targeted gene carriers. Zhao et al. (167–170)
wrote an excellent review on polyphosphoesters in drug and gene delivery. A report
described the application of PLGA:poloxamer and PLGA:poloxamine blend NPs as
new carriers for the gene delivery: plasmid DNA (171). They developed several formulations and found that all NP formulations provided continuous and controlled
release of the plasmid with minimal burst effect. In addition, the release rate and
duration were dependent on the composition of the particle matrix. Zahr et al. (72)
have created a layer-by-layer stepwise self-assembly of the polyelectrolytes
poly(allylamine hydrochloride). Poly(styrenesulfonate) was used to create a
macromolecular nanoshell around drug NPs (approximately 150 nm in diameter).
Dexamethasone was chosen as a model drug. The polymeric nanoshell on the
200
Pathak et al.
surface of the drug NPs provides a template upon which surface modifications can
be made to create stealth or targeted DDSs (172).
Kaul et al. (173) in their study used PEG-modified gelatin NPs as long-circulating
intracellular delivery systems with a mean particle size of 300 nm. They could efficiently encapsulate hydrophilic macromolecules including plasmid DNA. These
particles were internalized by tumor cells and were found near the nucleus after 12
hours. The PEGylated gelatin NPs were also very efficient in expressing GEP. Drug
loading and drug-release rates from NPs are important parameters for the formulation
of NPDDSs to optimize the therapeutic efficacy of the encapsulated drug (174–178).
Panyam et al. (178) showed that solid-state solubility of the drug in polymer could be
an important determinant that could influence the drug loading in NPs as well as
release characteristics of the encapsulated drug. In a recent study by Kaul et al., plasmid DNA was formulated with gelatin and PEGylated gelatin with average diameter
of 200 nm for targeted systemic delivery to solid tumors. The results showed 61%
transfection efficiency, which was attributed to a biocompatible, biodegradable
long-circulating carrier system (179).
Dendrimers are emerging as a new generation of nonviral vectors for gene
delivery because of two distinctive features: their structures can be controlled and
their chemistry can be adapted for various requirements such as drug or gene delivery (180,181). They also belong to the polyamine group of nonviral vectors, which
includes poly-l-lysine, polyethylenimine, and polyamidoamine. The versatility of
these vectors has been exploited to attach various ligands such as transferrin, sugars,
and antibodies for receptor-targeting dendrimers. A new family of composite dendrimers with lipidic amino acids was synthesized by Bayle et al. and used for transporting DNA both single- and double-stranded as well as RNA (182). Polyamidoamine
dendrimer (PAMAM) represents one of the most efficient polymeric gene carriers.
Zhang et al. (183) have shown their utility for gene delivery.
Xia et al. (177) report a novel DDS recently by preparing monodispersed NPs
consisting of interpenetrating polymer networks (IPNs) of polyacrylic acid (PAAc)
and isopropylacrylamide by a seed and feed method. The aqueous dispersion of IPN
NPs was found to be a unique NPDDS due to its abrupt inverse thermoreversible
gelation at around 33°C. The IPN and drugs were thoroughly mixed as an aqueous
solution at room temperature and formed a drug-delivery gel at body temperature.
The drug-delivery model was found very useful because such a dispersion and drug
was mixed without chemical reaction and the liquid can be injected into a body to
form in situ a gelled drug depot to release the drug slowly (177).
Benita et al. (184) described the application of PLGA–PEI NPs for gene delivery
to pulmonary epithelium. Diwan et al. (185) have shown that antigen delivery in
biodegradable NPs can facilitate induction of strong T-cell response, particularly of
the TH1 type, at an extremely low dose of CPG oligonucleotides. Such reduction in
dose would be advantageous for minimizing the potential side effects of these novel
adjuvants (185).
Rhaese et al. (186) reported NPs consisting of DNA, human serum albumin, and
polyethylenimine to be a carrier for the nonviral gene delivery. They could achieve
optimum transfection efficiency, and displayed a low cytotoxicity when tested
in cell culture. They recommended these carriers for delivering DNA for IV
administration.
An interesting strategy for the treatment of various vascular diseases uses
poly (methylidene malonate 2.1.2) NPs which is a biocompatible polymer that
enhances the peri-adventitial adenoviral gene delivery (187).
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
201
Gene therapy strategies have been proposed for a vast and diverse range of
disorders for which currently available treatments are deemed unsatisfactory.
Effective delivery of genes into cells has been considered a major hurdle in achieving
successful gene therapy. A number of delivery systems based on viral (188) or nonviral vectors (189,190) have been devised; none of them have proven to be completely
satisfactory. Viral vectors can only introduce genes, not other macromolecules such
as siRNA or antisense nucleotide. Furthermore, side reactions such as host immune
response and insertional mutagenesis leading to death, carcinogenesis, or germline
alternations have led to serious concerns about the use of viruses as gene transfer
vectors (191–194). Several analytical techniques were reported to be useful for the
characterization and establishing the structure/function relationship of polyamidoamine/DNA dendrimers as nanoparticulate drug/gene-delivery systems. Braun et
al. (195) used dynamic and electrophoretic light scattering technique for particle size,
and phase analysis light scattering for zeta potential. Ethium bromide displacement
assay has been utilized for determining the interaction between the dendrimer and
DNA, and extent of gene uptake. Circular dichroism spectroscopy was used for characterizing helical structure of DNA within dendrimer DNA complexes. Fourier
transform infrared spectroscopy was used as a complimentary technique to further
investigate the secondary structure of DNA component complexes. Isothermal titration calorimetry was employed to investigate the thermodynamics of binding of
dendrimers and DNA complexes. Differential scanning calorimetry was applied to
evaluate the thermal stability of DNA/dendrimer complexes (195). Zaitsev et al. (96)
used a strategy based on the formation of polyelectrolyte NPs and later deposition of
negatively charged polyelectrolytes onto a DNA core. They showed that these
negatively charged particles exhibited colloidal stability and high transfection
efficiency in an in vivo model (196).
The mechanistic pathways for gene expression are limited by at least five
major barriers: in vivo stability, cell entry, endosome escape, cytosolic transport, and
nuclear entry. The nuclear membrane restricts the transport of the plasmid DNA
and the efficiency of the DNA transfer from the cytoplasm to the nucleus has been
estimated to be about 10 −4. In order to obtain the high gene expression, the genes
introduced into the cells must be reduced to a compact size which can pass through
nuclear pores. The collapsing of the DNA into NPs of reduced negative or increased
positive charges (i.e., DNA condensation) has received considerable interest due to
its biological importance in DNA packaging in virus heads (197).
Gene delivery usually takes advantage of the endocytic pathway of the cell.
The cells continually ingest a part of their plasma membrane via endocytosis to
form endocytic vesicles. The cells can take up solvent and solute by the endocytosis
activity later. Endocytic vesicles incorporating DNA-bearing particles are transferred to endosomes and then to lysomes, where liberation of DNA somehow takes
place. However, the plasma membrane is directly linked to and functionally
integrated with the underlying cytoskeleton; the endocytosis at the membrane
would require the rearrangement of actin and tubulin.
Nonviral vectors hold several advantages over modified virus employed
for gene delivery, in terms of improved and predictable safety profile, a high
DNA-carrying capacity, increased versatility, the ease of large-scale production,
and quality control. Nevertheless, their efficiency lags behind that of viral systems.
Those nonviral vectors efficient in transfection are often toxic because of their nondegradable property (189). The gene transfer efficiency of nontoxic vectors, for
example, biodegradable cationic polymers, is often not satisfactory (198). Li et al.
202
Pathak et al.
have reported the synthesis, characterization of poly(d,l-lactide-co-4-hydroxy-lporline) polymer for the purpose of gene delivery, and they studied degradation,
cytotoxicity, as well as pDNA release kinetics and sustained gene expression of this
polymer-based system. They showed the usefulness of these polymers with
multiple advantages (190).
Gupta and Gupta (191) have shown the application of the Pullulan, a watersoluble, neutral linear polysaccharide for gene delivery. They showed that these
NPs had high transfection potential, could release DNA efficiently, and were stable
against degradation by DNAse. As specific ligands can be bound to the NP surface,
these particles offer the possibility for additional targeting strategies.
Kabanov et al. (199) suggested an interesting approach using polymer genomics
in their recent publication. Pluronic, the A-B-A amphiphilic block copolymers of
poly(ethylene oxide), can upregulate the expression of selected genes in cells and alter
the genetic response to antineoplastic agents in cancer. They reported that these block
copolymers alone as well as in combination with polyethylenimine can upregulate
the expression of the reporter genes in stably transfected cells. This underscores the
ability of selected synthetic polymers to enhance transgene expression through a
mechanism that augments improved DNA delivery into cell. Pluronic is genetically
benign when combined with an antineoplastic agent Doxorubicin. It drastically alters
pharmacogenomic responses to this agent and prevents the development of multidrug resistance in breast cancer cells. They proposed the need for a thorough
assessment of pharmacogenomic effects of polymer therapeutics to maximize the
clinical outcomes and understand the pharmacological and toxicological effects of
polymer-based drugs and delivery systems (199).
Tabatt et al. (200) compared the cationic SLN and liposomes for gene transfer
as a carrier. They found that DNA binding differed only marginally in these two
systems. They concluded that cationic lipid composition seems to be more dominant
for the in vitro transfection performance than the kind of arranged colloidal structure. Hence, cationic SLN extends the range of highly potent nonviral transfection
agents by one with favorable and distinct technological properties (200). Wissing et
al. (201) published a review which describes the use of NPs based on solid lipid for
the parenteral application of drugs. Different types of NPs based on solid lipids such
as SLNs, nanostructured lipid carriers, and lipid drug nanoconjugates are introduced
and structural differences are pointed out. Different production methods including
the suitability for large-scale production are described along with the stability issues,
and drug-incorporation mechanisms into the particles are discussed in detail. The
biological activity of parenterally applied SLNs and biopharmaceutical aspects such
as pharmacokinetic profiles as well as toxicity aspects are reviewed (201).
NANOPARTICULATE SYSTEMS: KNOWN AND UNKNOWN RISKS
An excellent review was recently published by Hoet et al. (202) with an extensive
literature survey in which they suggested that the particles in the nanosize range
could certainly enter into the human body through lungs, gastrointestinal system,
mucosa, and the skin. It is possible that some particles penetrate deep in the dermis
and gradually may be taken up by the body. The chances of penetration depend on
the size and surface properties of the particles and also point of contact. The distribution of particles in the body systems is also a function of the surface characteristics of the NPs. There might be a critical size involved beyond which the movement
of the particles might be restricted. The pharmacokinetic behavior of different types
Pharmaceutical Applications of Nanoparticulate Drug-Delivery Systems
203
of the nanosystems requires a detailed investigation and a database of health risks
associated with different NPs needs to be created. The increased risk of cardiopulmonary disease requires specific measures to be taken for every newly developed
nanoparticulate product. There is no universal NP to fit all the cases; each NP system
needs to be treated individually when a health risk is expected. The tests currently
used to verify safety of materials should be applicable to identify hazardous
NPs, and more stringent and efficient testing procedures are needed to evaluate the
nanoparticulate systems, especially when used as a food component or as DDSs
(119,203–206). A study reported by Bilati et al. (207) discussed the processing
and formulation issues related to PLGA protein-loaded NPs prepared by doubleemulsion method. They evaluated the effect of some typical formulation factors and
processing conditions on the mean size and the drug entrapment efficiency of PLGA
NPs. They found that the parameters that generally increase the entrapment
efficiency are (i) high molecular weight of the polymer, (ii) the presence of uncapped
carboxylic end groups when PLGA is used, (iii) the use of methylene chloride
instead of ethyl acetate, and (iv) an increased nominal drug loading. An interesting
study regarding the GI uptake and transport of SLN to the lymphatic system was
reported by Bargoni et al. (208). They showed that particle size is a critical determinant of the fate of NPs administered orally; larger particles may be retained for
longer time in Peyer’s patches, whereas smaller particles are transported to the
thoracic duct. Another study by Passirani et al. (209) reported that NPs bearing heparin or dextran covalently bound to polymethylmethacrylate was found to be in the
circulation for a long time. The potent capacity for opsonization of the polymethylmethacrylate core was hidden by the protective effect of either polysaccharide. In
the case of heparin NPs, the stealth effect was probably increased by its inhibiting
properties against complement activation. Silveira et al. (210) reported a new type
of NPs where they used polyisobutylcyanoacrylate and cyclodextrin combination
as polymeric carrier. Owing to the presence of many lipophilic sites belonging to the
cyclodextrins which were firmly anchored to the structure of the NPs, these types of
carriers were very useful to enhance and increase the loading of the lipophilic drugs
with probable less side effects. But, the risks involved need to be verified and
established by appropriate sensitive methods to ensure the safety of the NPDDS.
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14
Lipid Nanoparticles (Solid Lipid
Nanoparticles and Nanostructured Lipid
Carriers) for Cosmetic, Dermal, and
Transdermal Applications
Eliana B. Souto and Rainer H. Müller
Department of Pharmaceutical Technology, Biotechnology, and
Quality Management, Freie Universität Berlin, Berlin, Germany
INTRODUCTION
The skin, together with the mucosal tissues of the digestive, respiratory, and
urogenital tracts, forms the frontier of the body, separating the external environment from internal organs. The epidermis forms the barrier between those “worlds,”
that is, between the most varied conditions of temperature and humidity of the
external environment, in contrast to the very stable internal environment of the
living tissues and body fluids. Thus, epidermis behaves as the protective organ
showing functions of physical protection of the body, sensation, and temperature
control, as well as physical barrier against the exchange of chemical substances into
and out of the human organism.
These days, the pharmaceutical industry supplies the costumer with many
physicochemically different dosage forms intended for skin application, ranging
from powders to liquids, as well as semisolid formulations. When developing such
formulations, the pharmaceutical technologists must have in mind many concerns,
such as stability, compatibility, and costumer acceptability of the vehicle and active
ingredient, as well the bioavailability of this latter in cases of dermal and transdermal treatment (1).
It is a fact that preparations for topical, dermal, and transdermal use have the
unique feature that their physical properties are almost as important as any pharmacologically active ingredient that they contain. The composition of the vehicle
and active compounds deserves special emphasis, particularly because the intimate
contact with the skin is always accompanied by possible risks of adverse reactions.
The treatment of skin lesions is often conducted conservatively, mainly intended to
soothe the skin during the progress of natural tissue repair. The principles of formulation closely resemble those applicable to cosmetic products and in both cases a
good deal of caution is necessary to ensure the success of the therapy.
Concerning dermatological and transdermal therapy, the precise quantity of
active ingredient which is necessary to achieve a given response cannot be predicted
with certainty. This lack of precision is largely due to variability of skin penetration
by the active ingredient, which is related to the thickness of the epidermis and its
keratin layer, as well as to mechanical removal of the applied formulation if the
affected area is not covered by a dressing.
To overcome many of these above-mentioned drawbacks, attempts have been
made to introduce lipid nanoparticles into the cosmetic and pharmaceutical fields.
213
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Souto and Müller
Solid lipid nanoparticles (SLNs) (2) and nanostructured lipid carriers (NLC) (3) are
novel colloidal delivery systems with many cosmetic and dermatological features,
such as adhesive properties when applied to the skin (4). These properties bring
many other advantages such as occlusion and skin hydration, absorption-increasing
effects, active penetration enhancement, and controlled-release properties (4,5).
SLN and NLC systems differ because SLNs were developed by exchanging the
liquid lipid (oil) of oil-in-water (o/w) emulsions by a solid lipid (6), which can
bring many advantages in comparison to a liquid core (7). The concept of NLC was
developed by nanostructuring the lipid matrix, in order to give more flexibility for
modulation of drug release, increasing the drug loading and preventing its leakage.
This approach was accomplished by mixing solid lipids with liquid lipids (NLC
concept), instead of highly purified lipids with relatively similar molecules (SLN
concept). The result is a less ordered lipid matrix with many imperfections, which
can accommodate a higher amount of drug (5,8–10). The mixture obtained with
solid and liquid lipids needs to be solid at least at 40°C to make sure that it does not
melt at room and body temperature if used for drug delivery.
Advantages of using lipids as carrier systems for skin administration are
related to their physiological and well-tolerated nature, which reduces the risk of
toxicological problems and local irritancy (5,6,11). A range of very different lipids
with generally recognized as safe status has been used to produce SLNs and NLCs,
ranging from highly purified lipids, for example, tristearin (12) in SLNs, to mixtures
of mono-, di-, and triacylglycerols in NLCs, including monoacid and polyacid
acylglycerols.
When transforming the bulk lipid into a nanoparticulate form (i.e., SLNs or
NLCs), a melting point depression is, in general, observed. This melting point
depression is described in the Thomson equation which itself is derived from the
Kelvin equation (13). An additional melting point depression occurs when a foreign
compound (active ingredient or surfactant molecule) is dissolved in the lipid matrix,
for example, surfactant will partition from the water phase to the lipid phase,
and active-loaded lipid nanoparticles show a melting point depression when a
molecular-dispersed active ingredient is present.
When comparing the lipid materials as bulk and as nanoparticulate systems,
melting and recrystallization temperatures can be different, especially in chemically
polydispersed lipids. When producing lipid nanoparticles, the difference can be
between 10°C to 20°C or even more. For example, lipids, such as trilaurin (C12),
which has an original melting point of 43°C to 47°C, show a very pronounced supercooling when formulated as lipid nanoparticles. In such compounds, crystallization
will not take place anymore at room temperature, and it will only happen if cooled
down to freezing temperatures (14).
Lipid molecules show different three-dimensional structures, that is,
polymorphic forms: unstable α, metastable β′, and as the most stable the β modification (Table 1).
In cases of mixtures of acylglycerols, polymorphic transitions occur from β′ to
βi and then to β, which means that another intermediate polymorph form is present.
Most bulk lipids are obtained in β modification or at least predominantly in β modification. In many cases, the production of lipid nanoparticles leads to a change in the
relative fraction of the polymorphic forms. Depending on the chemical nature of the
lipid and on the parameters established for the production of lipid nanoparticles,
different fractions of α and β′ modification will be obtained. This can lead again to
a reduction of the melting point, or more precisely change in form and shift of the
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
215
TABLE 1 Three-Dimensional Structure of the Crystal Order in the Three Main
Polymorphs from Monoacid Triacylglycerols
α modification
Crystal system
Subcell
Hexagonal
Orthorhombic
β′ modification
Monoclinical
Orthorhombic
β modification
Triclinical
Triclinical
Source: From Refs. 15, 16.
melting peak. In contrast, it might be that the created polymorphic forms are not
long-term stable, that is, there is a gradual transformation to more stable modifications, which means increasing the content of β′/βi and finally β. This phenomenon
is not desired and the changes in lipid structure can cause drug expulsion during
storage and changes in the release profile of the incorporated active molecules. To
avoid such undesired effects, lipid mixtures should be chosen which transform relatively fast to more stable modifications directly after particle production (i.e., within
one to three days). It is also perfectly acceptable or it can even be planned to trigger
drug release, when the generated fractions of the different polymorphic forms
remain unchanged during storage and do not transform back to the β modification
of the bulk lipid.
Drug expulsion from lipid matrices is a well-known phenomenon from
suppositories (17). Concerning the different possibilities for how the active ingredients can be incorporated within the lipid nanoparticle matrix, one can mention: (i)
the replacement of host molecules in the lattice by a guest molecule or (ii) its incorporation in between the host molecules. However, for this latter hypothesis the
guest molecule needs to have a size less than 20% of the host molecule. Active molecules can be localized in between the lipid lamellae which then results in an increase
of the lattice spacing d (i.e., interatomic distance defined by the Bragg’s equation;
see Ref. 18). There is also the possibility that the active ingredient is present in the
form of amorphous clusters, mainly localized in the imperfections of the crystal. In
general, the accommodation of the active molecules is improved with the increase
of the number of imperfections in the lipid crystals (19). Active loading can be
increased by using rather crude lipid mixtures, such as the ones used in cosmetic
products. An even further improvement of active loading can be achieved by
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controlled nanostructuring of the lipid matrix, that is, creating as many imperfections as possible. Depending on the nature of the lipids used for blending and the
lipid matrix, different types of NLCs will be obtained. The crystalline nature of the
lipid matrix is, therefore, the dominant factor determining the active loading, the
release profile pattern, and also the long-term stability of SLNs and NLCs regarding
the release and subsequent bioavailability of the active ingredient. Thus, the
crystalline status needs to be closely monitored when developing a SLN or NLC
formulation.
HISTORICAL BACKGROUND AND PRODUCTION OF SOLID LIPID
NANOPARTICLES AND NANOSTRUCTURED LIPID CARRIERS
The first lipid particles were produced by the research group of Speiser, the father of
the nanoparticles, in Zurich (20) ( J. Paris Match). High-speed stirring was applied
for the production of an o/w emulsion between a melted lipid phase and a hot
aqueous surfactant solution. The obtained emulsion was then cooled and the inner
lipid phase formed solid particles. However, in general, the use of stirring techniques
has some drawbacks, such as the relatively broad size distribution of particles and
the fact that relatively high concentrations of surfactant molecules are usually
required to obtain a mean particle diameter in the nanometer range. The product
developed by Speiser was called “lipid nanopellets” and was intended for oral
delivery. The patent obtained by Speiser was not followed up by the owner
Rentschler, and patent protection no longer exists in a number of countries.
A similar process was developed and patented by Domb, who prepared particle dispersions applying a sonication procedure. The product developed by
Domb was called “lipospheres” (21), and they also found no broad application in
pharmaceutical products.
In 1991, the patent application of the first generation of lipid nanoparticles—
SLNs—was submitted describing the nanoparticle production by high-pressure
homogenization (HPH) (22) and also via microemulsion technique (23).
HPH is a technique broadly used in different research areas, and it is also
established in pharmaceutical production, for example, for the production of emulsions for parenteral nutrition (Intralipid®, Lipofundin®, Lipovenoes®) (24–26). Using
this technique, the problems of other nanoparticles (polymeric nanoparticles), such
as the lack of scaling up and large-scale production lines, has been overcome.
In addition, HPH leads to a product being relatively homogeneous in size, that is,
possessing a higher physical stability of the aqueous dispersion; in general,
polydisperse dispersions show a greater tendency to aggregate or coalesce.
Furthermore, it is a simple and very cost-effective production technique. There are
basically two different production methods: the hot and the cold HPH techniques.
For the hot HPH technique, the lipid is melted at approximately 5°C to 10°C
above its melting point, the active ingredient is dissolved or finely dispersed in the
melt and then the active-ingredient containing lipid melt is dispersed by stirring in
a hot surfactant solution. The obtained preemulsion is homogenized applying a
pressure between 200 and 500 bar and two to three homogenization cycles. After the
homogenization, a hot nanoemulsion is obtained; cooling leads to recrystallization
of the lipid and formation of lipid nanoparticles.
For the cold HPH technique, the lipid melt containing the active ingredient is
cooled, and after solidification is ground using a mortar mill. The obtained lipid
microparticles are further dispersed in a cold aqueous surfactant solution. The
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
217
resulted presuspension is homogenized in the solid state at or below room temperature by cooling the high-pressure homogenizer. The sheer forces and cavitation
forces in the homogenizer are strong enough to break the microparticles directly
into lipid nanoparticles.
For the production of lipid nanoparticles via the microemulsion technique,
the lipid is melted, the surfactant, cosurfactant, and water are added in such concentrations that a microemulsion results (27). The size of the microemulsion region
in the phase diagram is a function of temperature, that is, the microemulsion can be
converted to a different system when, for example, reducing the temperature.
Therefore, for particle production, the microemulsion needs to be kept at the
elevated temperature during the process. The hot microemulsion is diluted into
cold water, leading to a “breaking” of the microemulsion and subsequent formation of an ultrafine nanoemulsion. The dilution with water and the reduction of
temperature narrowing the microemulsion region are the reasons for breaking of
the microemulsion (28). One disadvantage of this procedure is the dilution of the
particle suspension by water, obtaining concentrations usually below 1% of particle
content. When processing to a final dosage form, a very large amount of water
needs to be removed.
Other approaches for the production of lipid nanoparticles have been adapted
from polymeric nanoparticle production procedures, for example, the solvent
emulsification-evaporation method described by Sjöström and Bergenståhl (29),
the solvent displacement method described by Fessi et al. (30), and the emulsification-diffusion technique patented by Quintanar-Guerrero et al. in 1999 (31). The
novel phase-inversion-based technique has been described by Heurtault et al.
(32,33) for the production of SLNs.
The solvent emulsification-evaporation is a method analogous to the production of polymeric nanoparticles and microparticles by solvent evaporation in o/w
emulsions. The lipid is dissolved in an organic solvent which shows no miscibility
with water [e.g., cyclohexane (29,34), chloroform (34), or methylene chloride (35,36)].
The organic solution is dispersed in aqueous surfactant phase, the solvent is then
removed by evaporation (34). It can be applied for the incorporation of hydrophilic
molecules such as peptides and proteins, which must be previously dissolved into a
water phase preparing in this case a water-in-oil-in-water (w/o/w) emulsion (36).
In the solvent displacement method, an organic solvent which is miscible with
water is used. In this case, the lipid material is previously dissolved in a semipolar
water-miscible solvent, such as ethanol, acetone, or methanol (37–40).
In the emulsification-diffusion technique, benzyl alcohol (41) or tetrahydrofuran (42), which is previously saturated with water, is used to ensure the initial
thermodynamic equilibrium between the two liquids (water and solvent). Then, the
water-saturated organic solvent is used to dissolve the lipid, and after this an o/w
emulsion is prepared. Owing to the saturation of the organic solvent with water,
no solvent diffuses from the droplets into the external water phase from the o/w
emulsion. Removal of solvent from the droplets and particle formation is achieved
by adding additional water to the emulsion and extracting the solvent (41).
A very interesting approach based on the physics behind it is the phaseinversion-based technique, which is a two-step method. First, all components are
placed on a magnetic stirrer using a temperature program from room temperature
to, for example, 85°C. This is followed by progressive cooling to 60°C. Three
temperature cycles (85–60–85–60–85°C) are applied to reach the inversion process
defined by temperature range. In step 2, an irreversible shock is induced by dilution
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Souto and Müller
with cold water. This fast cooling–diluting process leads to the formation of stable
nanoparticles (43).
MORPHOLOGY AND STRUCTURE OF SOLID LIPID NANOPARTICLES
AND NANOSTRUCTURED LIPID CARRIERS
Different models have been described in the literature for how active molecules can
be incorporated into SLNs and NLCs (4). For each of the carriers, three basic types
are described (Fig. 1).
The type of SLNs depends on the chemical nature of the active ingredient and
lipid, the solubility of actives in the melted lipid, nature and concentration of
surfactants, type of production (hot vs. cold HPH), and the production temperature.
Therefore, three incorporation models have been proposed (45):
1. SLN Type I or homogeneous matrix model,
2. SLN Type II or drug-enriched shell model, and
3. SLN Type III or drug-enriched core model.
The SLN Type I or homogeneous matrix model is derived from a solid solution of
lipid and active ingredient. A solid solution can be obtained when SLNs are
produced by the cold homogenization method. A lipid blend can be produced
containing the active in a molecularly dispersed form. After solidification of this
blend, it is ground in its solid state thus avoiding or minimizing the enrichment of
active molecules in different parts of the lipid nanoparticle.
The SLN Type II or drug-enriched shell model is achieved when SLNs are
produced via the hot HPH technique, and the active ingredient concentration in
the melted lipid is low. During the cooling process of the hot o/w nanoemulsion,
the lipid will precipitate first, leading to a steadily increasing concentration of
active molecules in the remaining lipid melt with increasing fraction of lipid solidified. An active-free lipid core is formed; when the active reaches its saturation
solubility in the remaining melt, an outer shell will solidify containing both active
FIGURE 1 (See color insert.) Basic types of solid lipid nanoparticles and nanostructured lipid carriers.
Abbreviations: NLC, nanostructured lipid carrier; SLN, solid lipid nanoparticle. Source: From Ref. 44.
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
219
and lipid. The enrichment in the outer area of the particles causes burst release. The
percentage of active ingredient localized in the outer shell can be adjusted in a
controlled way by altering the production parameters. A typical example of an
active-enriched shell model is the incorporation of coenzyme Q10 (46,47).
The SLN Type III or drug-enriched core model can take place when the active
ingredient concentration in the lipid melt is high and at or relatively close to its
saturation solubility. Cooling down of the hot oil droplets will in most cases reduce
the solubility of the active in the melt; when the saturation solubility is exceeded,
active molecules precipitate leading to the formation of a drug-enriched core.
Regarding the models described for NLCs, Figure 1 also shows three different
structures:
1. NLC Type I or imperfect crystal model,
2. NLC Type II or amorphous model, and
3. NLC Type III or multiple model.
NLC type I is defined as the imperfect crystal model, because once in its matrix
there are many imperfections which are able to accommodate the active molecules.
This model is obtained when mixing solid lipids with small amounts of liquid lipids
(oils). Owing to the different chain lengths of the fatty acids and the mixture of
mono-, di-, and triacylglycerols, the matrix of NLCs is not able to form a highly
ordered structure (5).
NLC type II is called the amorphous model because it is created when mixing
special lipids which do not recrystallize anymore after homogenization and cooling,
such as hydroxyoctacosanylhydroxystearate and isopropylmyristate. These lipids
are able to create solid particles of amorphous lipid structure, which can avoid the
occurrence of crystallization, minimizing drug expulsion, because the matrix is
maintained in the polymorphic α form.
NLC type III is described as the multiple model. This model has been
developed to improve the loading capacity of several drugs, such as the ones whose
solubility in liquid lipids is higher than in solid lipids (48,49). This type is derived
from w/o/w emulsions, which consist of an oil-in-fat-in-water dispersion. Very
small oil nanocompartments are created inside the solid lipid matrix of the nanoparticles generated by a phase-separation process (5). This model is obtained when
mixing solid lipids with liquid lipids (oils) in such a ratio that the solubility of the
oil molecules in the solid lipid is exceeded. The melted lipid and the hot oil are
blended; thus, the two lipids must show a miscibility gap at the used concentrations, at approximatel 40°C. A hot o/w nanoemulsion is produced at a higher
temperature (approx. 80°C), then the lipid droplets are cooled. When reaching the
miscibility gap, the oil precipitates forming tiny oil droplets in the melted solid
lipid. Subsequent solidification of the solid lipid as solid nanoparticle matrix leads
to fixation of the oily nanocompartments.
DEVELOPMENT OF COSMETIC AND TRANSDERMAL
PRODUCTS BASED ON SOLID LIPID NANOPARTICLES
AND NANOSTRUCTURED LIPID CARRIERS
The scientific literature reports several approaches for the development of products
based on SLNs and NLCs (4,5).
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Incorporation of Solid Lipid Nanoparticles and Nanostructured Lipid
Carriers into Semisolid Preparations
Similar to liposomes, polymeric nanoparticles, and microsponge systems, aqueous
SLN or NLC dispersions can be added to semisolid preparations such as lotions,
creams, and hydrogels (50,51). The advantage of this procedure is the association of
a well-established topical formulation with several attractive advantages of the lipid
nanoparticles in the same final product.
The incorporation of SLNs or NLCs into a cream consists of the addition of the
lipid nanoparticles as a highly concentrated dispersion, that is, with 50% solid (particle) content to a freshly prepared o/w cream or during the production of such cream
(49,52). In the first case, a part of the water in the cream formulation is replaced by a
highly concentrated SLN or NLC dispersion, and after that the production process is
run normally. The lipid nanoparticles are sufficiently stabilized to avoid their coalescence with the inner oil droplets of the emulsion. If the production process of the
emulsion is performed at a temperature higher than the melting point of the lipid
nanoparticles, these latter will melt but will recrystallize during the cooling at the
end of the process. The second approach is more elegant, but in this case the cream is
produced as usual, however with a reduced water content in order to compensate for
the water added with the aqueous SLN or NLC dispersion. After the production of
the cream, the concentrated SLN or NLC dispersion is added by stirring at room
temperature. This process avoids the melting of the nanoparticles, avoiding therefore
undesired changes within the internal particle structure.
If the aim is the addition of SLNs or NLCs to hydrogels, the procedure is even
simpler. The hydrogels can be previously prepared, and after that SLN or NLC dispersion is diluted within the semisolid formulation (50). Alternatively, a concentrated
SLN or NLC dispersion is added before the gelation process. However, it should be
noted that electrolytes frequently used to jellify synthetic polymers can destabilize the
lipid nanoparticles. The surface electrical charge (zeta potential) is reduced leading to
aggregation of suspended particles. This phenomenon should be considered when,
for example, adding electrolytes in the form of sodium hydroxide for the preparation
of carbomer gels. The type of neutralizing agent affects the aggregation of SLNs and
NLCs, and it can be avoided or minimized when using Tristan® and Neutrol® TE as
neutralizing agents (49). As soon as the gel is formed, aggregation will not occur
anymore because SLNs and NLCs will be physically stabilized and entrapped in the
gel network. In general, SLN and NLC dispersions, which are unstable because of
suboptimal surfactant combination, can be further stabilized after their incorporation
into hydrogels. This opens the opportunity to use stabilizers, which are not that
efficient in providing a long-term stability of the aqueous dispersion, but are regulatory accepted. The SLN and NLC dispersions need to be stable only for the few hours
until they are incorporated into the semisolid formulation, where they will be
stabilized by the gel network.
Production of Gels Based on Solid Lipid Nanoparticles and
Nanostructured Lipid Carriers
The production of gels based on SLNs and NLCs consists of topical formulations
having only lipid nanoparticles in their composition. A lipid nanoparticle dispersion is previously prepared, and then the gelling agent is added, preferentially
nonelectrolyte agents, such as cellulose derivatives (49) or other natural gums (53).
For many gel preparations, an amount of 4% to 10% of lipid nanoparticles is
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
221
sufficient. To make production more profitable, a concentrated particle suspension
(e.g., 40%–50%) can be prepared and then diluted to the desired final concentration.
This approach has certain advantages, such as the release of actives being controlled
only by the lipid nanoparticle matrix. The desired properties can be adjusted in a
very controlled way by attenuating the lipid nanoparticle features.
Production of Creams Based on Solid Lipid Nanoparticles and
Nanostructured Lipid Carriers
The production of creams based on SLNs and NLCs consists of the production of
aqueous dispersions of lipid nanoparticles in a high concentration, up to 50% or
60%. When applying higher lipid concentrations, ointments with a bicontinuous
structure will be formed, that is, the homogenization process leads to well-defined
particles and not to an ointment-like system.
It is known that the viscosity of the preparations increases with the lipid concentration (51), that is, at about 40% to 45% the formulations are cream-like, above
50% they become paste-like, and when increased to a solid content of 80% or 90%
the formulations are solid and can be cut with a knife. Cream-like and paste-like
formulations are suitable for dermal application of actives, whereas the solid-like
systems are usually of high interest for oral administration of lipid particles to
exploit the absorption-enhancing effect of lipids (54). For the production of such
formulations, a highly concentrated stock suspension (50%–60% lipid nanoparticles) is first prepared by HPH. Then, additional melted lipid is added stepwise and
dispersed again, building a concentration of up to 80% or 90% of solid content. In
the melted state, the obtained product is still liquid and it has a relatively low viscosity. After cooling down, the system becomes cream-like, paste-like, or solid-like.
FEATURES OF SOLID LIPID NANOPARTICLES AND
NANOSTRUCTURED LIPID CARRIERS AND SKIN EFFECTS
Physical Stability in Aqueous Dispersions and in Creams
Physical stability of the aqueous SLN and NLC dispersions, that is, the absence of
particle aggregation and creaming, is a prerequisite for the formulation of cosmetic
and pharmaceutical products based on SLNs and NLCs. Owing to their small size
in the nanometer range, lipid nanoparticles are naturally stable, and creaming of
aqueous dispersions or nanoparticle sedimentation might not occur. The particles
are kept in suspension by the Brownian motion of the water molecules. Owing to
their surface electrical charge, the particles are stabilized by electrostatics repulsion,
and the physical stability is even further enhanced when sterically stabilizing
polymers, such as Tween 80 or poloxamer, are used as surfactants in the formulation. The high lipid nanoparticle stability as aqueous dispersion has been reported
for more than three years (55). Furthermore, if necessary SLN and NLC can be
incorporated in creams, gels, lotions, or body milks, to increase their physical
stability as reported above.
Apart from aggregation and creaming, another type of instability could be
questioned, that is, the dissolution of lipid nanoparticles in the liquid oil of the
above-mentioned semisolid formulations. In contrast to liposomes, SLNs and NLCs
have the advantage that the stability can be proven quantitatively by differential
scanning calorimetry (DSC). In the DSC measurements, the melting of the nanoparticles and the melting energy in Joule per gram can be determined. Comparing the
melting energy on the day of nanoparticles incorporation into the cream with the
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melting energy obtained after a certain storage time, will allow the calculation of the
total percentage of nanoparticles still present in the formulation. This is not possible
to perform with liposomes. It is easy to prove the existence of liposomes in a product after certain storage time by electron microscopy, however the quantification
of their number is not possible or it is extremely difficult to perform (in this case
using, e.g., quantitative electron microscopy analysis). This is a major obstacle for
introducing a liposomal cosmetic formulation into the market in Japan. However,
this is not a major problem with SLNs and NLCs because the exact amount after a
certain storage time can be easily analyzed quantitatively by DSC (56).
Loading Capacity, Entrapment Efficiency, and
Controlled-Release Properties
Important parameters to evaluate the suitability of a carrier system are the
loading capacity and the entrapment efficiency for the active ingredients. A full
range of model active ingredients has been incorporated into SLNs and NLCs.
An updated list has been recently provided by our group (57). Loading capacity
of 10% to 20% was obtained for tetracaine and etomidate with physical stability
of the particles remaining (58) (please note that the loading capacity is calculated in
percent of the lipid mass). Vitamins A and E and their derivatives have been incorporated into SLNs up to 25% (56). In this study, the incorporation was limited by
the fact that higher percentages did not lead to any more solid particles. However,
if these active ingredients are blended with even higher melting lipids, then the
loading capacity can be further increased. In case of full miscibility of the active
ingredient with the lipid, loading capacities of 50% and more can be achieved.
The loading capacity can be 100% in cases where the active ingredient is lipophilic and solid itself.
The entrapment efficiency is the percentage of active ingredient which is
entrapped inside the lipid particles. For lipophilic active ingredients, the entrapment
efficiencies are typically between 90% and 98%. The lowest values observed were
approximately 80%, for example, tetracaine (59) and clotrimazole (60). For hydrophilic compounds, the loading capacity and the entrapment efficiency are obviously
lower. Values of about 50% have been obtained for the extremely hydrophilic model
compound Iotrolan, an X-ray contrast agent (61). Iotrolan was used as a model compound because of its extremely high hydrophilicity; two parts of Iotrolan dissolve in
only one part of water. Model drugs for peptides and proteins (e.g., lysozyme) have
also been incorporated. It could be proven that lysozyme remained chemically intact
after incorporation into the SLN and was still active (62). With hydrophilic peptides,
loading capacities up to 20% have been achieved for prolonged release from the
particles in vivo in the animal model.
The release of incorporated ingredients from the lipid nanoparticles can be
modulated according to the needs from very fast to very slow. Figure 2 compares
the release of clotrimazole obtained from SLNs and NLCs using Franz diffusion
cells (60).
In NLC formulations, lipid nanoparticles have a liquid core, and clotrimazole
is incorporated in the oil less tightly in comparison to the solid lipid matrix of SLN
formulations. In the latter, drug molecules are incorporated into the crystalline
matrix and their diffusional mobility is decreased (60).
In general, diffusion through the carrier is the main mechanism of controlled
release as described by Fick’s law of diffusion (5). However, drug diffusion
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
223
SLN
NLC
30
Drug released [%]
25
20
15
10
5
0
0
5
10
15
Time [h]
20
25
FIGURE 2 Clotrimazole release profiles from tripalmitin-based solid lipid nanoparticles and nano
structured lipid carrier. Abbreviations: NLC, nanostructured lipid carrier; SLN, solid lipid nanoparticle.
Source: From Ref. 60.
coefficient cannot be considered constant, but it is dependent upon drug concentration. Owing to a large drug loading, the degree of diffusion can be decreased.
There are too many molecules trying to diffuse and they limit their own permeation (hindering effects). Also, the previously described incorporation models for
SLNs (drug-enriched core or shell) underline these findings. Drug loading is very
important with regard to release characteristics (6). Generally, the increase of drug
loading leads to an acceleration of the drug release. However, in particular cases,
increasing the drug loading may slow down the release, which can be explained
by possible drug crystallization inside the nanoparticles.
Figure 3 shows the redistribution effect occurring during the SLN production
by the hot HPH technique (6). According to characteristics, such as drug solubility
and its partitioning coefficient, dispersing the drug-containing lipid melt in a hot
aqueous surfactant solution will lead to distribution of drug into the aqueous phase.
If the aqueous phase contains a higher surfactant concentration, in most cases this
leads to a better solubility of the lipophilic drug in the water phase (e.g., by solubilization) and thus more pronounced distribution of drug to the water phase.
The solubility of the active ingredient in the water phase can be further
increased by choosing higher production temperatures, that is, in this case more
active will partition to the water phase (6). The opposite effect will occur when
the obtained hot nanoemulsion is in the cooling process, that is, the solubility of
the active in the aqueous phase decreases. Lipid starts to precipitate forming a
lipid core with a lower active concentration than in the original drug-containing
lipid melt. Further cooling leads to further reduction of the active solubility in the
water phase and redistribution back to the lipid phase, however due to the formation of the solid lipid core, only the outer shell is accessible for the active. In this
case, most of the active will be released in the form of a burst. The extent of the
burst release can be, therefore, modified by controlling the amount of active in the
outer shell of the obtained particles. This phenomenon has been observed as a
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Souto and Müller
FIGURE 3 Partitioning effect of drug during production of lipid nanoparticles by high-pressure
homogenization technique. Source: From Ref. 6.
function of production temperature and surfactant concentration. Burst release
can also be avoided applying the cold HPH process (63,64).
Chemical Protection of Incorporated Substances
It is known that stabilization of chemically labile actives against degradation (e.g.,
hydrolysis and oxidation) can be achieved using a solid matrix such as the one of
polymeric particles (65). This is also valid for the lipid nanoparticles. The stability of
retinol and coenzyme Q10 could be enhanced when incorporated into lipid nanoparticles made from a mixture of Compritol®888 ATO and Miglyol®812 (66,67).
Ketoconazole could also be protected to a relative extent using the same lipid (68).
Owing to the fluid character of carriers such as liposomes and emulsions, the
active ingredients will partition between the liquid oil and water phases (56).
Permanent exchange of molecules between these two phases will decompose the
active molecules in the water phase, and simultaneously nondegraded compound
will partition from the oil to the water phase (based on partitioning coefficient after
Nernst). The partitioning ratio between nondegraded and degraded active will be
maintained (Fig. 4, upper). In comparison to a solid matrix (SLN or NLC), the active
will be in this case fixed inside the lipid particles. Exchange between the inner and
outer phases will not happen or it will occur very slowly (Fig. 4, lower). These observations emphasize the protection and chemical stabilization effects of incorporated
actives into SLNs and NLCs (69).
Occlusive Effects and Skin Hydration
Owing to the small particle size of SLNs and NLCs, these carriers show adhesive
properties (52,70), properties that have also been observed when using liposomes.
When in contact with the skin, lipid nanoparticles create a thin film with very narrow
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
225
FIGURE 4 Chemical protection of labile actives in emulsion droplets versus lipid nanoparticles.
Free partitioning of nondegraded and degraded active molecules in the emulsion system (upper),
but hindered by the solid state of the lipid nanoparticles (lower). Source: From Ref. 66.
interspaces between the particles. The formation of this film can be sensed when
applying an SLN or NLC formulation onto the skin. Even relatively small amounts of
lipid nanoparticles in a cream (e.g., 5%) can create this film (56).
Figure 5 shows that the relatively high occlusivity is a special feature of lipid
nanoparticles (Fig. 5, lower), when compared with lipid microparticles (Fig. 5,
upper) of identical lipid content. The rather low occlusivity of microparticles is
attributed to the large spaces between the micrometer carriers still allowing the
evaporation of water. The small spaces (capillaries) between the nanoparticles are
hydrodynamically unfavorable and limit water loss (small diameter of pores, higher
resistance to water vapor flow). This film hinders water evaporation, which means
that an occlusive effect leads to increased skin hydration.
This property might be very interesting for skin protection during the day,
when applying placebo SLNs or NLCs. The skin constantly comes into contact with
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FIGURE 5 Mechanism of the occlusion effect depending on the particle size. A solid lipid microparticle dispersion (diameter 1 µm, upper) in comparison to an aqueous solid lipid nanoparticles or
nanostructured lipid carrier dispersion (diameter 230 nm, lower). Source: From Ref. 66.
a wide variety of chemicals which are normally present in the atmosphere. These
include water and dissolved mineral salts present as free electrolytes together with
the atmospheric gases. Skin’s permeability to these different substances varies
widely. Occlusion of the skin leads to increased hydration and subsequently to the
smoothing of wrinkles—an effect utilized in many cosmetic products. On the basis
of this property, SLN- and NLC-containing products are expected to have also antiaging effects, especially when preparing them with skin-caring active ingredients,
such as several vitamins and ceramides. Ceramides can be blended with higher
melting lipids leading to solid nanoparticles, and when applied they will promote
restoration of the damaged protective lipid layer of the skin (4).
Penetration Enhancement of Incorporated Substances
Owing to occlusion properties of lipid nanoparticles and subsequent increased
skin hydration, these carriers can improve the penetration of the incorporated
actives into the skin. The penetration-enhancement effect of lipid nanoparticles has
been tested with tocopherol and tocopherol acetate. Lipid nanoparticles loaded
with those actives have been applied to the skin and the penetration compared to
alcoholic solution of the same compounds by applying a stripping test (cumulative
penetration as a function of the number of strips). Penetration observed with lipid
nanoparticle formulations was twice as high as obtained with the referenced
alcoholic solution. In addition, these data also have shown that active ingredients
incorporated into lipid nanoparticles were obviously released from the carriers
when applied to the skin (66).
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227
SOLID LIPID NANOPARTICLES AND NANOSTRUCTURED LIPID
CARRIERS AS VEHICLES FOR COSMETIC, DERMAL, AND
TRANSDERMAL ACTIVES
The introduction of lipid particles into pharmaceutical technology is reported to be
very old, and recently attempts have been made to introduce lipid nanoparticles
into the market by cosmetic and pharmaceutical companies. The first product is
already on the market and it has been introduced by the company Yamanouchi
(e.g., Nanobase®) in Poland, covered by the Yamanouchi patent (71). This product
contains active-free lipid nanoparticles and it is intended for cosmetic purposes.
Even though cosmetic formulations do not necessarily have any physiological
function, they resemble topical applications for dermatological use in many aspects.
For example, the purpose of a lipstick is essentially decorative, but its formulation
embodies a number of excipients commonly used in pharmaceutical products. Having
considered the ways in which therapeutic applications are formulated, it is only
necessary to examine the modifications of these techniques employed in the cosmetic
field, with special reference to fragrance, feel, freshness, appearance, and skin persistence. Factors such as personal taste and fashion also have a strong influence when
people purchase the product. Thus, in order to develop a successful product, the
formulation must exhibit both functional attributes and esthetic appeal.
Besides lipsticks, facial makeup products can also be developed that have lipid
nanoparticles. They serve to improve the uniformity of coloring, add new color,
disguise blemishes, and impart a smooth matt finish. The face exposes the greatest
single area of the body surface to climatic influence and in view of the major esthetic
significance of its appearance, there is a real need for protection. The relative thinness
of the facial epidermis adds further importance to such protection. Also cleansing
creams, which are oil-miscible preparations useful to remove makeup residues
without the energetic degreasing that may result from using soap and water, can be
formulated with SLNs and NLCs. The aqueous phase of the dispersion has surfactant properties to remove the makeup while the lipid nanoparticles adhere to the
skin protecting it. Other examples are face powders (lipid nanoparticles can be
spray-dried) for a basis coloring or masking to cover minor defects, lipsticks, hair
dressings, nail preparations, depilatories, and shaving aids.
According to market studies, an important factor for a cosmetic to be purchased
is not only that the packaging must have an esthetic appeal, but so must the appearance of the formulation itself. White products are preferred by the consumer, lipid
nanoparticles can be used to weaken, for example, yellowish actives (coenzyme Q10,
vitamin C). This whitening effect is of special interest for actives which are degraded
with degradation products possessing a color, even if they do not influence the
product quality (4). Whitening and lightening properties of lipid nanoparticles are
currently being explored (71–74).
Deodorants and antiperspirants can also have lipid nanoparticles particularly
if formulated with antiseptics, for example, hexachlorophene or trichlorophane,
which appear to increase the permeability of the sweat duct and thus reduce the
amount of the secretion reaching the surface of the skin.
Aging of the skin has been attributed to a change in the circulating sex
hormones, but alterations of the collagen and elastic tissues, which are accelerated
by chronic UV exposition, is probably more significant. It is, therefore, likely that
UV-absorbing cosmetics should have a truly protective role in this context.
Molecular UV blockers have been incorporated into lipid nanoparticles (75–80).
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Souto and Müller
A side effect of these sunscreens is penetration into the skin leading to skin irritation or even allergic reactions (81). Particulate UV blockers frequently used are
titanium dioxide particles which have also been loaded into NLCs (82). There is
also an ongoing controversial discussion whether and to what extent titanium
dioxide particles penetrate into the skin (83,84) where they can potentially interact
with the immune system (85). After incorporation of molecular UV blockers into
lipid nanoparticles, a synergistic effect of the UV scattering caused by the SLN
themselves and the UV absorbance of the molecular sunscreen was observed (86).
This opens the question of reducing the concentration of the molecular sunscreen,
and simultaneously its potential side effects, while also maintaining the UV-protective level. Apart from reduction of side effects, this result is of commercial interest
for expensive UV blockers. In addition, it was found that penetration of molecular
sunscreens into the skin was reduced when comparing SLNs to a traditional o/w
emulsion system of similar composition (78).
The only element common to nearly all cosmetic products is the employment
of a perfume to impart a pleasing fragrance. Lipid nanoparticles can also be used for
prolonged release of perfumes, fragrances, and insect repellents, in comparison to
o/w emulsions and Eau de Toilettes (84). The release can be slowed down by
incorporating the perfume in a solid matrix instead in a liquid lipid particle (oil droplet) (75). This is very interesting if the aim is to create a once a day application with
continuing scent. Fundamentally, perfumes are derived from the essential oils of
botanical origin and most formulae still rely to a considerable extent in the use of
natural oils. Thus, NLCs are suitable to deliver oily fragrances due to the presence of
an inner liquid core in their structure. Although oily fragrances are quite expensive,
using a prolonged release system such as NLC, lower amounts of oil will be needed;
in addition, they can be used in soaps and less expensive products. Prolonged release
was also observed for insect repellents, such as lemon oil (75,88–90).
Dermal and transdermal formulations may also be compounded as freeflowing solids (powders), semisolids, or liquids. The decision to employ a particular
dosage form is mainly dependent on considerations of functional suitability. The
effectiveness of both topical and transdermal systems is related to the extent of
percutaneous absorption of the drug. Thus, there is a tendency to view these systems
as being closely related in terms of functionality. However, when systems such as
SLNs and NLCs are placed on the skin to deliver the incorporated drugs, these can
act: (i) in the local tissues immediately beneath the application site, (ii) in the deep
regions in the vicinity of the application site, and (iii) in the systemic circulation. Of
course, when moving down this progression, the drug-delivery challenge becomes
increasingly difficult. To achieve systemic circulation, drug release over long periods must be provided. The main feature of lipid nanoparticles intended for dermal
and transdermal delivery of drugs is their controlled-release properties (10,91). The
release profile accomplished by lipid nanoparticles is dependent on their structure
(60). As discussed previously, depending on the production method (hot vs. cold
homogenization), the composition of the formulation (i.e., surfactant, lipids), the
solubilizing properties of the matrix for the drug nanoparticles with a different
structure will be obtained (5,63,64). Depending on the matrix structure, the release
profiles will vary from very fast, medium, or extremely prolonged release (60). The
degree of the initial burst release, for example, desired for topical applications, could
be explained by the dissolution properties of the drugs during the production
process being a function of lipid, surfactant, and production temperature. The
understanding of this mechanism allows the controlled production of lipid
nanoparticles with a defined initial dose.
Lipid Nanoparticles for Cosmetic, Dermal, and Transdermal Applications
229
To reach pharmacologically adequate systemic levels in transdermal therapy,
a sufficient amount of drug needs to cross the skin from the application site to the
circulation. Conventional transdermal systems are ointments and adhesive systems
of precisely defined size, for which ideally there would be no local accumulation.
However, in those cases the drug molecules are forced to cross through a small diffusional window defined by the contact area of the transdermal system. Consequently,
irritation and/or sensitizing effects underlying this system might be observed. To
overcome such inconvenience, the use of lipid nanoparticles might be beneficial
particularly due to the fact that they are composed of physiological and wellaccepted lipids. Furthermore, skin lipids (fatty acids, ceramides) can also figure in
the lipid nanoparticle composition. In addition to chemical stabilization of
incorporated drugs (68), drug penetration can be improved by occlusion and further
hydration effects.
CONCLUDING REMARKS
There is no doubt that lipid nanoparticles combine advantages of other carrier systems, such as emulsions, liposomes, and polymeric nanoparticles. Considering their
special properties, SLNs and NLCs will find applications in cosmetic products but
also in pharmaceutical formulations, that is, for dermal and transdermal delivery.
Over the last decade, dermal and transdermal delivery of drugs has become
one of the most promising areas of pharmaceutical research and due to intensive
efforts, thousand of patents have been filed dealing with new concepts and new
technologies. Apart from their special features, major advantages of SLNs and NLCs
are definitely the possibility for large-scale production, the cost-effective production
method, and the relatively low cost of excipients and other components. Owing to
the worldwide patent protection, the product exclusivity can be guaranteed giving
no advantage to competitors. Registered trade names worldwide offer the possibility for sale using a product name optimized considering marketing aspects.
Pharmaceutical and cosmetic preparations are made in a wide range of batch
sizes and the scale of operations strongly influences the final quality of the obtained
product. Wherever those formulations are prepared, raw materials have to be stored,
the ingredients and the final product have to be transferred from one point to
another. They have to be transported and processed in contact with various
constructional materials and finally stored in bulk before transfer to final containers. When dealing with aqueous SLN and NLC dispersions, these need to be physicochemically stable to avoid failure of the scale of production. Large-scale production
is possible for SLNs and NLCs, which means that not only the equipment is
available, but also the obtained batches are physicochemically stable (82,92).
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15
Nano-Carriers of Drugs and Genes
for the Treatment of Restenosis
Einat Cohen-Sela†, Victoria Elazar †, Hila Epstein-Barash†,
and Gershon Golomb
Department of Pharmaceutics, School of Pharmacy, The Hebrew University of
Jerusalem, Jerusalem, Israel
INTRODUCTION
Restenosis
Percutaneous coronary interventions (PCIs) are widely employed for the revascularization of arteries obstructed by an atherosclerotic plaque in patients with
symptomatic coronary artery disease, which usually presents as angina or myocardial infarction (1). The PCI procedures include balloon dilation, endoluminal stenting,
excisional atherectomy, intravascular brachytherapy, and laser ablation. The successful treatment of stenotic coronary arteries by PCI is limited by the occurrences of
restenosis, which continues to be a serious complication (2–6). Restenosis is characterized by three stages of response to the vessel wall injury: (i) acute elastic recoil, (ii)
negative remodeling, and (iii) neointimal proliferation (7). Coronary artery stenting
following balloon angioplasty significantly decreased restenosis rate by solving the
problems of elastic recoil and vessel negative remodeling (5,8–10), however, in-stent
restenosis, due to neointimal proliferation, remains the major limiting factor of PCI.
An inflammatory healing response is triggered by the mechanical damage to the
arterial wall (11–14). At first, platelets are activated and attached around the site of
injury, followed by adhesion of inflammatory cells. Cytokines and growth factors
are secreted by adjacent platelets, macrophages, and/or endothelial cells, including platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), and
interleukin-1β (IL-1β) (15), leading to proliferation and migration of smooth muscle
cells (SMCs) towards the arterial lumen, with subsequent secretion of abundant
extracellular matrix forming neointima and narrowing the artery.
The multifactorial pathogenesis of restenosis allows several options where
pharmacological agents might be applied to prevent the disease process (16,17).
A large number of clinical trials have investigated various systemic drug therapies
in an attempt to reduce restenosis. Pharmacological therapies can be divided
into categories based on mechanisms of action: prevention of thrombus formation,
prevention of vascular recoil and remodeling, and prevention of inflammation and
cell proliferation (16). None of the human trials demonstrated any beneficial effect
on the incidence of restenosis. Systemic pharmacological approaches have failed in
humans due to inability to achieve the required dose at the site of injury without
causing systemic side effects (18). As a response to the failure of the systemic administration, the concept of local drug delivery emerged, offering the advantage of
high local concentration and minimizing systemic side effects due to the relatively
lower systemic concentrations. Until recently, locally delivered drugs have been
unsuccessful in humans due to rapid washout of the drug (19), indicating the need
Equal contribution to this work.
†
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for a controlled administration of the drug for an adequate period of time. The most
successful approach to date is drug-eluting stents, delivering medication directly to
the site of vascular injury from polymeric coated stents (20–24). The first approved
and commercially available drug-eluting stent was the Cypher (Cordis, a Johnson &
Johnson Company; New Brunswick, New Jersey, U.S.A.), containing Rapamycin
(Sirolimus), a naturally occurring macrolide antibiotic and a potent immunosuppressive agent (25,26). Taxus (Boston Scientific Corporation; Natick, Massachusetts,
U.S.A.) was the second approved drug-eluting stent containing Paclitaxel (PXL),
a microtubule-stabilizing agent with potent antiproliferative activity (27,28).
However, despite their clinical success, the long-term efficacy and safety of these
two drug-eluting stents are yet to be confirmed.
Nano-Carriers
Nano-carrier systems, such as liposomes, oil-in-water emulsions, and polymeric
nanoparticles (NP), have been extensively investigated over the last few decades as
a way to modify the biodistribution of drugs (29–36). An encapsulated drug follows
the distribution of the carrier, rather than depending on the drug’s physicochemical
properties and molecular structure. The therapeutic efficacy of pharmacological
agents is dependent on their biodistribution, as well as their elimination route and
kinetics; thus, nano-carriers can be used to improve the therapeutic index of drugs.
Nano-carriers have attractive biological properties, since carrier-incorporated pharmaceuticals are protected from the inactivation effect of external conditions, yet do
not cause undesirable side reactions (37–39). In addition, nano-carriers may be utilized to achieve targeted therapies; targeting to a desirable site may lead to a more
effective therapeutic drug action and prevent side effects (40).
Liposomes
Liposomes are microscopic phospholipid vesicles with a bilayered membrane
structure and have received a lot of attention during the past 30 years as pharmaceutical carriers of great potential (41,42). Liposomes are formed by the self-assembly of phospholipid molecules in an aqueous environment. Liposomes are used as
biocompatible carriers of drugs, peptides, proteins, plasmid DNA, and antisense
oligonucleotides or ribozymes for pharmaceutical, cosmetic, and biochemical
purposes (42). The enormous versatility in particle size (43,44) and in the physical
parameters of the lipids affords an attractive potential for constructing tailor-made
vehicles for a wide range of applications (45–63). Liposomes can be divided into
three subgroups: classical (conventional) liposomes that are short circulating (64);
long circulating liposomes, typically membrane-modified with polyethyleneglycol
(PEG) (37,39,65–68); and immunoliposomes, liposomes with surface-attached
ligands capable of recognizing and binding to cells of interest (69,70).
Nanoparticles
NP are solid colloidal particles in the nano-size range and, when formulated from
biocompatible polymers, they can be used as drug carriers for therapeutic applications. Depending on their design, NP can be divided into two major subgroups:
nanospheres (NS) and nanocapsules (NC). The NS are composed of a matrix
structure, and NC are characterized by a polymeric rate-limiting membrane (71). The
drug may be chemically bound to the particle-forming polymer, adsorbed on the NP
surface, or entrapped in the NP (72,73).
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The NP can be prepared by methods involving either polymerization of
dispersed monomers or dispersion of preformed polymers. The use of preformed
polymers is preferable, since they are well characterized and contain no residual
monomers or polymerization reagents (74). The choice of the polymer depends
upon the therapeutic application of the system, desired drug release, and biocompatibility. The NP may be fabricated using biopolymers (gelatin, albumin,
casein, polysaccharides, lectins, etc.) and synthetic polymers (polycianoacrylates,
polyesters, polyanhydride, etc.). The biopolymers are not capable of providing
protracted release kinetics for small drug molecules, thus limiting possible applications to delivery of biological macromolecules or drugs for which immediate action
is desirable. The polymers most commonly used in this field of research are the
synthetic aliphatic biodegradable polyesters, polylactide (PLA), polyglycolide
(PLG), and their co-polymer (PLGA). The PLGA polymers have been used widely
as biomaterials for medical applications over the last 30 years and are regarded as
biocompatible and nontoxic (33).
Nano-Carriers in the Treatment of Restenosis
In designing a delivery system, it is important to consider the route of administration (local or systemic) and medication dose, as well as mechanism and duration
of the desired drug action. Liposomes and NP as drug carriers can be utilized for
the treatment of restenosis. The nanocarriers can be designed to have varying
degradation and drug release profiles according to specific requirements (73), as
well as tissue and cell localization properties, allowing control of their pharmacological action selectivity. Optimal drug delivery requires preferential localization to
the site of injury, while maintaining a reservoir of drug sufficient for desired activity
duration, in order to avoid uptake into noninjured segments of the artery and
nonspecific cells (75,76). Several modifications of liposomes were proposed to cause
selective cell targeting for the cardiovascular system (77). Modifications promoting
targeting to the injured artery can be divided according to the specific ligands
and receptors on cells involved: (i) GPIIb-IIIa receptor on activated platelets (78),
(ii) tissue factor (TP) on vascular endothelial cells, (iii) upregulated PDGF receptor
for SMCs migration/proliferation, and (iv) E- and/or P-selectine both on endothelial
cells and platelets (79,80). Specific targeting for activated cells causes reduction of
side effects in the nonactivated noninjured site due to lower or no expression on
resting vascular cells (81).
LOCAL DRUG DELIVERY IN RESTENOSIS
It is accepted that the ultimate antirestenotic therapy will utilize the concept of
local drug delivery, considering the regional nature of the restenotic process. This
hypothesis is strongly supported by the virtual failure of systemically administrated
pharmacological agents to control restenosis (17,82), on one hand, and the success
of drug-eluting stents (27,28), on the other. Local delivery of drugs directed against
particular steps in the pathogenesis of restenosis enhances efficacy by increasing
drug concentration to therapeutic levels in the immediate vicinity of vascular injury
with minimal or the absence of systemic side effects (83–90). In addition, therapeutic
agents with short half-life or chemically unstable agents are delivered directly to
their site of target with minimal loss of therapeutic activity prior to their uptake.
Furthermore, the incorporation of the drug into a carrier system allows controlling
its release, thus fitting the drug release to the pathology demands (73).
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Liposomes for local drug delivery offer the advantages of extensive uptake by
a variety of cells and specific localization in the cells (43,44). By controlling liposome
size, one can control the delivery (48,91–94). Small vesicles (<150 nm diameter) bind
to cell surface receptors and are transported by receptor-mediated endocytosis.
After internalization, the vesicle fuses with lysosomes, which induce the breakdown of the lipids and release of their contents (95–105). Large particles (>150 nm
diameter), on the other hand, are taken up principally by phagocytosis, which is
usually limited to phagocytic cells but can be induced in many other cell types with
appropriate ligands (106). In both cases, the liposome could either be degraded in
the low pH environment, or it could fuse directly with the endosomal or lysosomal
membrane. Furthermore, the incorporation of drug within liposomal vector allows
combination of rapid burst release and, at the same time, sustained release by simple
controlling of the liposomal formulation.
In local drug delivery, the ultrasmall size of NP is an important advantage,
since it enables higher arterial uptake because of the apparent ease of access to the
arterial wall (107). Fishbein et al. (108) have shown that rhodamine B containing
polymeric NP was traced in rat carotid arteries 24 hours after intraluminal delivery,
whereas no fluorescent dye was present in the arterial segments treated with an
equal amount of rhodamine B in solution (Fig. 1), indicating that the NP were able
to prevent the rapid elimination of the free drug. Furthermore, NP offer the possibility of intra-arterial delivery via porous catheters that could be included as a final
step in interventional cardiac catheterization (109–111).
FIGURE 1 (See color insert.) Confocal
images of balloon-injured rat carotid arteries after intaluminal delivery of rhodamine
solution and rhodamine-containing nanoparticles. The arteries were harvested 90
minutes (A and D), eight hours (B and E),
and one day (C and F) after delivery.
Abbreviation: NP, nanoparticles. Source:
From Ref. 108.
Nano-Carriers of Drugs and Genes
239
Gene Therapy for the Treatment of Restenosis
The local nature of the restenosis phenomenon makes it an attractive target for gene
therapy, given that the tissue reaction that develops is directly accessible by the
intervention itself (112). The success of gene therapy in the prevention of restenosis
depends on the identification of appropriate molecular targets, a suitable vector
system chosen for efficient vessel wall targeting, and methods for vascular gene
delivery without producing undue damage or distal tissue ischemia (113). The
appropriate genetic modification, performed locally at the time of angioplasty, could
induce a long-term benefit in potency by fundamentally redirecting the healing
response at its roots (114).
Several molecular species have been implicated in the stimulation of vascular
smooth muscle cells (VSMCs) to initiate the intimal hyperplasic process (115,116):
PDGF, basic fibroblast growth factor (bFGF), transforming growth factor beta
(TGF- β), angiotensin II, and activator protein-1 (AP-1) (117–120).
Another strategy is focused on the cell cycle machinery of cells involved in
proliferation and neointima formation (121,122). The VSMC proliferation depends
on the increased expression of certain cell cycle regulatory proteins at critical points
along the cell cycle, including proliferating cell nuclear antigen (PCNA), transcription factor E2F, nuclear factor-κB, cell division cycle 2 (cdc2) kinase, the cyclins,
cyclin-dependent kinase (cdk), and nuclear protooncogenes (c-myc and c-myb),
molecules that appear to be critical in the mitogenic signaling events (123–128). It
has been hypothesized that by blocking the gene expression or function of one or
more of these proteins, one could prevent the progression of VSMC through the cell
cycle and inhibit neointimal growth. Thus, different regulatory molecules can be
specific targets for gene therapy (2,129).
Gene therapy can be defined as a transfer of nucleic acids to the somatic cells
of an individual with a resulting therapeutic effect, such as synthesis of missing or
defective proteins (introduction of intact gene sequences) or expression blockade of
disease-related genes (antisense technology, decoy oligonucleotides, and ribozymes)
(130–133). Somatic gene therapy has emerged as a promising approach for the
prevention of restenosis (134). The aim of antisense-short double-stranded DNA or
RNA oligonucleotides strategy is to inhibit protein synthesis by selectively blocking
the initiation of translation or increasing the mRNA degradation by RNase H
(131,132,135). Natural phosphodiester antisense oligomers are susceptible to rapid
degradation by nucleases. To overcome this problem, a number of chemically
modified nuclease resistant analogs, such as phosphorotioated oligodeoxynucleotides (PT-ODNs) and chimeric oligomers (morpholino and peptide nucleic acids)
have been designed (136–138). Some success in altering intimal hyperplasia in
restenosis has been achieved using antisense oligonucleotides in cell culture and in
animal models (139,140). Treatment with PT-ODNs against c-myc, using a polymerbased delivery system, resulted in only partial inhibition of VSMC growth in vitro
and failed to reduce overall intimal hyperplasia in rat arterial injury model (141).
Introduction of c-myb PT-ODNs, incorporated in the same carrier system, using the
same animal model, was found to be effective in vivo (142). However, in further
studies, the results were not confirmed.
Transfection of cis-element double-stranded oligodeoxynucleotides (decoy)
has been reported as a powerful tool in a new class of antigene strategies for gene
therapy. Synthetic double-stranded DNA with high affinity for a target transcription
factor may be introduced into target cells as a “decoy” cis element to bind the
transcriptional factor. Transfection of such “decoy” oligonucleotides results in
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attenuation of authentic cis–trans interaction, leading to the removal of the transfactors from the endogenous cis-element with subsequent modulation of gene expression (143). Several studies have provided evidence of the in vivo application of this
molecular approach as a therapeutic strategy against cardiovascular disease (133).
A major problem encountered with gene therapy and, particularly, the therapeutic use of oligonucleotides, is the low cellular permeability due to their large
molecular size and high charge density and lysosomal degradation. The uptake of
foreign genetic material (naked DNA, RNA, or oligonucleotide) by mammalian
cells is an inefficient process (118,144). To improve their cellular delivery, several
vector systems were developed (145,146). Viruses are among the most precise and
stable gene vectors. Retroviral, adenoviral, and adeno-associated vectors have the
advantage of exploiting the natural mechanisms of the body’s cell receptors, thus
increasing their efficacy and leading to a stable and long-term expression of the
genetic material in the cells where they have been inserted (147,148). However,
endogenous virus recombination, oncogenic effects, immunogenicity, and unknown
long-term effects lead to serious limitations regarding their use in gene therapy
(145,149–153). Moreover, the recent severe toxicity and side effects of acute T-cell
leukemia in some patients that emerged in a recent clinical trial of retroviral vector in
X-linked severe combined immunodeficiency highlights the importance of nonviral
gene delivery (149,154).
Liposomal Gene Delivery
Nonviral transfection systems, such as cationic liposomes, hemagglutinating virus of
Japan (HVJ)-liposomes, and biodegradable polymeric NP, are generally preferred over
viruses because they are nonimmunogenic, relatively easy to assemble, form stable
complexes with plasmid DNA, provide protection of the plasmid DNA from degradative nucleases, and are amenable to scale-up for industrial production (155,156).
Their drawback is that they are less efficient than the viral vectors (113,157).
Cationic Liposomes
Complexing recombinant DNA with cationic liposomes consisting of cationic and
neutral lipids, also known as cationic lipoplexes, are relatively efficient in delivering
DNA into cells (158), however, the complex can be inactivated in the presence of
serum, and there have been reports of instability upon storage (159). In addition,
cytotoxicity of cationic liposomes remains a concern (160,161). The degree of
successful gene transfer is highly dependent on the cationic lipid type, liposomal
formulation, and cell type (162–165). Cationic vectors bind to negatively charged
DNA, resulting in a condensation reaction and the formation of stable complexes in
the nano range. Fusion with the cell membrane or endocytosis allows incorporation
of DNA into the cell (166). Recent reports suggest that new liposomal formulations,
individually optimized for the targeted tissue, with better protocols and/or continuously administered (poly-)-cationic liposomes may substantially increase transfection efficiencies (167–169). Transfection results have been different in various cell
types and cell lines; thus, each reagent apparently has celltype-specific and additional species-specific characteristics (170). Newer lipids are being developed in
order to enhance vascular gene transfer while minimizing toxicity. New compounds,
such as (±)-N-(3-aminopropyl)-N,N-dimethyl-2,3-bis(dodecyloxy)-1-propanaminium
bromide (GAP-DLRIE), and dioleoylphosphatidylethanolamine (DOPE), have been
shown to increase the vascular delivery of plasmid DNA by 15-fold compared with
previous cationic liposomes (171), screened nonviral formulations transfecting
Nano-Carriers of Drugs and Genes
241
primary VSMCs in vitro with different concentrations and combinations of nonviral
vectors as well as varying DNA/vector ratios and adjuvants. Uptake efficiencies
ranged from 0.1% to 20% of transfected VSMC by changing liposomal parameters.
Use of cationic liposomes (Lipofectamine) for introduction of PT-ODNs, targeting human PCNA mRNA, in vitro influenced the oligonucleotide uptake in
human VSMC and human atherosclerotic plaque, did not enhance antisense effect,
but increased the magnitude of specific VSMC inhibition compared to the naked
sequences (172). Nuclear factor-κB (NF-κB) decoy oligonucleotides transfected by
cationic liposome method (Lipofectamine and TfX50) inhibited tumor necrosis
factor-α (TNF-α)-induced expression of interleukin-6 (IL-6) and intracellular
adhesion molecule-1 (ICAM-1) in mouse endothelial cells (173) and proliferation
of rabbit VSMC (174) in vitro. 1,3-Dioleoyloxy-2-(N-[5]-carbamoyl-spermine)propane (DOCSPER), a recently developed polycationic spermin with no addition
of neutral lipid, has shown the best efficiency with less toxicity. In vivo studies,
comparing 1,3-dioleoyloxy-2-(6-carboxy-spermyl)-propylamide (DOSPER) and
Lipofectamine (DOSPA/DOPE 3:1 W/W) conducted in the pig balloon angioplasty
model using plasmids coding for the antibacterial peptide cecropin A, showed
greater reduction in neointima formation using DOSPER cationic liposomes in comparison to Lipofectamine as nonviral vector (175). The mechanism of enhancement
of gene transfection rate using DOCSPER is not well understood. Nonviral vectors
facilitate gene transfer in different ways, even in the same cell type, due to complex
extracellular and intracellular events involved during the transfection process
(176,177). Membrane binding, internalization, endosomal release, uncoating, and
nuclear translocation constitute basic cellular events during nonviral gene transfer
and may be involved in different degrees depending on the vector used. Combination
of anionic liposomes/DNA and calcium ions has been suggested by Patil et al. (178)
to enable efficient transfection with safer profile. Increased transfection of nonviral
vector can be achieved in vitro as well as in vivo by controlling liposomal formulation and transfer conditions (175,177,179).
Fusigenic Liposome Hemagglutinating Virus of
Japan-Liposomes (Virosom)
The HVJ or Sendai virus is a member of the murine paramyxovirus family, containing a single-stranded RNA virus genome with an envelope (165,180). The HVJenvelope contains two glycoproteins: HN (hemagglutinating neuraminidase) and
F (fusion protein). These HVJ-envelope proteins are involved in cell fusion. The
HVJ virus is an enveloped large particle ranging from 300 to 600 nm in diameter.
The viral particle, which is negatively charged and attached to sialic acid (the HVJ
receptor), fuses with cell membrane, and releases its genome into cytoplasm directly,
rather via the endocytosis (181).
The HVJ-liposome gene transfer technology was developed in late 1980s
(e.g., Ref. 182) and early 1990s (e.g., Refs. 183 and 192) to introduce nucleic acids,
oligodeoxynucleotide (ODN), and protein with high efficiency. The molecules
included in HVJ-liposomes are delivered directly into various types of mammalian
cells by means of the virus cell fusigenic character of HVJ. The HVJ-liposomes are
constructed by a combination of inactivated viral particles and cationic liposomes to
produce a nonviral gene transfer system (165,182,183). The HVJ-liposome nucleic
acids or ODN are more efficient and safer than other nonviral vectors (184,185).
Moreover, ODN delivered by HVJ-liposomes showed rapid accumulation in the
nucleus, which persisted up to two weeks. In the second generation, HVJ-liposomes
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were modified from cationic to anionic. Anionic HVJ-liposomes showed a 5- to
10-fold higher gene expression in several cell types (186). The HVJ-liposomes have
low immunogenic and low pathogenic profile (safety experiments were conducted
in monkeys that demonstrated the safety, feasibility, and therapeutic potential of the
HVJ-liposome vector for humans) (123,187). The HVJ-liposomes’ versatility offered a
wide range of different molecules and high transfection efficiency into a variety of
cells both in vitro and in vivo (117,133,144,186–194).
Intact gene sequences were transferred using HVJ-liposomes for prevention of
experimental restenosis. Plasmids, containing cDNA of endogenous neointimal
growth inhibitors, such as nitric oxide synthase (NOS), prostacyclin synthase (PGIS),
tissue factor pathway inhibitor (TFPI), were delivered to the vessel wall in vivo by
means of HVJ-liposomes (186,188,189). Transfection of NOS inhibited neointima
formation by 70%, 14 days after balloon injury of rat carotid artery (188). Introduction
of cDNA encoding human PGIS into endothelium-denuded rat carotid arteries
resulted in strong PGIS expression in neointimal cells at day 7 and significant inhibition of vascular lesion formation generated at day 14 after balloon injury (neointimal/medial area: 1.2 ± 0.4), in comparison to vehicle control group (1.7 ± 0.5;
P < 0.01) (189). One week after angioplasty and TFPI gene delivery to rabbit iliac
artery, TFPI protein was detected in neointima and media of the vessel wall. At four
weeks, the minimal luminal diameter was significantly greater (P < 0.01) and the
mean percentage of stenosis was significantly lower (37 ± 18 vs. 83 ± 18%, P < 0.01)
in the TFPI-transfected group than in the other three control groups (186).
The HVJ-liposomes were utilized to enhance the efficiency of cellular uptake
and the stability of most commonly used phosphorothioated antisense oligonucleotides, while minimizing their nonspecific toxicity in several in vitro and in vivo
studies (123,190–192). Cyclin B1 and cdc2 kinase antisense oligonucleotides, transferred into injured rat carotid arteries by HVJ-liposomes, were localized primarily in
cell nuclei and caused partial, but specific and significant reduction of intimal
hyperplasia up to eight weeks after transfection, whereas cyclin B1/cdc2 antisense
combination completely abolished neointima formation (123). Transfection of antisense angiotensin-converting enzyme (ACE) oligonucleotides resulted in inhibition
of neointimal growth by 50% without effect on blood pressure, heart rate, and serum
ACE activity (193). Sustained inhibition of neointima growth was produced by a
single administration of HVJ-liposomes containing antisense cdk2 kinase (65%
neointima inhibition) or combinations of cdk2/cdc2 (85%) and cdc2/PCNA (50–
60%) oligonucleotides (191,192). Transfection of decoy ODNs targeting transcription factor E2F using HVJ-liposomes inhibited the induction of c-myc, cdc2, and
PCNA expression, as well as VSMC proliferation in vitro and markedly suppressed
neointima formation compared with control-treated vessels in the model of rat
carotid injury (neointima/media ratio for E2F-transfected segments: 0.291 ± 0.061
vs. untransfected ones: 1.117 ± 0.138) (194). HVJ-liposomal transfer of AP-1 decoy
ODN resulted in significantly decreased VSMC growth (P < 0.05) and migration
(P < 0.01) in cell culture and attenuation of intimal hyperplasia in rat vessel wall in
vivo. Pretreatment of rat carotid artery with AP-1 decoy ODN before balloon injury
was more effective than post-treatment (P < 0.01) in inhibiting neointimal formation
(117,193,194).
Polymeric NP for Gene Delivery
Biodegradable NP also present a potentially advantageous drug delivery system
to transfer genes for restenosis prevention, due to their nonviral nature and their
Nano-Carriers of Drugs and Genes
243
ability to produce local activity of derived gene products (195). Antisense ODNs
utilized to inhibit the function of growth regulatory or cell-cycle genes (c-myb,
c-myc, PCNA, cdc2, and cdk2) of SMCs have been shown to decrease intimal
thickening in experimental restenosis (139,140,142,191,196,197). Nevertheless, fully
phosphorothioated antisense ODNs analogs were tested, which demonstrate high
affinity for various cellular proteins, resulting in nonspecific effects. In addition,
high concentrations of phosphorothioated ODNs inhibit DNA-polymerases and
RNase H, which impairs their effectiveness as antisense agents (131). Cohen et al.
(198) synthesized a novel partially phosphorothioated antisense ODN sequence
for the purpose of minimizing the nonspecific effects of the fully phosphorothioated ODNs. The ODN was designed to downregulate PDGFR-β, which plays a
pivotal role in the enhancement of SMCs migration after balloon angioplasty (199).
The antisense was incorporated into a PLGA-based nanopaticulate system, and the
antirestenotic efficacy was examined in the rat carotid artery model (Fig. 2A). The
purpose of the encapsulation in NP was to enable the intracellular delivery of
the ODN, which exhibit low cellular permeability due to its large molecular weight
and high charge density.
The double emulsion–solvent evaporation method (200) was employed for
the preparation of the PLGA-antisense NP, yielding NP with a spherical shape,
homogeneous size distribution (~300 nm), high encapsulation efficiency (81%),
and sustained release of PDGFR-β antisense for over a month. In order to achieve
high encapsulation, calcium ions were added to the formulation, that minimized
the escape of the negatively charged ODNs to the exterior phase (Fig. 2B), similar
to plasmid DNA encapsulation as was previously reported by the same group
(Fig. 2C) (201).
Partially phosphorothioated antisense sequences were found to be more
specific than the fully phosphorothioated analogs in vitro. The in vivo efficiency
was examined by locally administrating treatment and controls to rat carotid
arteries, immediately after balloon injury. The rats were sacrificed after 14 days.
Scrambled ODN sequence (naked and encapsulated) did not show any effect, which
indicates that the antirestenotic effect was sequence-specific. A significant antirestenotic effect of the naked antisense sequence and the antisense-loaded NP was
observed in comparison to blank NP that had no deleterious effect on the arteries,
while the naked antisense showed a strong trend towards a greater effect than the
nanoparticulate antisense, although this trend did not reach statistical significance.
It was suggested that the dose of the ODN delivered from the NP was markedly
lower in comparison to the naked antisense, since the release of the antisense from
the NP was very slow. It was also proposed that the uptake into the artery was in
favor of the dissolved material rather than the suspended NP. It can be concluded
that while a more specific and effective antisense sequence was designed, the formulation should be further improved to meet the specific requirements of the disease.
The failures of human clinical studies (130,202–204) suggest that the promise of
gene therapy in restenosis was not fulfilled.
Local Drug Delivery of Drugs
Tyrphostins
Tyrphostins are a family of structurally related drugs that possess inhibitory activity
against receptor-bound and cytoplasmic protein tyrosine kinases (PTK) (205,206).
Derivatization of the basic structure of these compounds gives rise to inhibitors
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FIGURE 2 (A) (See color insert.) Inhibition of neointimal formation 14 days after balloon injury to rat
carotid artery and intraluminal instillations with 20 µM (1 nmole) antisense or scrambled partially phosphorothioated oligodeoxynucleotides (PT-ODNs) encapsulated in PLGA nanoparticles (total of
50 µL suspension). Photomicrographs of representative histological sections. Verchoeff’s elastin
stain, magnification ×12.5. (B) (See color insert.) Fluorescence micrographs of blank and PT-ODNloaded PLGA nanoparticles. The PT-ODNs were covalently labeled with FITC prior encapsulation.
Note fluorescence signal in over 90% of the particles. (C) Scanning electron micrographs of blank
(1) and plasmid DNA loaded (2) PLGA nanoparticles, prepared by double emulsion–solvent evaporation method. Gold coating (200 seconds), magnification ×5K, 30 kV, bar = 10 µM. Abbreviations:
A, adventitia; L, lwnen; M, media; N, neointima; NP, nanoparticles; SC-NP, scrambled NP; AS-NP,
antisense NP. Source: From Refs. 198 and 201.
Nano-Carriers of Drugs and Genes
245
exhibiting a high degree of specificity to PTK of a given type, such as PDGFR-β
(207,208). The antirestenotic efficacy of locally delivered inhibitors AG-1295 and
AGL-2043, formulated in PLA-based NP, has recently been evaluated (108,110,209).
These two tyrphostins are low MW, aromatic compounds, whereas AGL-2043 has
some solubility in water due to the presence of a polar moiety in its structure
and AG-1295 is practically insoluble in water. Tyrphostin AG-1295 demonstrated
SMC-specific inhibitory effect on cell growth in vitro and ex vivo (111,210) via inhibition of PDGFR-β-autophosphorilation, which supports the suggested mechanism
of action. The AGL-2043 is a novel tyrphostin characterized by a higher affinity
to the receptor and increased inhibitory potency (207). The tyrphostins were
successfully entrapped in the PLA-based NP by the nanoprecipitation method
(211), which involves a spontaneous gradient-driven diffusion of water miscible
organic solvents into continuous aqueous phase (110,212,213). The release of both
tyrphostins was studied using the external sink method. The AGL-2043 exhibited a
more rapid release in comparison to AG-1295, apparently due to partial adsorption
of the former on the NP (110,214). Rapid drug release associated with substantial
drug distribution onto NP surface has previously been reported for other semi-polar
compounds formulated in NP by nanoprecipitation (215).
Selected NP formulations of AG-1295 and AGL-2043 were evaluated in animal
models of restenosis (108,110). Neointimal formation was significantly reduced
by locally delivered 90 nm NP of both AG-1295 and AGL-2043 in comparison to
control animals, but not by 160 nm NP at the same drug doses (108,110). The insignificant effect of the 160 nm particles might be associated with the less efficient
ingress in the arterial tissue (108) and possibly with reduced uptake by SMCs (216).
The small NP were characterized by a rapid elimination in the first 90 minutes
followed by a slower late elimination (108). The early faster elimination of the
smaller particles was probably caused by the easier migration to the adventitia,
facilitating their elimination through the vasa vasorum. Their reduced late washout
was probably due to their better penetration into deep arterial structures and creation of a drug depot that is relatively inaccessible to leaching by blood flow. The
inferior antirestenotic effect of the large particles is in good accord with their low
tissue drug level at advanced time points. The NP of AGL-2043 exhibited higher
antirestenotic efficacy compared to AG-1295 loaded NP of comparable size and
molar drug loading, which may be attributed to the lower potency of the second
molecule. The pig model is considered to be more relevant for the pathophysiology
of human restenosis (217). Further examination of the antirestenotic effect of NP
formulations of the two tyrphostins was performed in the pig model, and a significant antirestenotic effect was observed in comparison with unloaded NP of the
same size (110).
Dexamethasone
Dexamethasone (DXM) is a glucocorticosteroid drug that was found to significantly reduce neointimal proliferation in the rat carotid artery model, following
periadventitial delivery (218). There are several possible mechanisms for the
antirestenotic effect of glucocorticosteroids, including inhibition of PDGF production (219), IL-1β transcription (220) (a cytokine that stimulates SMCs proliferation),
and reduction of the inflammatory response (221). In light of these results, Guzman
et al. have decided to examine the antirestenotic efficacy of locally delivered DXM
PLGA-based NP (109).
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The result of the in vivo efficacy study in the rat carotid artery model demonstrated a significant decrease in the intima/media ratio with DXM NP, in comparison to blank NP and DXM NP injected intraperitoneally (IP) (109,222). The DXM
levels in the arterial wall were detectable up to two weeks following intra-arterial
administration of DXM NP, while no detectable DXM levels were observed beyond
three hours after IP administration, indicating the successful localized delivery of
the NP (109,222). The antirestenotic effect of DXM delivered in NP was compared to
silicone perivasular delivery. The efficacy of the silicone implant was found to be
higher by more than two-fold, in spite of the higher drug dose delivered into the
artery by the NP (109,218,222). It was hypothesized that the inferior effect of the NP
was due to their less efficient arterial uptake and shorter residence time in the artery,
indicating the need for formulation improvement.
2-Aminochromone U-86893 (U86)
The U-86 is an antiproliferative cytostatic drug that was found to have an antichemotactic and antimutagenic activity on SMCs in vitro (223). A significant antirestenotic
effect in balloon-injured rat carotid arteries was achieved by systemically injected U86. However, due to rapid plasma clearance and poor oral availability, the IV doses
required to achieve the therapeutic effect were very high and prolonged (223).
Humphrey et al. (224) examined the inhibitory effect of PLGA-based U-86 NP on
neointimal proliferation in porcine coronary arteries subjected to balloon injury. The
emulsification solvent evaporation method was employed to formulate NP of
110 ± 40 nm. The release was biphasic and relatively prolonged; over 40% of the
loaded drug was released over the first 24 hours, followed by a slow phase with
exponentially decreasing release rate over the next two weeks. It was found that 9 mg
of U-86 encapsulated within 60 mg of PLGA NP, locally administered by direct
intramural infusion via the Dispatch® catheter, significantly reduces neointimal
hyperplasia development in severely balloon-injured porcine coronary arteries. The
antiproliferative effect of locally delivered nanoparticulate U-86 was limited primarily
to those arteries that had extensive IEL/medial ruptures. No effect was observed in
moderately injured vessels, which was probably due to insufficient retention of active
drug in these vascular sites.
The Effect of Formulation Characteristics and
Delivery Conditions on Arterial Uptake
In order to optimize the intra-arterial localization of therapeutic agents using the
NP nano-carrier system, several groups examined the performance of NP as a
function of formulation characteristics as well as delivery conditions.
Particle size was found to be an important determinant of the arterial uptake.
Song et al. (225) reported on a three-fold lower arterial uptake of large PLGA-based
NP (266 nm) in comparison to small-sized NP (100 nm) in the ex vivo dog femoral
model. Westedt (226) demonstrated a size-dependent NP penetration into an intact
rabbit vessel wall: while smaller NP (100–200 nm) were deposited in the inner regions
of the vessel, larger NP (514 nm) accumulated primarily at the luminal surface of the
aorta. It was demonstrated by the authors that the residence properties of tyrphostinsloaded NP in rat carotid arteries were influenced by their size. The smaller NP
(90 nm) showed better ingress into the arterial tissue than the large NP (160 nm), as
indicated by a 3.4-fold higher initial drug concentration in the tissue, more than
eight-fold higher amount of NP remaining associated with the artery following the
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rapid NP washout phase and higher local drug levels maintained over 14 days
following administration (108). Furthermore, the 90 nm NP, but not the 160 nm NP,
significantly reduced neointimal hyperplasia in the rat carotid model, establishing
for the first time a correlation between NP size and antirestenotic efficacy (108,110).
Drug loading in the NP carrier was also found to be a significant factor in the
arterial uptake of NP. It was demonstrated that higher U-86 loading reduced the
arterial U-86 levels in both ex vivo and in vivo studies (76), which can be attributed
to faster release kinetics of the drug from the formulation containing the higher
drug loading. An increase in the NP concentration in the infusion solution increased
the arterial U-86 levels up to a certain NP concentration beyond which the arterial
uptake of the NP was limited, probably due to charge neutralization (76). The same
group also investigated the influence of surface modifications of PLGA-based NP
containing U-86 on their arterial uptake (76,227). In general, cationic surface modifying agents demonstrated increased arterial drug levels. The NP surface, modified
with a cationic compound didodecyldimethylammonium bromide (DMAB), demonstrated 7- to 10-fold greater U-86 arterial levels in comparison to unmodified NP
in different ex vivo and in vivo studies. The mechanism responsible for the enhanced
arterial uptake was assumed to involve an alteration in the NP surface charge (76).
The aforementioned data confirm that NP properties, such as size and surface
chemistry, play a critical role in their arterial uptake and deposition; however, their
successful application also depends on the delivery conditions, as well as on possible encountered anatomic barriers. In particular, for NP delivered to the artery by
infusion through a balloon catheter delivery system, it was reported that the infusion duration and pressure influence the NP arterial localization. Repeated short
infusions of NP suspension were found to be two-fold more effective in terms of
drug arterial levels than a single prolonged infusion, which was not significantly
different in efficacy from a short single infusion (227). In a different study, an
enhancement of particle deposition at the delivery site was observed as a consequence of an infusion pressure increased from 2 to 4 atm. Recently, it was also
reported that the localization of the NP in a coronary artery of a pig was influenced
by the catheter type (228). The efficiency of arterial localization of NP was evaluated
in the porcine coronary model using two catheters differing in their work principle.
In the Dispach® catheter, the drug is infused into small closed chambers formed
when the balloon is inflated and can diffuse into the arterial wall under pressure.
The infusion using this catheter can be prolonged, since the blood continues to flow.
The second catheter was the Infiltrator® in which the micro-injector-ports on the
surface of the balloon penetrate the arterial wall upon inflation of the balloon and
allow intramural delivery. This catheter cannot be used for prolonged infusion since
it blocks the blood flow completely. The Dispach® was found to be more efficient
than the Infiltrator®; it was suggested that that the micro-injector-ports may have
failed to penetrate the intimal layer. Possible anatomic barriers to any particle transportation through the wall tissue are the intact internal and external elastic lamina.
Moreover, the presence of an atherosclerotic plaque on the inner luminal surface
strongly influences the particle penetration into the arterial wall (226). This observation points to at an additional aspect of local NP delivery strategies in restenosis
that should be taken under consideration.
In summary, while NP present a promising drug carrier system for arterial
local drug delivery, NP formulation properties, as well as their delivery protocol,
should be carefully optimized to achieve the antirestenotic effect. The lack of clinical
trials suggests that the promise from the studies was not fulfilled.
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SYSTEMIC DRUG DELIVERY
The failure of the numerous systemic treatment attempts to prevent arterial restenosis in humans raises doubts regarding the feasibility of the systemic approach.
However, an efficient systemic treatment could potentially address problems
inherent in the local treatment strategy, regardless of its initial success rate. As
opposed to local therapy, a systemic drug delivery system may allow the treatment
of multiple lesions with a single injection and provide an option for convenient
administration of repeat doses for the treatment optimization, as well as flexibility
in choosing the type and number of stents to be deployed for varying lesion lengths,
artery sizes, and anatomic locations.
Numerous studies established the potential of systemic treatment with liposomal and nanoparticulate formulations (95–97). By controlling their formulation
characteristics and composition, these nano-carriers can present a highly efficient
drug device with specific release profile and definite target sites (38,39,64,91,92,94,
229–232). In addition, due to their small size and biocompatible components, liposomes and NP are safe for parenteral administration. Furthermore, liposomes and
NP can be surface-modified with tissue-specific ligand to achieve arterial localization, following systemic administration; however, to date, no significant developments have been reported for targeted delivery systems in vivo.
Systemic nano-carriers therapy for restenosis includes systemic administration intended for arterial localization, attained either passively (by controlling the
formulation properties) or actively (by attaching the ligand on the nano-carrier
surface), and systemic administration aimed at a systemic target.
Systemic Delivery for Arterial Localization
Doxorubicin
Doxorubicin is a drug clinically used for the treatment of cancer; it is often administered as a liposomal formulation pegylated or in ultrasmall nanosize in order to
reduce its uptake by the mononuclear phagocyte system (mps) and reduce the drug
toxicity (233). Doxorubicin damages DNA by intercalation of the anthracycline
portion, metal ion chelation, or by generation of free radicals. It has also been shown
to inhibit DNA topoisomerase II, which is critical to DNA function. The cytotoxic
activity of the drug is not cell-cycle specific. NK911 are core-shell NP formed by
PEG-based block copolymer encapsulating doxorubicin (234). When NK911 components are dissolved in the aqueous phase, they form stable NP (polymeric micelles)
with an average diameter of 40 nm. NK911 NP were found to selectively penetrate
through tumor vessel walls, having enhanced permeability (234). Enhanced vascular permeability that was first recognized in tumor tissues (235) was later on also
observed in inflammatory and infected tissues (236–238). Uwatoku et al. (239)
hypothesized that balloon-injured coronary arteries also have enhanced permeability and suggested that they could be a good target for NK911. They confirmed that
balloon injury causes a marked and sustained increase in the vascular permeability
for at least seven days using Evans Blue staining. The NK911 accumulated in the
vascular lesion, where permeability was increased; the tissue concentrations of doxorubicin were up to four-fold higher three hours following IV injection of NK911 in
comparison to free doxorubicin. Low drug levels were observed in the uninjured
contralateral arteries used as a control. The accumulation was attributed to two
factors: the size of the NP, which was adequate for enhanced accumulation, and
the NP negative surface charge, which created an attraction of the formulation to
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the positively charged luminal surface of the injured blood vessel. The IV injected
NK911 immediately and three and six days after the balloon injury, but not doxorubicin alone, significantly inhibited restenosis in the rat carotid artery four weeks
after the injury in both single- and double-injury models. It was demonstrated by
immunostaining that the antirestenotic efficacy of doxorubicin NP was achieved by
the inhibition of SMCs proliferation, rather than due to enhancement of apoptosis
or inhibition of inflammatory cell recruitment. The RNA protection assay demonstrated that the expression of several cytokines was inhibited by NK911, but not that
of apoptosis-related molecules. However, even in light of the encouraging results
using the NK911 particles, the clinical potential of this formulation is limited due to
the high toxicity of doxorubicin.
Paclitaxel
The PXL is a potent antiproliferative drug clinically used to treat cancer. The PXL
promotes polymerization of the α- and β-units of tubulin and causes abnormal
stabilization of microtubules required for G2 transition into M-phase (240,241). In
low doses, the structural changes in the cytoskeleton result in a nearly complete
inhibition of growth and proliferation and migration of SMCs for a long period. The
most common delivery of PXL for the prevention of restenosis is local via drug-eluting stents (28). Kolodgie et al. (242) chose to examine the feasibility of systemic
administration of nanopariculate PXL for the treatment of restenosis. The rationale
was that systemic delivery of PXL would provide a more uniform stented arterial
segment exposure to the drug, treatment of multiple lesions, and readily adjusted
target dose. Nevertheless, it should be noted that the NP were administrated locally
to the iliac bifurcation. A systemic IV infusion was only employed for the repeat
dose at 28 days. The NP used was a novel albumin-stabilized 130 nm NP formulation of PXL (nPXL) (243). Using albumin as the stabilizer solved the hypersensitivity
reactions (despite prophylaxis) and long infusion rates of 3 to 24 hours associated
with the former nPXL containing Cremophor EL as the surfactant (244). Doses of
nPXL were administrated as a 10-minute intra-arterial infusion through a balloon
catheter at the iliac bifurcation to NZ white rabbits, immediately after bilateral iliac
artery stent implantation. The nPXL reduced neointimal growth at 28 days produced in a dose–response manner, accompanied by incomplete healing. The efficacy
was not sustained after 90 days following a single dose of nPXL of the highest dose
tested (5 mg/kg). However, a second 3.5 mg/kg dose administrated intravenously
28 days after stenting led to persistent reduction of neointimal formation at 90 days
with almost complete healing. It was suggested that the novel PXL formulation may
offer an antirestenotic therapy with reduced toxicity that may overcome some of the
limitations of the drug-coated stents. However, there are no proven advantages
over the clinically used, highly successful, Taxus, drug-eluting stent.
Systemic Delivery for a Systemic Target
Better insights and understanding of the involvement of innate immunity as a
vital factor in the progression of restenosis allowed the evolution of a new systemic
therapy, directed against the immune cells. Innate immunity triggers the healing
response that leads to neointimal formation (245–247). Monocytes/macrophages
play a key role in the inflammation cascade, resulting in restenosis of blood vessels
(248,249). Emerging experimental and clinical data indicate that the inflammatory
cell number within the vessel wall is a powerful predictor of the extent of cellular
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proliferation and intimal thickening (250,251). Monocytes/macrophages comprise
up to 60% of neointimal cells after stent-induced arterial injury in rabbit, porcine,
and nonhuman primate models and in human autopsy specimens (251). It was
recently shown by Fukuda et al. (252) that the circulating monocytes count in human
patients increased and reached its peak two days after stent implantation, and that
the maximum monocytes count after stent implantation showed significant positive
correlation with in-stent neointimal volume at six months follow-up in patients
after stent implantations.
Macrophages belong to the MPS. They originate from pluripotent stem cells of
bone marrow that are precursors for all hematopoietic cells, that is, lymphocytes,
erythrocytes, osteoclasts, neutrophils, and mononuclear phagocytes (253). After
certain differentiation steps in the bone marrow, the committed stem cells give rise
to monocytes, which move to the circulation, migrate to distinct tissue compartments, and differentiate into macrophages. Macrophages are an essential part of the
immune system; in normal steady state in the body, they form a constant population
with a balance between renewal and cell death. In inflammation, the distribution
and development of macrophages changes significantly (254–256). Following a
stimulus such as inflammation, the amount of circulating monocytes and their
migration to the site of inflammation increases significantly. It was hypothesized
by the authors that systemic inactivation of monocytes and macrophages may lead
to attenuation of neointimal formation. In order to target therapeutic agents to
monocytes and macrophages, the authors used liposomes. Liposomes are readily
taken up by cells of the MPS (formerly known as RES), macrophages, in particular,
and to some extent, neutrophils, by the process of phagocytosis. Cell-specific
delivery system for monocytes and macrophages depletion could be beneficial
for the attenuation of the restenotic processes, while causing minimal toxicity to
non-phagocytic cells (245).
Monocyte depletion was previously achieved by administration of other
particulates, silica, asbestors (257), and carrageenan (258). These methods resulted in
partial depletion as well as unwanted effects on non-phagocytic cells. Rooijen (259)
was the first to use a liposomal formulation containing a bisphosphonate (BP) in
order to deplete systemic monocytes. Liposomes for the treatment of restenosis via
monocytes/macrophages depletion should possess optimal size distribution, since
liposomes size control circulation time. The general trend for liposomes of similar
composition is that increasing size translates into more rapid uptake by MPs (91,94),
in particular by circulation monocytes.
Liposomes surface charge is also an important factor; charged liposomes
are cleared more rapidly than neutral, and negatively charged liposomes are eliminated faster than positively charged liposomes (63,229,260). The effect of lipids
composition, choosing lipids with specific transition phase, is an additional important factor that is intricately interrelated with the effect of charge. Circulation time
is dependent on membrane packing and permeability considerations, as the inclusion of high-phase transition lipids will increase circulation lifetime. The presence
of cholesterol in liposomes probably has one of the most important roles in the
maintenance of membrane bilayer stability and, consequently, long circulation time
in vivo (261). In the absence of cholesterol, liposomes are destabilized by high-density lipoprotein (HDL) particles and, upon release, their components can be readily
eliminated from the circulation (262,263). Liposomes with cholesterol display negative
correlation between clearance and stability in plasma (262). Designing an optimal
liposomal formulation for the treatment of restenosis requires consideration of all
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these factors together. Negatively charged liposomes at the range of 200 nm are
preferable due to the compromise between high phagocytic capacity by monocytes
and lower systemic toxicity. In order to achieve optimal release profile and plasma
stability, incorporation of cholesterol is essential. Conventional liposomes, negatively charged with an average size of 20 nm containing cholesterol, were utilized to
achieve better and improved phagocytosis capacity and, at the same time, suitable
release pattern and plasma stability. Thus, the MPS cells, mainly monocytes, are the
target site for the systemic treatment of restenosis.
Bisphosphonates
The BPs, bone-seeking agents, are a family of drugs that inhibit bone resorption via
osteoclast inactivation and are used clinically in several calcium-related disorders,
such as tumor osteolysis and osteoporosis. The BPs are hydrophilic-charged molecules that do not penetrate cells (264,265). After administration to patients or animals,
they accumulate mainly in bone tissue, and are cleared rapidly from the circulation
into the urine (264–266). The in vivo effect on bone by BPs is mediated by the phagocytosis of the bone-adsorbed BP by osteoclasts. Osteoclasts and macrophages share a
common hematopoietic progenitor cell in the bone marrow. Liposomes, a particulate
dosage form, can be used to enhance the intracellular delivery of BPs into phagocytic
cells, such as, monocytes/macrophages, in cell culture and in animals (259,267).
Negatively charged liposomes were prepared by thin lipid film hydration.
The liposomes were composed of distearoylphosphatidylcholine (DSPC), phosphatidylglycerol (DSPG), and cholesterol (CHOL) at a ratio of 3:2:1, respectively
(Fig. 3A). Anionic liposomes are nontoxic, and, after phagocytosis by monocytes/
macrophages, the lipid bilayers of the liposomes are disrupted under the influence
of the lysosomal phospholipases in the macrophage. The drug, which is dissolved in
the aqueous compartments, is released into the cell. Furthermore, a free BP, released
by leakage from liposomes or released from dead macrophages, will not enter cells
in amounts that are able to disturb their metabolism. This approach, named the
liposome-mediated macrophage “suicide” technique, was intensively used to eliminate macrophages from different compartments of the body in animals to study the
role of macrophages in pathological and immunological conditions (268).
In macrophages, cell line (RAW264) and human monocytes (primary culture),
highly endocytotic cells, it was found that encapsulation of BPs in liposomes
enhances their inhibitory activity 20- to 1000-fold compared with free drug (Fig. 3B)
(269–272). The SMCs and endothelial cells (ECs) are insensitive to the liposomal
drug delivery (245). Enhanced phagocytosis into cells may be achieved by negative
and positive charge of the liposomes (45,59,63,229,267). Positively charged lipids
are not approved by the FDA for clinical use. The surface charge density of the
liposomal bisphosphonates (LBPs) has been optimized for minimal leakage and
effective intracellular delivery of encapsulated drugs (260,271).
Activated macrophages secrete cytokines and growth factors, leading to
enhanced inflammation, and induce SMC proliferation and migration towards the
arterial lumen, forming neointima and narrowing the artery (273–280). The BPs
have numerous biochemical effects on cellular metabolism, ranging from the inhibition of general cell metabolism to modulation of cytokine secretion (281). Nonamino
BPs as clodronate and etidronate have anti-inflammatory properties (282–284). They
reduce cytokine secretion upon stimuli. In contrast to the non-amino-BPs, clodronate and etidronate, the amino-containing BPs, alendronate and pamidronate at a
dose that does not kill the cells, showed proinflammatory properties on macrophage
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FIGURE 3 (A) Cryo-TEM microscopy of DSPC:DSPG:CHOL liposomes obtained by thin film
hydration method, and extruded through 0.2 µm polycarbonate membranes. (B) Effect of liposomal
formulations: bisphosphonates (BPs) (alendronate and clodronate), empty and free drugs on RAW
264 cell survival. Curves represent percent of cell inhibition with different BP concentrations. Cell
count in buffer only was determined to be 100% (n = 3). (C) (See color insert.) Low- and high-power
photomicrographs of hypercholesterolemic rabbit iliac artery stents at 28 days (Verhoeff staining) of
control (I and II) and of animals treated with LA 3mg/kg (III and IV) and 6mg/kg (V and VI). (D) Bar
graph showing reduction in intimal area and increased luminal area in treated animals. (E) Bar
graph showing reduction in stenosis (%) in LA-treated animals (n = 16 arteries/group, *P < 0.05).
Abbreviation: LA, liposomal alendronate. Source : From Refs. 269 and 292.
Nano-Carriers of Drugs and Genes
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functions by inducing the secretion of cytokines from macrophages (285). It has
been shown that BPs also diminish secretion of reactive oxygen species by human
neutrophils, polymorphonuclear leukocytes, and macrophages in vitro (286–288).
Encapsulation of BPs in liposomal formulations enhanced their potency.
Liposomal clodronate (LC) was more than ten times more potent an inhibitor of
cytokine secretion from RAW 264 than the free drug (289).
The PDGF-BB is a strong chemoattractant for vascular SMCs involved in
neointima formation secondary to vascular injury (290,291). In vivo studies conducted in the authors’ laboratory revealed that PDGFβR activation (i.e., tyrosine
phosphorylation) is markedly increased to 135% of baseline levels in balloon-injured
arteries of untreated rats, whereas it was barely detectable in LC-treated rats (i.e.,
below baseline activity) (269). Macrophages are a rich source of growth factors and
cytokines, which facilitates SMCs migration to the injured vessel (248). Suppression
of these mediators activation in treated animals (rat and rabbit) corresponds to a
substantial reduction of PDGF-BB protein levels in the lesion, which can explain the
reduced SMC migration and neointimal formation in treated animals. The authors’
data are in conjunction with reduction in arterial and blood cytokines, IL-1β, TNFα,
NFκB and MMP-2 activity, following injury (245). The systemic inactivation results
in reduced expression of local inflammatory mediators, leading to reduced activation and proliferation of SMC and decreased neointimal formation.
It was reported by the authors that LBPs deplete blood monocytes (FACS),
tissue macrophages, and total WBC (Coulter) count in the rabbit balloon angioplasty
model (with or without stenting) (245,269). Systemic administration of LBP reduced
circulating monocytes as well as tissue infiltration and accumulation of macrophages. The WBC count increased slightly 48 hours after surgery, with no significant difference between controls and the LBP groups (269). Monocyte number at 24
and 48 hours after balloon injury and stenting was significantly lower in LBPstreated animals. Blood monocytes depletion and elevating WBC was partial and
transient, lasting six days after IV injection of LBPs. Site-specific reduction of macrophages numbers was observed at the injured arterial site in LBPs-treated rabbits,
three and six days after injury (245). Furthermore, the number of tissue macrophages in liver and spleen were reduced by LBPs at 6 days after treatment.
Liposomes loaded with the fluorescent marker, Rhodamine, with or without a
BP were utilized to support the notion that LBPs exert their effects systemically
(245,269). Depletion of systemic monocytes that carry the liposomes to the injured
site should result in reduced arterial uptake of liposomes. Indeed, fluorescent-labeled
LBPs significantly reduced the fluorescent signal in the injured arterial wall, whereas
massive fluorescent was detected without the BP (245).
The LBPs were injected at different dosage regimens to rats and hypercholesterolemic rabbits to examine the therapeutic effect.
The rat carotid artery injured by a balloon catheter has been widely used as a
model of angioplasty. The rat model is a “proliferation” model without foam cells.
This form of injury causes immediate coagulation and thrombosis cascade in which
platelets adhere, spread, and degranulate on the denuded surface of the artery, and,
approximately 24 hours later, SMC begin to proliferate. The LBPs, clodronate and
alendronate, were injected to male sabra rats, 15 mg/kg and 3 mg/kg, respectively.
Marked neointimal formation and decreased luminal area were observed in control
animals. The LC and liposomal alendronate (LA) suppressed intimal growth when
administered IV on day −1 and +6. The N/M ratios were reduced by 60% and 69%
for LC and LA, respectively (245,269,292).
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The hypercholesterolemic rabbit model simulates the human model, since
foam cells are involved. Liposomal clodronate and alendronate were injected one
day before balloon angioplasty and six days after (15 mg/kg and 1.5 mg/kg, respectively), resulting in significant attenuation of intimal hyperplasia and luminal
stenosis 28 days after surgery (245,269). Stenosis was reduced from 75 ± 8% in the
control: empty liposomes, saline and free BP, to 41 ± 8% with LC treatment and 68 ±
5% with LA (3 mg/kg) treatment. Stenting procedure in the Iliac artery resulted in
abundant concentric neointimal formation composed of SMC and foam cells, with
both intraluminal and outward neointimal growth. Luminal stenosis 28 days after
stenting was 58 ± 11%. The LA (3 mg/kg) significantly reduced neointimal formation compared with control groups. No significant difference was observed between
animals treated with 3 or 6 mg/kg LA or 15 mg/kg LC (Fig. 3C).
Different dosage regiments were examined; multiple doses of LC (15 mg/kg)
or LA (1.5 mg/kg) one day before balloon angioplasty and six days after were found
to have the same effect as one dose at day −1. Changing the time of injection from −1
to one injection six days after surgery had no effect. Treatment of LA, a single dose,
at the time of injury required dose adjustment, elevation of dose from 1.5 to 3 mg/
kg (269). A drug potency effect relationship of reducing restenosis was established
by the authors, alendronate > pamidronate > ISA-13-1 > clodronate (266,293).
Nonamino BPs as clodronate are several orders of magnitude less potent than the
amino BP, alendronate, in inhibiting osteoclasts and consequently bone-related
disorders, such as tumor osteolysis and osteoporosis (265). Consistent effect was
established with monocytes/macrophage inhibition (245).
Having established the antirestenotic effect of BPs encapsulated in liposomes,
the authors further investigated the use of polymeric NP for the same purpose, in
order to examine whether a polymeric nanoparticulate formulation containing BPs
could exert a comparable effect on LBPs. Similar to liposomes, a successful targeting
to monocytes/macrophages requires knowledge of the particle characteristics that
promote their phagocytosis. The physicochemical properties of NP, such as size and
surface chemistry, also play an important role in opsonin adsorption and the elimination of NP from the circulation (294,295). Phagocytosis may be increased with
increasing surface charge of the particle, particularly when the charge is negative
(296). In addition, larger-sized particles are more inclined to be phagocytosed than
smaller ones (297). Encapsulation of hydrophilic agents, such as BPs, in a nano-carrier based on a lipophilic polymer is challenging due to the agents’ affinity for the
external aqueous phase (40,225,298,299). The encapsulation efficiency of hydrophilic
BPs in NP may be optimized by the in situ formation of their salts with low solubility in water, such as poorly water soluble complex with calcium. The double emulsion–solvent evaporation method (200) was used for the NP preparation, and their
antirestenotic effect was examined in animal models. Unlike other NP formulations
examined for restenosis therapy releasing the therapeutic agent in a slow release
manner, the release rate of the BPs was designed to be rapid inside the phagocytic
cell, protecting the drug in circulation for a sufficient period till phagocytosis. The
NP of an aminobisphosphonate (ISA) were found to release 88% of the drug within
20 minutes and the release was completed within two hours. The ISA NP with
encapsulation yield of 69% and average size of 392 nm, both subcutaneously and
intravenously injected one day before balloon angioplasty (15 mg/kg), resulted in
significant attenuation of intimal hyperplasia and stenosis at 14 days in the rat
carotid injury model compared with control animals treated with free ISA, carrier
solution or blank NP. Alendronate NP with encapsulation yield of 55.1% and
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average size of 223 nm, subcutaneously administered (1.5 mg/kg) one day before
and six days after balloon angioplasty, resulted in a reduction of intimal hyperplasia and stenosis at 28 days in the rabbit model.
The effect of BPs encapsulated in nano-carriers on SMCs proliferation is indirectly mediated by the inhibitory effect on monocytes/macrophages. By systemic
administration of BPs encapsulated in nano-carriers, systemic monocytes and tissue
macrophages depletion is achieved, reducing the number of monocytes/macrophages available, and thereby reducing their accumulation in the arterial wall and
the subsequent contribution to SMCs migration and proliferation. This approach,
biologic targeting, is fundamentally different than any other treatment modality in
restenosis. While other nano-carrier formulations, even if administrated systemically, are aimed to effect locally inhibiting local events related with the restenotic
course of action, the encapsulated BPs present a systemic therapy to a systemic
process, regardless of the procedure and the devices used. If effective in a clinical
setting, it may be an easily administered, cost-effective modality that allows flexibility in choosing the type and number of stents to be deployed and may serve as an
adjunct therapy in high-risk patients.
SUMMARY
To date, local drug delivery is considered to be the most favorable treatment for
restenosis, with drug-eluting stents being the leading approach in clinical practice;
however, their long-term efficacy and toxicity should be further examined. In addition, the antirestenotic effectiveness in high-risk groups, as well as in cases of small
vessels, long lesions, and ostial or bifurcation lesions is yet to be established.
Another strategy for the prevention of restenosis by the local route investigated over the last few decades is nano-carriers, including liposomes and NP. These
carriers offer a potentially improved delivery system, since they incorporate the
advantages of local drug delivery and, in addition, enable targeting to specific cells.
Liposomes and polymeric NP were found to be highly efficient in the local delivery
of both pharmaceutical agents and gene products. In the field of gene therapy for
restenosis, numerous liposomal formulations were examined. The formulations
were modified by either surface charge or nonimmunogenic viral vectors for the
enhancement of their penetration through the arterial wall as well as their incorporation into specific cells and specific intracellular localization. However, thus far, all
the approaches in gene therapy for the prevention of restenosis failed to produce
satisfactory results in clinical trails. The use of NP for gene therapy in restenosis is
in its early stages. The systemic approach for restenosis prevention (with a free drug
in solution) failed to produce satisfactory results in clinical trials. This is because
therapeutic drug levels in the injured artery were not achieved following systemic
administration.
The strategy of most systemic experimental therapies for restenosis is to
achieve arterial localization following the systemic delivery of a pharmaceutical
agent. The single unique approach of systemic administration aimed at a systemic
target, termed “biological targeting,” is the BPs-encapsulated nano-carriers. This
latter approach was developed in view of the paradigm change that restenosis is a
systemic disease manifested in local hyperplasia, which calls for a systemic intervention. Encapsulated BPs are targeted to deplete circulating monocytes, which are
converted into macrophages at the site of injury, and play an important role in
the course of restenosis. Such targeting requires specific formulation properties,
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including relatively larger particle size and high surface charge. Both liposomes and
NP were found to be effective carriers for targeting of BP to monocytes, showing
highly promising results as an antirestenosic therapy in rat and rabbit restenosis
models.
Nano-carriers with their diverse size, high drug entrapment capacity, the
ability to control their characteristics, and drug release profile present an advantageous delivery system for therapeutic agents in restenosis, suitable for both local
and systemic administration. The systemic approach, if addressed properly, could
provide significant advantages over local drug delivery.
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16
Ocular Applications of Nanoparticulate
Drug-Delivery Systems
Annick Ludwig
Department of Pharmaceutical Sciences, University of Antwerp,
Antwerp, Belgium
INTRODUCTION
The absorption of topically applied ophthalmic drugs is very poor because of
efficient mechanisms protecting the eye from harmful materials and agents. These
protective mechanisms, such as reflex blinking, lachrymation, tear turnover, and
drainage, result in the rapid removal of foreign substances from the eye surface.
Moreover, the very tight epithelium of the cornea also compromises the permeation of drug molecules. Consequently, the contact time of conventional eye drops
is only about five minutes, and typically less than 5% of the applied dose penetrates
passively across the cornea and reaches the intraocular tissues (1).
Thus, frequent instillations of eye drops are necessary to maintain therapeutic
drug level in the tear film or at the site of action. However, the frequent use of highly
concentrated drug solutions may induce toxic side effects after absorption via the
blood vessels in the conjunctival stroma or nasal mucosa into systemic circulation.
Moreover, cellular damage at the ocular surface could occur (2–4).
The rationale for the development of various particulate systems for sustained
drug delivery is based on possible entrapment of the particles in the ocular mucus
layer covering the eye surface, and the interaction of bioadhesive polymer chains
with mucins. This will increase the precorneal residence time of the drug, allowing
for an extension of the absorption time. Furthermore, by controlling drug release
and enhancing drug absorption, this could also improve corneal drug penetration.
One attractive feature of utilizing nanoparticles for ocular drug delivery is that it is in
a liquid dosage form which can be easily administered by patients (4–8).
Another significant challenge is to deliver therapeutic doses of drugs to treat
diseases affecting the posterior segment of the eye. It is difficult to deliver drugs to
the posterior segment by topical application because of the diffusional distance and
the counter-directional intraocular convection from the ciliary body to Schlemm’s
canal. The most logical way is to deliver drugs by intraocular injections thus bypassing anatomical and physiological barriers. However, drugs of short half-lives (e.g.,
antibiotics, antiviral drugs) would require repeated injections, which could increase
the risk of retinal detachment or hemorrhage. Therefore, biodegradable nanoparticles were developed for intraocular administration in order to obtain a controlled,
sustained drug release and thus reducing the number of injections required (9).
The choice of the polymer for preparing particulate systems will depend on
the physicochemical properties of the drug. The goal is to achieve high drug loading in order to minimize the required volume to be instilled into the eye. Corneal
uptake and transport of nanoparticles should be facilitated by small particle size
(100–200 nm) (8).
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Ludwig
Drug-release kinetics are regulated by the composition and preparation procedure of the particles, the molecular weight and degradation of the polymers, and
the physicochemical properties of the entrapped drug molecule (4,6–8). As excellent
reviews on the use of nanoparticles in ocular drug delivery have been published
previously, this chapter will mainly focus on the latest relevant research reported in
literature (5–9).
POLYMERS OF NATURAL ORIGIN
As they are biocompatible, polymers of natural origin such as proteins and polysaccharides have been investigated for use in the production of micro- and nanoparticles. A denaturation process induced by either heating or cooling and subsequent
chemical cross-linking procedures to create a denser particle matrix prepare protein
nanoparticles. Another preparation method is based on desolvation of the macromolecules, which leads to precipitation or the formation of a coacervate phase. A
cross-linking agent (e.g., glutaraldehyde) is added to harden the native particles
(5,6,10). Owing to the presence of charged groups, protein nanoparticles can be used
as a matrix in which the drug molecules may be physically entrapped or covalently
linked (11). Some examples of particles for ocular purpose reported in literature are
summarized in Table 1.
Various in vivo studies in rabbits reported that a prolonged effect of drugs
(pilocarpine, piroxicam) incorporated in albumin particles was observed when
compared to commercial preparations or aqueous and viscous solutions (4,5,11).
TABLE 1 Types of Ocular Particulate Dosage Forms Used in Animal Studies Which Were
Prepared from Proteins and Polysaccharides
Polymer
Drug
Albumin
Piroxicam
Albumin
Chitosan
Chitosan
Chitosan PEO insert
Pectin
Starch acetate
Carrageenan gelatin
Observations
Increased bioavailability compared to
commercial eye drops
Ganciclovir
After intravitreal injection in rats, the
residence time is prolonged (2 weeks);
no inflammation; good tolerance
Cyclosporin A Therapeutic concentrations in cornea and
conjunctiva during at least 48 hrs
Fluorescent
High amounts of nanoparticles into corneal
label
and conjunctival epithelia
Ofloxacin
Cmax in aqueous humor increased
compared to plain PEO inserts
Piroxicam
2.5-fold higher bioavailability in aqueous
humor compared to commercial eye
drops
—
After 3 hrs incubation, 8% of cultured
retinal epithelial cells took up
microspheres
Timolol
5.6- and 2.5-fold higher bioavailability in
aqueous humor compared to commercial
eye drops and in situ gelling system,
respectively
Note : Unless as indicated, in vivo tests were performed on rabbits.
Abbreviation: PEO, poly(ethylene oxide).
References
(12)
(13)
(14)
(15)
(16)
(17)
(18)
(19)
Ocular Applications of Nanoparticulate Drug-Delivery Systems
273
However, in one study, topical application of hydrocortisone-loaded albumin
particles in rabbits resulted in a lower tissue concentration of the drug as compared
to the application of a simple drug solution (5). This result is possibly due to the
strong binding of the drug to the particles. The retention of nanoparticles was found
to be higher in inflamed tissue as compared with the normal tissue (5).
Ganciclovir, the most widely used antiviral drug for the treatment of
cytomegalovirus retinitis, was formulated in nanoparticles for intravitreal administration in rats. A prolonged residence in the vitreous cavity (two weeks) showed no
evidence of inflammatory reaction in the retinal tissue and also did not affect the
vision (13).
The types of polysaccharides which have been investigated for the production
of ocular particulates are chitosan, pectin, and carrageenan.
Chitosan, a deacetylated chitin, is biodegradable, biocompatible, and nontoxic
(7,8). The polycationic chitosan was investigated as an ophthalmic vehicle because
of its probable superior mucoadhesiveness due to its ability to produce molecular
attractive forces by electrostatic interactions with the negative charges of the mucus.
Numerous studies indicated an increase in the precorneal retention time of chitosan
solutions (7). Moreover, chitosan appears to enhance the permeability of the cornea
transiently due to the interaction with tight junction structures (7).
For the preparation of chitosan nanoparticles, a number of techniques are
employed. These include covalent chemical cross-linking, ionic cross-linking (e.g.,
ionotropic gelation with tripolyphosphate), or desolvation. Chitosan microspheres
have been prepared by water-in-oil evaporation and spray-drying. By adjusting the
molecular weight and the deacetylation degree of the polymer used, one could
achieve the degradation and release rate required for a specific drug (7).
In one study, the animals treated with cyclosporine A-loaded nanoparticles
were found to show significantly higher corneal and conjunctival drug levels than
those treated with a suspension of cyclosporin A in a chitosan aqueous solution or
in water (2–6-fold increase) (14). In a recent study in rabbits, De Campos and coworkers showed that the amounts of fluorescent nanoparticles in cornea and conjunctiva
were significantly higher than those for a control solution. These amounts were
fairly constant for up to 24 hours. A higher retention of chitosan nanoparticles in the
conjunctiva as compared to that in the cornea was observed. Confocal microscopy
studies suggest that nanoparticles penetrate into the corneal and conjunctival
epithelia by a paracellular/transcellular pathway which is different from the pathway used by the poly(alkylcyanoacrylate) (PACA) and poly(epsilon-caprolacton)
(PECL) nanoparticles (7,15).
To improve the release kinetics of ofloxacin, chitosan microspheres loaded
with ofloxacin were also incorporated in poly(ethylene oxide) (PEO) inserts. The in
vivo results in rabbits did not demonstrate a biopharmaceutical improvement
compared to plain PEO inserts. However, a significant increase in the Cmax value in
the aqueous humor was observed which could partly be due to the permeabilityenhancing effect of chitosan across the cornea (16).
Pectin nanoparticles loaded with fluorescein showed an increase in the
precorneal retention time from 0.5 to 2.5 hours, when compared with a fluorescein
solution. In vivo tests with piroxicam-loaded pectin particles indicated a 2.5-fold
increase in the amount of drug concentration in the aqueous humor as compared
to a commercial eye drop solution (17).
Tuovinen et al. (18) reported the first study on enzyme-sensitive microparticles made of the natural biodegradable potato starch acetate for retinal targeting.
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Ludwig
After a three-hour incubation, about 8% of the cells of a retinal pigment epithelium
(RPE) culture took up microparticles. The viability of cultured RPE cells was at least
82% after 24-hour incubation with the microparticles. The degradation of potato
starch acetate in RPE cells is catalyzed by esterases and amylases. Considering the
low toxicity, it seems that these microparticles are suitable for drug delivery to the
posterior segment of the eye (18).
Carrageenans are a group of water-soluble sulfated galactans extracted from
brown seaweed. Microspheres containing gelatin, lambda carrageenan, and timolol
were prepared by spray-drying. The different ratios of carrageenan and gelatin
proved to be useful in modulating the drug-release profiles and mucoadhesive
properties. After administration of a microsphere (50/50 mixture) suspension, the
bioavailability (AUC values) of timolol in the aqueous humor in rabbits were 5.6
and 2.5 times higher in comparison with commercial eye drops and in situ gelling
system, respectively (19).
ACRYLATES
Poly(alkylcyanoacrylates)
In the past, two biodegradable and biocompatible polymers, PACAs and poly(alky
lmethacrylates), were quite popular for use in the preparation of drug carriers of
particle size range from 200 to 500 nm. The use of these polymers is based on their
mucoadhesive or bioadhesive properties. The difference between the two polymers
lies mainly in their degradation rate: polymers with a longer side chain are degraded
more slowly, resulting in a slower release of the drug incorporated in PACA
particles (4 –6).
The drug is either incorporated in PACA particles during the polymerization
process or adsorbed to the nanoparticle surface after the particle is formed. The ultimate particle size of the particles and the degree of drug loading are dependent on
several preparation parameters such as the pH of the preparation medium and the
type of stabilizer or surfactant used (4–6,8).
PACA nanospheres were shown to adhere to the eye surface, and were taken
up by a transcellular pathway in the outer cell layers of conjunctiva and cornea. This
could be explained by either endocytosis or lysis of the cell wall induced by particle
degradation products. Although the concentration of the drug in ocular tissues was
shown to be higher in inflamed tissues, a condition where the permeability of the
cell membrane is increased, no full penetration across the cornea into the anterior
chamber was observed (5).
Many studies confirmed that the use of nanoparticles with these polymers can
improve the clinical effects of a therapeutic agent (i.e., in glaucoma therapy or
antibiotic therapy) and also minimizes its side effects (4 –6). Sustained drug release
from the polymer matrix and prolonged therapeutic effect were observed, except in
the situation where the drug had a high affinity for the polymer. The increased biological response was attributed to improved ocular penetration and bioadhesion.
The precorneal residence time of PACA nanoparticles could further be increased by
incorporating the particles into a polyethylene glycol (PEG) gel or coating the particles with PEG (20,21). Acyclovir-loaded, PEG-coated polyethyl-2-cyanoacrylate
showed a 25-fold increase in the drug level in aqueous humor when compared with
an aqueous solution of the free drug. This result is probably due to a longer interaction of the nanoparticles with the corneal epithelium and the penetration-enhancing
effect of PEG (21).
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Ocular Applications of Nanoparticulate Drug-Delivery Systems
Acrylate Derivatives
Polyacrylic acid (PAA) and carbomers are used in the preparation of viscous eye
drops, artificial tears, hydrogels, and inserts. The mucoadhesive properties of these
polymers are mainly due to hydrogen bonding and interpenetration of the polymer
chains and the mucus layer at the eye surface (4).
The hydration state and pH of microspheres in the lacrimal fluid were shown
to be a factor affecting the residence time of the microspheres in the eye. In one
study, radiolabeled PAA microspheres made with Carbopol® 907 cross-linked
with maltose was applied either in dry form or in a prehydrated form. The clearance rate of the microspheres applied in the dry form was found to be higher than
the prehydrated form probably due to incomplete hydration in the lachrymal
fluid. Furthermore, the clearance of hydrated microspheres at pH 7.4 is higher
than at pH 5.0 because the presence of protonated carboxyl groups at pH 5.0 in
PAA permits enhanced bioadhesion through hydrogen bond formation with
mucins (Table 2) (22).
In a recent study, De et al. (28) demonstrated the biocompatibility and adhesive
properties of PAA nanoparticles on human corneal epithelial cells. A controlled
release of the partly ionically entrapped brimonidine was obtained due to the
cation-exchange properties of the polyanionic PAA matrix of the nanoparticles (28).
A major problem of PACA nanoparticles is the low loading capacity for
hydrophilic drugs. To improve drug loading, attempts are made to increase the
hydrophilicity of the particle surface by copolymerization of methylmethacrylate
(MMA) and sulfopropylmethacrylate (SPM). Bound drug molecules are released
from the carrier by competitive replacement by other ions. These copolymer
nanoparticles loaded with the muscarinic agonists arecaidine propargyl ester (APE)
and (S)-(+)aceclidine were evaluated in rabbits. A twofold increase in drug absorption was obtained when the nanoparticles were instilled together with bioadhesive
hyaluronan (23).
Various cationic acrylic copolymers were also examined to prepare nanoparticles
with a positive charge in order to facilitate an effective adhesion of the delivery system
to the ocular surface (8). Pignatello et al. (24–26) formulated nanoparticles with
TABLE 2 Overview of In Vivo Studies Carried Out in Rabbits Using Nanoparticles Prepared
with Acrylate Derivatives
Type of polymer
Drug
Carbopol 907
111Indium
Copolymers
MMA–SPM
Eudragit RL100
and RS100
Propargyl ester
(APE)
(S)-(+)aceclidine
Ibuprofen;
Flurbiprofen
PNIPAAm
Epinephrine
Observations
References
Faster elimination of microspheres in dry
form than when instilled as a
dispersion
Combination with mucoadhesive
polymers increased bioavailability by
twofold
Higher drug levels in the aqueous humor
compared to aqueous solution due to
sustained release and increased
precorneal retention
Eightfold longer decrease of IOP after
administration of cross-linked particles
compared to eye drops
(22)
(23)
(24–26)
(27)
Abbreviations: APE, arecaidine propargyl ester; IOP, intraocular pressure; MMA, methylmethacrylate; PNIPAAm,
poly-N-isopropylacrylamide; SPM, sulfopropylmethacrylate.
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Ludwig
Eudragit® RL and RS using a modified quasiemulsion solvent diffusion technique and
solvent evaporation method. Using these nanoparticles, they demonstrated that sustained release and increased absorption of the incorporated nonsteroidal, antiinflammatory drugs were achieved. Furthermore, no inflammation or discomfort was observed
in the rabbit’s eye, suggesting that the nanoparticles were well tolerated (24).
Hsiue et al. (27) investigated the use of a thermosensitive polymer, poly-Nisopropylacrylamide (PNIPAAm), in controlled-release delivery systems for
glaucoma therapy. Cross-linked PNIPAAm nanoparticles containing epinephrine
were administrated to rabbits, and its effect on the intraocular pressure (IOP) was
monitored. The results indicated that a decreased pressure response which lasted
eight times longer than that of conventional eye drops was observed.
POLY(EPSILON-CAPROLACTON)
Another biocompatible and biodegradable polymer is PECL. It is slightly more
hydrophobic when compared with PACA (6,29). PECL nanoparticles and nanocapsules for ophthalmic use can be prepared by solvent extraction or solvent
evaporation method (6).
Upon instillation, the PECL particles form aggregates and reside in the
cul-de-sac, gradually releasing the drug. This hypothesis was put forward to
explain the much more pronounced reduction in IOP in glaucomatous rabbits
after administration of PECL nanoparticles when compared to PACA and PLGA
microspheres (4–6). PECL nanoparticles also seem to be able to penetrate the outer
corneal cell layers, but contrary to PACA, no cellular damage was observed. This
suggests that an endocytic mechanism may be involved. The penetration of the
particles is found to be size-dependent: nanoparticles but not microspheres were
found in the corneal cells (4–6).
Numerous in vivo studies demonstrated that enhanced corneal absorption
and prolonged therapeutic effects of drugs are possible when the drugs are incorporated in PECL nanoparticles. The drugs tested were: betaxolol, metipranolol,
carteolol, cyclosporin A, and indomethacin. Furthermore, nanocapsules seem to
display a better effect than nanospheres, probably because the drug entrapped was
in the unionized form in the oily core of the delivery system and could diffuse more
easily to the cornea (4–6,8).
A strategy designed to enhance the interaction with the mucus layer in the eye
has been investigated by Calvo et al. (30) who coated PECL particles with cationic
bioadhesive polymers. Compared with noncoated particles, the corneal and aqueous
humor indomethacin concentrations were doubled for chitosan-coated nanocapsules.
However, coating with polylysine did not seem to improve the bioavailability of the
tested drug. The rise in bioavailability was attributed to the structural similarity
between chitosan and mucin, rather than to its positive charge (30). Chitosan-coated
PECL particles exhibit an important interaction with the mucus layer, but the penetration in the corneal epithelium is lower when compared with that of the uncoated
particles (7).
De Campos et al. (31) compared the effect of coating PECL nanocapsules with
PEG versus chitosan on the interaction of PECL nanocapsules with the ocular
mucosa. The in vivo study showed that the nanoparticles entered the corneal epithelium by a transcellular pathway. The penetration rate and depth were dependent
on the coating composition. PEG coating enhanced the passage of the PECL particles
277
Ocular Applications of Nanoparticulate Drug-Delivery Systems
nanocapsules across the whole epithelium, whereas the chitosan coating favored the
retention in the superficial layers of the epithelium (31). Consequently, the design of
colloidal carriers with different ocular distribution seems to be possible.
POLY(D,L-LACTIC ACID) AND POLY(D,L-LACTIDE-CO-GLYCOLIDE)
As poly(d,l-lactic acid) (PLA) and poly(d,l-lactide-co-glycolide) (PLGA) are
FDA-approved products (additives) for parenteral use, microspheres and nanoparticles made of these polymers were investigated primarily for the controlled
release of drugs after intravitreal or subconjunctival injection rather than for
topical application (4,5,9,29).
PLA is a synthetic, biocompatible, and biodegradable polymer and PLGA a
copolymer of poly(lactic acid) and poly(glycolic acid). The degradation rate depends
on the molecular weight, conformation, and polymer composition (9,29,32). The
drug is released out of the spheres by diffusion and by hydrolysis of the PLA/PLGA
matrix. In an attempt to optimize the mucoadhesive properties, PLA can be
copolymerized with other polymers (e.g., PECL or PEG), so that more appropriate
“tailor-made” polymers can be developed (33).
Several methods have been proposed for the preparation of the PLA/PLGA
particles. However, the most popular technique is the emulsification solvent
evaporation method (9,32). The uptake of PLGA nanoparticles in conjunctival
epithelium is dependent on the particle size: 100 nm particles exhibited the highest
uptake as compared to larger particles (800 nm and 10 µm). The saturable particle
uptake occurs via adsorptive-mediated endocytosis which is independent from
clathrin- and caveolin-1-mediated pathways (8,34).
Some examples of PLA/PLGA particles for ocular purpose reported in
literature are summarized in Table 3. Giannavola et al. changed the surface
properties of PLA nanoparticles loaded with acyclovir by incorporating PEGylated
1,2-distearoyl-3-phosphatidylethanolamine (DSPE-PEG) into the polymer instead
of coating the external surface as in the case of PACA nanoparticles (22,35). After
TABLE 3 Evaluation of Poly(D,L-Lactic Acid) and Poly(D,L-Lactide-Co-Glycolide) Particles for
Ophthalmic Use in Animal Studies
Drug
Application
Acyclovir
Instillation
5-Fluorouracil
Conjunctival
implantation
Instillation
Vancomycin
rhVEGF
Budesonide
Subretinal and
intravitreal in rats
Subconjunctival
injection in rats
Observations
References
A 12.6-fold increase of aqueous humor
AUC for PEG-coated nanoparticles
compared with drug suspension
Therapeutic scleral levels during 7 days
Low ocular toxicity and no irritation
A twofold increase of aqueous humor AUC
with respect to drug solution and
sustained release
Dose-dependent angiogenic response
(35)
Sustained drug level in retina and other
ocular tissues during 14 days
(39)
Note: Unless otherwise indicated, all in vivo tests were performed on rabbits.
Abbreviation: AUC, area under the curve.
(36)
(37)
(38)
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Ludwig
administration of the nanoparticle suspension in the rabbit’s eye, the AUC0–6
values in the aqueous humor were found to be significantly greater for the PEGcoated PLA nanospheres than for the uncoated particles. A drop of 48% in AUC
value was observed when PEG-coated particles were administrated to the eye
after removal of the mucus layer. This decrease was attributed to the absence of
PEG–mucin interactions. Thus, differences in bioavailability of the drugs could be
correlated with the different interactions between the nanoparticle’s surface and
the corneal epithelium.
Higher and prolonged vancomycin concentration was shown in the aqueous
humor by incorporating the drug in PLGA microspheres. However, increasing the
viscosity of the microsphere suspension by adding a viscosifying agent (hydroxypropyl cellulose) did not seem to improve any further drug absorption, probably
because of the rapid ocular clearance of the highly hydrophilic drug. No irritation
or ocular discomfort was observed in rabbits (37).
Microspheres have been investigated for use in delivering many ophthalmic
drugs for intravitreal or subconjunctival injections. This includes retinoic acid,
adriamycin, 5-fluorouracil, dexamethasone, cyclosporin A, and ganciclovir (9,33). It
has been shown that intracellular drug delivery to retinal epithelial cells may be
possible by coating the PLA microspheres with gelatin, probably via phagocytosis
(9). Recent studies demonstrated that PLA/PLGA microspheres may be useful in
sustained or even extended slow drug delivery targeted to the posterior segment of
the eye for the treatment of chronic diseases or gene therapy (38,39).
As the integrity and activity of proteins and oligonucleotides is preserved
when encapsulated within PLA, the nanoparticles were evaluated for delivering
these molecules to the retina (9). After intravitreal injection in rats, transretinal
movement of the nanoparticles with a preferential localization in the retinal epithelial
cells was observed. A mild transient inflammatory reaction after injection was
reported. The presence of nanoparticles within the retinal epithelium cells could be
detected even after four months of a single injection. This suggests that a steady and
continuous delivery of drugs could be achieved (40). However, interference with
visual acuity due to floating of nanoparticles in the vitreous cavity could be a drawback. Before clinical implementation of the nanoparticles, the possible effects of the
nanoparticles on the retinal function and the vision have to be investigated (9,40).
LIPIDS
Instead of using macromolecules, several lipids were proposed for use to prepare
nanoparticles. Solid–lipid nanoparticles (SLNs) consist of a biocompatible lipid
core and an amphiphilic surfactant as an outer shell. The advantages of SLNs are:
scalable manufacturing process using hot or cold high-pressure homogenization,
easy modulation of the drug-release profile, no organic solvents, the wide range of
lipid/surfactant combinations, and the high drug payload (41). Cavalli et al. (42)
evaluated the use of SLNs as carriers for tobramycin. Compared to commercial eye
drops, the tobramycin-loaded SLNs produced a significantly higher bioavailability:
a 1.5-fold increase in Cmax and fourfold increase in AUC. The SLN dispersion was
well tolerated, with no evidence of ocular irritation (42). The longer retention
observed for SLNs on the corneal surface and in the cul-de-sac is probably related to
their relatively small size. The nanoparticles are presumed to be entrapped and
retained in the mucus layer.
Ocular Applications of Nanoparticulate Drug-Delivery Systems
279
CONCLUSIONS
The in vivo studies that were used to evaluate the use of nanoparticles for ocular
drug delivery were primarily performed on animals. From the results to date, one
could conclude that the nanoparticles should possess bio- or mucoadhesive properties in order to achieve a long precorneal retention time and to improve drug
absorption. The intraocular use of biodegradable, slow-releasing nanoparticles
looks very promising for drug delivery targeted to the tissues of the posterior
segment in order to treat chronic diseases or for gene therapy.
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17
Nanoparticulate Systems for Central
Nervous System Drug Delivery
Jean-Christophe Olivier and Manuela Pereira de Oliveira
Pharmacologie des Médicaments Anti-Infectieux, Faculty of Medicine and
Pharmacy, and INSERM, ERI 023, Poitiers, France
INTRODUCTION
The central nervous system (CNS) is isolated from the whole body by the blood–
brain barrier (BBB) which creates a strictly controlled extracellular fluid
environment protecting the brain parenchyma from the blood composition
variation and from blood-borne potentially CNS-toxic compounds. This tight
physiological barrier limits drastically drug diffusion towards the brain parenchyma and is considered as the bottleneck in brain drug development and as an
important limiting factor for the future applications of biotechnology-derived
neurotherapeutics (1). The BBB can be circumvented by intraventricular or intracerebral administration. These invasive and risky techniques are limited to the
treatments of restrained brain areas due to the poor tissue diffusion of injected
materials from the administration sites. Postsurgical administrations are the usual
administration routes for relatively large drug-delivery systems (DDSs), such as
microspheres and wafers, and are presently limited in clinical practice to the
adjuvant treatment of brain tumors after surgical resection (2,3). The noninvasive
access from the blood compartment via the BBB is the most convenient and safest
way to treat the entire brain space. There are indeed around 400 miles of blood
capillaries in the average 1300-g human brain, which constitute a large interface of
approximately 20 m2 surface area (150 cm2/g) for blood-to-brain exchanges.
Nanoparticulate DDSs (nano-DDSs), mostly liposomes and nanoparticles, have
been investigated for the brain delivery of therapeutic agents which poorly diffuse through the BBB. Owing to their nanometric size, these nanocontainers freely
circulate in blood capillaries and can be conveniently administered by the intravenous route. Table 1 summarizes the ideal properties that nano-DDS should possess
for brain drug delivery. The transit within the blood compartment requires the
nano-DDS to escape from the mononuclear phagocyte system (MPS). Once within
the brain vasculature, nano-DDSs have to trigger their translocation through the
BBB to deliver their content in the brain. This chapter reviews the various parameters that should be considered when designing nano-DDSs for brain delivery and
the present achievements in the field.
THE BLOOD–BRAIN BARRIER AND THE NEED FOR BRAIN-SPECIFIC
NANO-DRUG DELIVERY SYSTEMS
Normal Blood–Brain Barrier
A scientific consensus locates the BBB at the endothelia of brain capillaries which
are in close relationship with the surrounding pericytes, actrocytes, neurons, and
glial cells (Fig. 1). To summarize, the BBB results from the unique properties of
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TABLE 1 Ideal Properties of Nano-Drug Delivery Systems for Drug Delivery Across
the Blood–Brain Barrier
General for drug delivery
Nontoxic, biodegradable, and biocompatible
Amenable to small molecules, peptides, proteins, or nucleic acids (genes, antisense drugs)
Minimal nano-DDS excipient-induced drug alteration (chemical degradation/alteration,
protein denaturation)
Modulation of drug-release profiles
Scalable and cost-effective manufacturing process
General for free circulation in blood
No capillary filtration: particle diameter <1 µm
Physical stability in blood (no aggregation, no dissolution, no interaction with blood cells)
Avoidance of the MPS, prolonged blood circulation time
Particular for brain delivery
Noninvasive delivery from the blood circulation into the brain parenchyma via the BBB without
inducing BBB alteration, therefore by a transcytotic pathway
Particle diameter <100 nm to fit the loading capacity of transcytotic vesicles
Extra- or intracellular drug delivery to subset of brain cells or brain tumor cells
Abbreviations: BBB, blood–brain barrier; DDS, drug delivery system; MPS, mononuclear phagocyte system.
endothelial cells (1,4): (i) the tight junctions that connect adjacent endothelial cells
and physically restrict solute flux between the blood and the brain with the consequence that passive diffusion towards brain is limited to small lipophilic compounds
(optimal log Po/w is 1–3) of molecular weights below 400 to 500 Da, (ii) an elaborate
system of transport proteins that allows the selective influx transport of hydrophilic
solutes (generally by carrier-mediated transport or CMT) and macromolecules
(by receptor-mediated transcytosis or RMT) necessary for CNS maintenance and
extracellular fluid homeostasis, and that reject potentially CNS-toxic compounds or
metabolites [by active efflux transport (AET) proteins], both influx and efflux
transport proteins being the cause of the observed deviations from the precept of
lipophilicity-based solute penetrability into the brain, (iii) a metabolic barrier which
serves as a biotransformation and detoxification system, and (iv) a negligible pinocytotic activity. The high transendothelial electrical resistance value (∼2000 Ω cm2)
measured across the brain capillary endothelium is indicative of the very low ionic
permeability. As a consequence, BBB accounts for the restricted CNS access of more
than 98% of all small-molecule drugs and of 100% of large-molecule drugs (4).
Blood–Brain Barrier Alteration in Pathology
In various brain pathologies, including brain trauma, stroke, septic encephalopathy,
or neurodegenerative diseases, or in metabolic disorders such as diabetes mellitus,
the alterations of the BBB permeability result in leakage of normally restrained
plasma components and contribute to neuroinflammation and neuronal damage
(5). It is still controversial whether the BBB permeability increase results from the
opening of passageways through brain endothelial cells, defined as vesiculo-tubular systems or vesiculo-vacuolar organelles (6), or from tight junction degradation
(5). Little is known on the degree of BBB permeabilization in inflammatory brain
pathologies, but its actual impact on nano-DDS passive diffusion is likely to be low.
In a rat model of cerebral ischemia and reperfusion, polystyrene nanoparticles
(20 nm in diameter) were shown to extravasate into the brain interstitial fluid (7). In
the case of larger nano-DDSs such as liposomes, the uptake by infiltrating
Nanoparticulate Systems for Central Nervous System Drug Delivery
astrocyte foot
process
endothelial cell
Pinocytotic
vesicles
M
1
nucleus
mitochondria
blood
A
+
+
+
+
+
+
M
D
AET
+
+
+
+
+
2
B
CMT
-+- - +- +
-+
+
-+ +
+
+
+
+
+
+
AMT
+
Y
Y
-
Y
Y
Y
Y
Y
AMT
Y
RMT
Y
RMT
Y
Y
Y
Y
Y
tj
+
Y
4
C
Y
+
--
Y
RMT
Y
3
283
Y
Y
Y
Y
extracellular matrix
Y
Y
pericyte
axonal
ending
FIGURE 1 Schematic diagram of the neurovascular association forming the BBB together with the
molecular transport pathways across the BBB (1–4) and proposed transcytotic brain-delivery mechanisms for nano-DDSs (A–D). Brain endothelial cells are characterized with tight junctions (tj) that
lock the aqueous paracellular pathways, negligible pinocytotic activity, high transport, and metabolic
activity. They are in close relationship with pericytes, astrocyte foot processes, axonal endings, and
microglial cells (not represented). Passive diffusion across the endothelial cells is the passageway
of small lipophilic molecules (Mw below 400–500 Da) (1). Restrained by tight junctions and the
plasma membranes, hydrophilic compounds require specialized transport to reach brain: CMT systems for small molecules (2), RMT for macromolecules (3), or absorptive-mediated transcytosis
(AMT) for cationic molecules (4). Compounds traversing endothelial cells may be rejected back into
the blood compartment by AET systems as intact molecules or as metabolites (M). Proposed transcytotic pathways for nano-DDSs are AMT triggered by positively charged surfaces (A) or RMT triggered by natural substrate ligands such as transferrin (B) or peptidomimetic antibodies directed
against exofacial epitopes of transferrin or insulin receptors (C). The hypothetic “differential protein
adsorption mechanism” according to which the nano-DDS surface would adsorb, from blood, natural substrates of BBB receptors (e.g., apolipoproteins B or E, substrates low-density lipoprotein receptors) and then undergo endocytosis/transcytosis is also represented (D). Abbreviations: AET, Active
efflux transport; AMT, absorptive-mediate transcytosis; BBB, blood–brain barrier; CMT, carrier-mediated
transport; DDSs, drug delivery systems; RMT, receptor-mediated transcytosis.
macrophages is probably the mechanism for their increased distribution into
inflammatory brain tissues (8).
In the case of brain tumors, the BBB integrity is locally compromised by the
absence of tight junctions, allowing for tumor core penetration and retention of
drugs, macromolecules, or nano-DDS otherwise excluded from normal brain (9).
This phenomenon is known as the enhanced permeability and retention (EPR) effect
of tumors. Generally, growing tumor margins and adjacent normal tissue remain
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Olivier and Pereira de Oliveira
unreachable. In experimental rat brain tumors, the vascular pore cutoff size determined with long-circulating liposomes or microspheres ranged from 100 to 550 nm,
depending on the tumor cell line (10). By optimizing the EPR effect or by decreasing
peripheral distribution of antitumor agents, nano-DDS formulations were generally
more efficient than solutions. In rats, doxorubicin-loaded nanoparticles or liposomes
increased delivery to experimental brain tumors (11,12), leading to higher efficiency
(12). Doxorubicin formulated as pegylated liposomes (marketed under Doxil® or
Caelyx® trademarks) showed, however, a moderately increased efficiency against
high-grade gliomas in clinical practice, compared to the solution (13). Using confocal
imaging, MR imaging, and histological examination of experimental rat brain tumors,
Straubinger et al. (14) showed that intravenously administered liposomes lined
tumor capillaries or blood vessels and poorly spread within tumors. Under repeated
administrations of Doxil® liposomes, extensive regions of hemorrhage within the
tumor tissue occurred, suggesting a destruction of tumor vasculature or underlying
tumor cells. The opening of the tumor stroma would permit subsequent doses of
liposomes to diffuse more deeply into tumors and finally exert their cytotoxic
effect on margins. Uptake efficiency and intracellular delivery may be improved
with tumor-specific nano-DDSs, but the tumor cells that infiltrate the normal
brain parenchyma remain protected by intact BBB. The combination of nano-DDSs
able to cross BBB with tumor-specific nano-DDSs may be a means to optimize
antitumor therapy.
Candidate Drugs for Brain-Targeted Delivery with
Nano-Drug Delivery Systems
Generally, small-molecule drugs can be chemically designed or modified (e.g., by
prodrug synthesis) to be adequately lipophilic for passive diffusion through the BBB
and do not need DDSs for brain delivery. However, drug lipidization is not always
applicable or effective, especially when drugs are particular chemical entities that
cannot be modified without losing their pharmacological activities and/or are
substrates of efflux transport proteins present at the BBB, such as the P-glycoprotein.
Furthermore, drug lipidization generally results in increased metabolism and peripheral distribution, which necessitates higher doses, potentially at the cost of more
frequent adverse reactions. In such cases, or when small drug molecules undergo
metabolization in brain endothelial cells, nano-DDS formulations should be considered as a means for improving brain delivery. Another group of molecules that
necessitates nano-DDS formulations are the new large-molecule therapeutics potentially efficient to treat CNS: peptides, proteins, such as neurotrophic factors, antisense
drugs, or genes (plasmids). Owing to their poor stability in biological fluids, rapid
enzymatic degradation, unfavorable pharmacokinetic properties, and lack of diffusion towards the CNS, they could be advantageously formulated in brain-targeted
protective nanocontainers. Compared to conventional drugs, they possess a high
intrinsic pharmacological activity. The small dose needed for therapeutic efficiency
would easily fit the loading capacity of nano-DDSs and avoid the administration of
large amount of potentially toxic nano-DDS excipients. The choice between liposomes or nanoparticles will depend on formulation aspects, release kinetics, and
stability upon storage. Proteins are generally unstable in the presence of organic
solvent or solid interfaces. It may, therefore, be challenging to maintain their biological activities upon entrapment within nanoparticles that need solvents (polymeric
NP) or heating steps (lipid NP) for preparation. Formulation additives may improve
their stability (15), or liposome formulations may be more appropriate. Some drugs
Nanoparticulate Systems for Central Nervous System Drug Delivery
285
will require release into the brain interstitial fluid to be effective. Others, such as
plasmid genes or antisense oligonucleotides, need to be delivered intracellularly for
effectiveness, which means that after crossing the BBB nano-DDS should be also able
to cross plasma membranes of targeted brain cells.
NANO-DRUG DELIVERY SYSTEMS FOR NONINVASIVE
DRUG BRAIN DELIVERY
Basic Principles
According to the pharmacokinetic rule, the percent of injected dose of a drug that is
delivered per gram brain (%ID/g) is directly proportional to the BBB permeability–
surface area (PS) product and the area under the plasma concentration curve (AUC,
%ID min μL−1): %ID/g = PS × AUC (1). This rule also stands for nano-DDS brain
delivery. For optimal efficiency, nano-DDSs administered IV should remain in the
blood compartment, avoiding useless peripheral distribution, in order to reach the
brain vasculature and should possess an appropriate BBB PS in order to deliver
their content beyond the BBB into the brain parenchyma. Owing to their size, nanoDDSs cannot cross the endothelium of brain capillaries by passive diffusion through
normal BBB. Although effective for promoting liposome delivery into the brain
parenchyma (16,17) and explored for improving antitumor drug diffusion into brain
tumors (18), increasing BBB permeability using hyperosmolar solutions or the selective B2 bradykinin Cereport (labradimil or RMP-7) should be limited to short-term
therapy, as this promotes the non-selective entry of blood-borne compounds and
may result in seizures and permanent neuropathological changes. With preserving
the BBB integrity, the delivery of nano-DDSs into the brain parenchyma implies that
nano-DDSs should not only recognize specific sites at the BBB, but also trigger their
own transfer across the endothelial cells. Indeed intravenously administered
pegylated immunoliposomes directed against brain astrocytes displayed longcirculating properties (elimination half-lives of 8–15 hours), but were unable to cross
the BBB and missed their targets (19). On the basis of an already long research, both
steps are being rationally apprehended and key parameters are well defined for
liposomes. However, despite real achievements, optimization is still needed in the
case of nanoparticles (20).
Circulating in the Blood Compartment
If not especially designed to escape from the MPS uptake, intravenously administered nano-DDSs are rapidly cleared from the blood stream (blood half-lives are
generally around two to three minutes) and mostly accumulate in liver and spleen.
It is generally admitted that opsonization, the first step for MPS recognition, is
favored by hydrophobic surfaces, which promote protein adsorption, and negative
surfaces, which are activators of the complement system (21). In contrast, hydrophilic coating sterically stabilizes nano-DDSs and reduces opsonization and MPS
uptake. These rather simplistic views should, however, be tempered, as recent
evidence showed that sterically stabilized particles were efficiently opsonized and
activated the complement system, while keeping their stealth properties (22).
Whatever the mechanism may be, the incorporation of polyethylene glycol (PEG)
derivatized lipids into the phospholipids bilayer of liposomes resulted in longcirculating liposomes, generally referred to as sterically stabilized liposomes, with
apparent terminal half-lives up to 90 hours in humans (23). Long-circulating nanoparticles are however still elusive. Pharmacokinetic studies are often contradictory, if not
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Olivier and Pereira de Oliveira
lacking. In the case of pegylated poly(lactide) or poly(lactide-co-glycolide) nanoparticles, the up-to-date most studied polymeric nano-DDSs, surface coating with PEG
molecular weights of 5000 or over, were shown to be efficient at reducing opsonization, but blood half-lives were found to be very variable ranging from less than one
to six hours (20). The variability in density-related PEG conformation may explain
the rapid clearance of a significant fraction of intravenously injected “longcirculating” nanoparticles by the MPS (22). Contrary to liposomes, the variety in
nanoparticle core nature (polymers or lipids) that characterizes the research on
nanoparticles resulted in the dispersion of research effort which was probably at the
origin of the limited achievements in the long-circulating nanoparticle design.
Nano-Drug Delivery System Transport Through Brain Endothelial Cells
The reasoned approaches presently under investigation for triggering nano-DDS
translocation across BBB are based on AMT or RMT (Fig. 1). Unexpectedly, several
research teams found that surfactant-coated nanoparticles delivered their contents
to the brain. As they were not based on theoretical concept, these methods were
referred to as “empirical” approaches.
Absorptive-Mediated Transcytosis
Cationized proteins (albumin, immunoglobulins) have been demonstrated to
undergo absorptive-mediated endocytosis (AME) through the BBB in vivo (24).
AME is triggered by an electrostatic interaction between the positively charged
protein and the negatively charged plasma membranes. It is therefore not brainspecific, and cationized proteins also accumulate in liver, kidney, and/or lung
tissues. However, the relatively high capacity of AME should be favorable to brain
delivery (25). Preservation of appropriate pharmacokinetic profiles depends on the
degree of cationization (24). As an application of such a concept, cationic bovine
serum albumin-conjugated pegylated nano-DDSs [liposomes (26) and pegylated
polylactide nanoparticles (27)] were shown to undergo endocytosis in in vitro BBB
model. It is still to be demonstrated whether these cationic nano-DDSs will possess
appropriate pharmacokinetic profiles by the intravenous administration route.
Receptor-Mediated Transcytosis
Owing to a high stereospecificity for their substrate, CMT requires drugs with
molecular structures mimicking the endogenous nutrient and is not a realistic option
for nano-DDS brain delivery (1,25). Only RMT may be used for transcytotic delivery
of nano-DDSs through the BBB. Owing to the ubiquity of RMT, relative brain specificity may be achieved by targeting receptors that are overexpressed at the luminal
side of brain endothelial cells compared to other organs, for example, the receptors
of transferrin, insulin, insulin-like growth factor (IGF), or low-density lipoproteins
(LDL) (4). Nano-DDSs should be less than 100 nm in diameter to fit the loading
capacity of these transport systems and should have on their surface receptorrecognizing ligand (natural substrate or antibody) capable of triggering transcytosis. Transferrin ligand was investigated for liposome brain targeting with promising
results (28), but an apolipoprotein E-derived peptide failed to trigger liposome
endocytosis via LDL receptor (29). Antibodies directed against exofacial receptor
epitopes that do not interfere with the natural ligand-binding sites should be preferred in order to avoid potential competition between targeted nano-DDSs and
endogenous natural ligand (transferrin, apolipoproteins), pharmacological activity
Nanoparticulate Systems for Central Nervous System Drug Delivery
287
(insulin), or linkage to plasma-binding proteins (IGF) (4). The most advanced work
based on this “Trojan horse” concept has been carried out by Pardridge and coworkers with pegylated immunoliposomes (4). The receptor-specific targeting ligands
located at the tip of 1% to 2% of the PEG2000 strands are peptidomimetic monoclonal
antibodies able to trigger the activation of transferrin or insulin receptors. After
intravenous administration, these immunoliposomes delivered their content (smalldrug molecules, plasmids) into the brain parenchyma without damaging the BBB
(30–33). Owing to the presence of transferrin or insulin receptors on the neuronal
plasma membrane, plasmid-loaded immunoliposomes permitted neuronal nuclear
delivery, resulting in gene expression in the entire brain space (31–33). As RMT
receptors are also abundant in some peripheral tissues, ectopic expression in
nonbrain organs was eliminated using a brain-specific gene promoter (33). In a
6-hydroxydopamine rat model of Parkinson’s disease, the selective gene expression
in nigrostriatal neurons resulted in the restoration of tyrosine hydroxylase activity
and in reversal of motor impairment (33). As immunoliposomes deliver their content quickly, as shown by the relatively short-lasting plasmid expression in brain
(33), they would require monthly administrations to sustain a pharmacological
effect (31). This potential inconvenience for the treatment of chronic brain disorders
may be corrected using pegylated poly(lactide) immunonanoparticles with
sustained-release properties (20,34).
Empirical Approaches
Various works reported on the beneficial effect of surfactant-coated lipid or polymeric nanoparticles for the brain delivery of drugs. No brain-targeting ligand is
present on nanoparticle surface and the various mechanisms underlying increased
brain uptake of entrapped compounds remain hypothetic.
Solid–Lipid Nanoparticles
Solid–lipid nanoparticles are basically constituted of a liquid or solid–lipid core
(liquid glyceride or wax) surrounded by a solid shell made of an association of
several anionic surfactants (polysorbates, PEG fatty acid ester, PEG fatty alcohol
ether, poloxamers) and/or ionic surfactants (lecithin, etc.) and do not (generally)
necessitate organic solvents for their preparation (35). Experimental demonstration
of SLN “stealthness” is scarce and more often based on the presupposed antiopsonic effect of PEGylated surfactant rather than on experimental evidence (36).
Several experimental works showed that surfactant-coated solid–lipid nanoparticles could increase the brain uptake of loaded drugs that poorly diffuse into the
brain when administered as solutions. A prodrug of dioctanoylfluorodeoxyuridine
had a significantly higher brain distribution when administered as pluronic
F68-coated solid–lipid nanoparticle formulations compared to the solution (37). The
authors proposed as brain uptake mechanisms either an increase in gradient
concentration between blood and brain resulting from a higher plasma AUC
(compared to the drug solution), or an endocytosis by endothelial cells. In other
studies, intravenously administered to rats or rabbits, doxorubicin “stealth” SLN
coated with PEG2000 stearate also increased brain uptake compared to the commercial solution (36,38). In an in situ rat brain perfusion model, paclitaxel brain uptake
was increased when formulated in wax nanoparticles including polysorbate 60 and
Brij 78 as surfactants, compared to the control solution (39). Control nanoparticles
were shown to undergo a significant brain uptake (40) without altering BBB permeability (41,42). Unexpectedly, nanoparticle translocation through the BBB was not
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Olivier and Pereira de Oliveira
retained as the mechanism for increased paclitaxel brain uptake, and the authors
favored a modulation of paclitaxel efflux by P-glycoprotein transporter (39).
Polybutylcyanoacrylate Nanoparticles
Adsorbed onto polysorbate-coated polybutylcyanoacrylate (PBCA) nanoparticles
administered intravenously, compounds with poor brain diffusion (doxorubicin,
loperamide, tubocurarine, the hexapeptide dalargin) were successfully delivered
to the brain where they induced pharmacological effects (for review, see Ref. 43).
As these nanoparticles do not possess targeting ligands on their surface, the concept of “differential protein adsorption” was proposed by Müller (44). According
to this concept, some blood proteins preferentially adsorb onto nanoparticle surfaces depending on their physicochemical properties. Apolipoproteins B and E
would be the brain-targeting proteins that adsorb onto polysorbate-coated
nanoparticles and permit nanoparticle endocytotic delivery into the brain vessel
endothelia via LDL receptors (45). This mechanism may also account for polysorbate 80-coated poly(lactide) nanoparticle interaction with brain microvessels (46).
However, it was shown that in rats treated with polysorbate 80-coated PBCA
nanoparticles (polysorbate 80: 25 mg/kg, nanoparticles: 50 mg/kg) inulin spaces
increased by 10% (nonsignificant) after 10 minutes and by 99% (significant) after
45 minutes compared to controls (47). A nanoparticle-induced nonspecific BBB
permeabilization, resulting from the synergistic effect of nanoparticle toxicity and
high polysorbate 80 concentrations, was proposed as an alternative mechanism to
the brain translocation of nanoparticles across the BBB (48).
CONCLUSION
Technology now exists for designing safe brain-targeted long-circulating nanoDDSs. The PEGylated immunoliposomes based on the concept of molecular “Trojan
horse” that targets the transcytotic BBB receptors and ferries the liposomes though
the BBB permitted the efficient noninvasive, nonviral delivery of plasmid genes into
the brain parenchyma cells. They constitute the present most convincing demonstration of successful brain delivery of complex macromolecules by nano-DDSs
from the blood circulation through the BBB. The empirical use of simpler polymeric
or lipid nanoparticles also resulted in promising results. The mechanisms underlying their still hypothetic translocation through the BBB need further investigations
for validation. Thus, despite the complexity of the issue, there have been real
achievements in nano-DDS brain delivery, but the way is still long from bench
results to clinical applications.
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18
Nanoparticles for Gene Delivery:
Formulation Characteristics
Jaspreet K. Vasir and Vinod Labhasetwar
Department of Pharmaceutical Sciences, University of Nebraska Medical Center,
Omaha, Nebraska, U.S.A.
INTRODUCTION
The recent advances in the field of molecular biology have highlighted gene therapy
as a promising approach for the treatment of numerous diseases including genetic
disorders, such as cystic fibrosis, hemophilia; many somatic diseases such as tumors,
neurodegenerative diseases; and severe viral infections such as AIDS. Despite the
considerable interest generated in gene therapy and the phenomenal pace at which
research has advanced, “delivery” of genes to the target cells still remains the most
formidable challenge. Polynucleotide molecules (e.g., DNA or RNA) are large,
hydrophilic, macromolecules with a net negative charge. Unlike other drugs, these
are very labile in the biological environment and do not cross biological membranes
effectively. Thus, the need for an effective and safe gene-delivery system is quite
obvious. In general, gene delivery/expression vectors can be broadly categorized into
the viral and nonviral vectors. Viral vectors include the use of genetically engineered
retroviruses, adenoviruses, adeno-associated viruses, and other viruses that have been
used for gene-transfer procedures. Although the viruses are highly efficient for gene
transfer to cells, their pote ntial to induce drastic immune responses such as with adenovirus or the risk of insertional mutagenesis in the host genome with retroviral vectors
(1) has sparked a major debate over their safety for human gene therapy.
Thus, the current consensus is to develop suitable vector systems which are
minimally invasive (safe) and highly efficient for gene therapy in humans. This has
steered research towards the development of nonviral vectors for gene delivery.
Nonviral vectors include nanoparticles (NPs), liposomes, and complexes prepared
either using cationic lipids (lipoplexes) or polymers (polyplexes), and also mechanical methods such as electroporation or microneedle injections of plasmid, especially
for transfection through the skin surface.
Among polymeric gene expression systems, biodegradable NPs offer certain
advantages for gene delivery. These can be formed by encapsulating DNA into
polymers or by complexing DNA to the surface of preformed NPs (Fig. 1). Plasmid
DNA can be encapsulated in polymeric NPs (i) alone as naked plasmid DNA (2,3),
(ii) condensed with some cationic polymers (4), or (iii) in noncondensed form with
some protective excipients (5). Polynucleotide molecules can also be complexed
with positively charged entities (cationic surfactants or polysaccharides) grafted
on the surface of preformed polymeric NPs (6). Thus, there is a great degree of flexibility in developing biodegradable NPs as gene expression vectors.
POLY(LACTIC ACID)/POLY(D,L-LACTIDE-CO-GLYCOLIDE)-BASED NPS
Poly(lactic acid) (PLA) and poly(d,l-lactide-co-glycolide) (PLGA) are the biodegradable and biocompatible polymers used for formulating NPs and are approved
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P O L Y M E R IC N A N O P A R T IC L E S
FOR
G E N E D E L IV E R Y
CONDENSE DNA
WITH POLYMERS
(POLYPLEXES)
POLY ETHYLENIMINE
POLY-L-LYSINE
POLY AMIDOAMINE
POLY β AMINO ESTERS
CATIONIC DENDRIMERS
CHITOSAN
ENCAPSULATE DNA
IN
POLYMERIC MAT RIX OR RESERVOIR
POLY LACTIC ACID
POLY LACTIC-CO-GLYCOLIC ACID
POLY β AMINO ESTERS
CHITOSAN
COMPLEX DNA
TO THE SURFACE OF
POLYMERIC NANOPARTICLES
POLY ALKYLCYANOACRYLATE
POLY LACTIC ACID
POLY LACTIC-CO-GLYCOLIC ACID
CHITOSAN
(Grafted with
cationic polysaccharides or surfactants)
- -- - +
DNA
+
+
+
+
+
---
+
- -+
- -+
- -
+
+
+
+
Polymer
-
+
+
+
+
+
+
Polyplex
Matrix
Reservoir
+
+
+-
-
+ Cationic surfactant
- -
FIGURE 1 Schematic representation showing different types of polymeric nanoparticles for gene
delivery.
for human use by the U.S. Food and Drug Administration. PLA/PLGA NPs with
entrapped plasmid DNA are of special interest for gene delivery due to their nonviral and nonimmunogenic nature. Plasmid DNA is entrapped into the polymeric
matrix, which not only protects DNA from nucleases, but also allows a control over
the DNA-release kinetics from NPs. The main advantage of these NPs is the slow
release of DNA from the NPs which facilitates sustained levels of gene expression.
Moreover, the duration and levels of gene expression can be easily modulated by
altering formulation parameters such as DNA:polymer ratio, or polymer molecular
weight, and composition.
PLGA NPs are generally formulated using “water-in-oil-in-water” doubleemulsion solvent evaporation technique, using polyvinyl alcohol (PVA) as an emulsifier (2). In brief, an aqueous solution of plasmid DNA is emulsified into the
polymer solution in organic solvents (usually chloroform or methylene chloride)
either by high-speed homogenization or sonication. This water-in-oil emulsion is
then mixed into a second aqueous phase (usually an aqueous solution of surfactant/
emulsifier, typically PVA is used). The double water-in-oil-in-water emulsion is then
formed using sonication or homogenization. The organic solvent is then allowed to
evaporate by stirring the emulsion at room temperature for approximately 18 to 20
hours. This results in the formation of polymeric NPs with entrapped plasmid DNA
which are recovered by ultracentrifugation, washed with distilled water to remove
PVA and unentrapped DNA. The NPs are suspended in water and lyophilized to
form a dry powder. NPs prepared using this procedure are usually around 100 nm
in diameter and have a negative zeta potential (7). DNA loading in these NPs has
Nanoparticles for Gene Delivery
293
been reported around 0.5% to 2.5% (w/w) (3). The efficiency of DNA encapsulation
using this process depends on the polymer composition, molecular weight of the
polymer, and the nature and concentration of emulsifier used in the process (2).
There are some concerns regarding the stability of DNA during the encapsulation
process, due to the use of high-shear forces (generated using high-speed homogenization or sonication) and exposure of DNA to the organic solvents (8,9). These conditions may result in the transformation of DNA from its supercoiled form to the
open-circular or linear forms. Preservation of the structural integrity of DNA is
important, whereas the supercoiled form has the highest bioactivity and the linear
form is least active (10). Studies have demonstrated that there is only a minimal loss
of activity of DNA following nanoencapsulation as the DNA is mostly present in the
supercoiled or open-circular forms in the NPs (11).
A cryopreparation method for the microencapsulation of plasmid DNA has
been described, to reduce the loss of the supercoiled form of DNA during formulation and to improve the encapsulation of DNA (12). This method involves freezing
the aqueous phase of the primary emulsion containing the plasmid DNA before
subjecting it to homogenization. This has been shown to preserve the content of the
supercoiled form of DNA, as DNA frozen in the aqueous phase is exposed to minimum shear stress. Addition of saccharides as cryoprotectants in the primary emulsion can also prevent the structural loss of DNA. Saccharides prevent crystal
formation from the buffer salts (present in the primary emulsion) during lyophilization and thus consequently prevent nicking of DNA by such salt crystals.
Oster et al. (13) have described a gentle solvent displacement method, without
the use of high-speed homogenization for the encapsulation of DNA. A new class of
biodegradable polymers consisting of amine-modified PVA backbone grafted with
PLGA side chains has been used in this procedure. The tertiary amino groups in the
polymer backbone interact with DNA by electrostatic interactions and facilitate
NP formation due to their amphiphilic character. These NPs exhibited positive
zeta potential and high transfection efficiencies in cell culture comparable to
polyethylenimine–DNA complexes. The high transfection efficiency may be due to
the rapid polymer degradation rates and the preservation of DNA integrity and
bioactivity during the NP formulation procedure.
Owing to their small size, PLGA NPs are taken up by cells (mainly by endocytosis) and facilitate intracellular delivery of plasmid DNA. As NPs come in direct
contact with the cell membranes, the surface properties of NPs are critical in
determining their intracellular fate and can potentially influence the gene transfection efficiencies (Fig. 2) (2). These include the surface-associated PVA in NPs, hydrophilicity, and the surface charge (zeta potential) of NPs. It has been shown that for
NPs formulated using the double-emulsion solvent evaporation, a fraction of PVA
used in the formulation remains associated with the NP surface, and cannot be
removed even by multiple washings (14). This residual PVA on NP surface can alter
its physical properties and affect the cellular uptake of NPs. NPs with lower amounts
of surface-associated PVA show about threefold higher cellular uptake in vascular
smooth muscle cells than the NPs with higher residual PVA (15). This could be due
to shielding of the surface charge reversal of NPs by the presence of higher amount
of surface-associated PVA, which could affect the endosomal escape of NPs. Further,
the amount of PVA associated with the NP surface depends on its concentration
used as an emulsifier during formulation, its molecular weight, and degree
of hydroxylation (2). Cellular internalization of NPs also depends on their particle
size and thus has been shown to affect the gene transfection. The smaller size
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Vasir and Labhasetwar
NPs
Protein
Cellular uptake
PE
RE
Gene expression
• Particle size
mRNA
• Surface properties
DN A
Nucleus
• Nuclear transport of
DNA
•Structural integrity of
DNA released from NPs
Intra-cellular release of DNA
• Endosomal escape
• Polymer composition
• DNA:polymer ratio
• DNA loading in NPs
• Inclusion of excipients
FIGURE 2 (See color insert.) Formulation factors influencing nanoparticle-mediated gene expression. Abbreviations: NPs, nanoparticles; PE, primary endosomes; RE, recycling endosomes.
(less than 100 nm) of NPs showed 27-fold higher gene transfection than the larger
size (more than 100 nm) NPs (7). Thus, the smaller size with a uniform particle size
distribution is expected to increase the gene transfection efficiency of NPs.
Other important formulation parameters which influence the gene transfection ability of NPs include the molecular weight of polymer and its composition
(lactide to glycolide ratio) (2). NPs formulated with higher-molecular-weight polymer showed enhanced gene transfection. This may be attributed to the relatively
higher DNA loading and its release from NPs prepared with high-molecular-weight
polymer (Fig. 3). Higher viscosity and better emulsifying properties of the polymer
solution facilitate higher loading of DNA in NPs and also lead to lower particle size
of NPs. Polymer composition can affect the hydrophobicity of the polymer and thus
can affect the DNA loading and release of DNA from the NPs. NPs prepared using
more hydrophobic polymers (polylactides) demonstrated lower transfection than
those formulated using copolymers of polylactide and glycolide (Fig. 4). The slow
rate of release of DNA from the hydrophobic polymeric matrix may be responsible
for the lower levels of gene transfection.
Although the levels of gene expression with NPs are lower than that achieved
with lipid-based gene delivery, they are sustained for a prolonged period of time.
Further NP-mediated gene transfection is not affected by the presence of serum in
the cell culture media and thus PLGA NPs constitute a potential gene delivery
vector for in vivo gene delivery. Prabha and Labhasetwar (16) have shown slow
intracellular release of plasmid DNA from the PLGA NPs, which results in sustained retention of DNA inside the cells (Fig. 5). NPs loaded with wt-p53 gene demonstrated higher mRNA levels for p53 as compared to that with a liposomal
formulation at five days after transfection of MDA-MB-435S breast cancer cells. The
sustained p53 expression levels resulted in greater and sustained inhibition of cell
proliferation in vitro as compared to that with liposomal formulation or plasmid
Nanoparticles for Gene Delivery
295
FIGURE 3 Effect of molecular weight
of PLGA on (A) in vitro release of DNA
from NPs and transfection of NPs in
(B) MCF-7 cells. Cells (35,000 per
well in 24-well plate) were incubated
with NPs (444 µg/mL/well) for one day
after which the medium in the wells
was replaced with fresh medium (without NPs). Medium was changed on
alternate days thereafter. NPs showed
sustained gene transfection in MCF-7
cell line. Data as mean ± S.E.M.,
n = 6. Abbreviation: NPs, nanoparticles. Source: From Ref. 2.
DNA alone. Cohen et al. (3) have shown that despite the lower transfection levels
observed in vitro with NPs as compared to liposomal formulations, the in vivo gene
transfection with NPs was one to two orders of magnitude greater than that with
liposomes at seven days after an intramuscular injection in mice. Their studies demonstrated gene expression sustaining over 28 days in vivo with a single dose of
intramuscular injection of NPs. Such sustained gene expression is advantageous
especially if the half-life of the expressed protein is very short and/or a chronic gene
delivery is required for better therapeutic efficacy.
Although these NPs have been shown to release plasmid DNA at slower rates
inside the cell, it is important that the DNA released as a result of polymer degradation retains its bioactivity. There are reports showing that the degradation of PLGA/
PLA polymers into lactic and glycolic acids can produce highly acidic microenvironments in the NPs, which can compromise the stability and activity of DNA
released from NPs (17). Additional excipients such as polyethylene oxide (PEO) can
be coencapsulated with plasmid DNA in polymeric NPs in order to prevent the
generation of extremely acidic microenvironments inside NPs on polymer
degradation (18). Polymer blends of PLGA and polyoxyethylene derivatives have
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Vasir and Labhasetwar
FIGURE 4 Effect of polymer composition on (A) in vitro DNA release
from NPs and (B) transfection of NPs
in MCF-7 cells. Cells (35,000/well in
24-well plate) were incubated with
NPs (444 µg/mL/well) for one day,
and then the medium was replaced
with fresh medium (without NPs).
Medium was changed on every alternate day thereafter, and transfection
levels were determined at one, three,
five, and seven days post-transfection in MCF-7 cell line. This figure
represents lactide:glycolide ratio.
Data shown as mean ± S.E.M., n = 6.
Abbreviation: PLA, Poly(lactic acid).
Source: From Ref. 2.
1.2
*
Nanoparticles
DNA
Fluorescence intensity
1
0.8
*
0.6
0.4
0.2
0
0.08
1
3
Time (Day)
5
7
FIGURE 5 Quantitative determination of intracellular DNA levels
cells transfected with YOYOlabeled DNA-loaded nanoparticles demonstrated sustained and
increased intracellular DNA levels
as opposed to transient DNA
levels in the cells transfected with
naked DNA. Data are represented
as the mean ± the standard error
of the mean (n = 6; P < 0.001 for
points marked with asterisks).
Source: From Ref. 16.
Nanoparticles for Gene Delivery
297
been used to prepare NPs entrapping plasmid DNA, using a modified emulsificationsolvent diffusion technique (5). PLGA and poloxamer or poloxamine were mixed in
different ratios and dissolved in methylene chloride. An aqueous solution of plasmid
DNA can then be emulsified into this organic phase containing polymers using vortexing. The emulsion is then poured into a polar phase under moderate magnetic
stirring, to allow immediate precipitation of the polymer. Thus, this modified technique avoids the use of high shear forces during nanoencapsulation of DNA, which
can cause structural damages to the DNA. Moreover, the use of poloxamers and
poloxamine facilitated DNA encapsulation and resulted in an increase in the encapsulation of DNA within the NPs. The nature and HLB value (hydrophilic/lipophilic
balance) of the surfactant (poloxamer/poloxamine) can possibly affect the interaction of DNA with the hydrophobic polymers. Higher encapsulation of DNA with
polymer blends made with surfactants of intermediate HLB values can be attributed to the improved compatibility of hydrophilic plasmid DNA with the hydrophobic PLGA polymer. The rate of release of plasmid DNA from the NPs also
depends on the HLB value of the polyoxyethylene derivative used in the polymer
blend. NPs prepared using hydrophilic poloxamers (Tetronic 908, HLB = 30.0)
showed a continuous and fast release of DNA within one week, whereas the ones
prepared with a more hydrophobic surfactant (Tetronic 904, HLB = 14.5) showed
slow release of DNA over a period of two weeks (Fig. 6).
Attempts have been made to improve the encapsulation of DNA in biodegradable NPs. Condensing plasmid DNA prior to encapsulation can increase the
encapsulation as compared to encapsulation of uncondensed DNA. Plasmid DNA
condensed with polycations such as poly-l-lysine (PLL) or with cationic dendrons
has also been encapsulated in the PLGA NPs. This allows for the protection of DNA
inside the polymeric matrix and controlled delivery of plasmid DNA. Ribeiro et al.
(4) have reported the encapsulation of plasmid DNA condensed with cationic
PLL-based lipidic dendrons (dendriplexes) into PLGA NPs by double-emulsion
method. The size of PLGA-dendriplex particles was a function of the molar charge
ratios at which DNA was condensed with the dendriplexes. Plasmid DNA could be
FIGURE 6 Release profiles of plasmid DNA from PLGA:Tetronic 908 (squares) and PLGA: Tetronic
904 (triangles) blend NPs and from the control PLGA (stars) NPs (mean ± S.D., n = 3). Abbreviations:
PLGA, poly(D,L-lactide-co-glycolide); NPs, nanoparticles. Source: From Ref. 5.
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Vasir and Labhasetwar
fully condensed with cationic dendrons at high (10:1) molar charge ratios, resulting
in dendriplexes with a positive zeta potential and a DNA loading of 15.0% (w/w).
PLGA NPs can be modified to generate cationic surfaces, to which negatively
charged DNA molecules can be condensed. An emulsion–diffusion–evaporation
technique using PVA–chitosan blend as a stabilizer has been described to produce
cationically modified PLGA NPs (19). PVA helps in the stabilization of the particles,
whereas chitosan owing to its positive charge provides a cationic surface to the NPs.
Plasmid DNA can be complexed to the surface of these preformed cationic NPs by
means of electrostatic interactions, thus avoiding the involvement of DNA during
the NP formulation and hence preserving DNA integrity. Blends of the polymers
PLA/PLGA with polyethylenimine (PEI) have also been used to prepare positively
charged NPs by a diafiltration method (20). The use of diafiltration technique
avoids the use of any surfactant in the NP formulation as the hydrophilic nature of
PEI effectively reduces the interfacial energy between the hydrophobic NP surfaces
and the aqueous media. Particle size and the zeta potential of these NPs can be
controlled by the amount of PEI used in the polymer blends. Imine groups of PEI
can be used to complex DNA to the surface of cationic NPs. The NPs produced
using this method showed high adsorption capacity for plasmid DNA and a high
dispersive stability.
Further, the surface of polymeric NPs based on PLGA/PLA can be functionalized to improve their biodistribution and also to conjugate targeting ligands which
can direct NPs to specific cells/tissues where gene delivery is desired. Surface modification of NPs is achieved either by adsorbing amphiphilic excipients onto preformed
NPs or by covalently linking excipients to the core-forming polymer prior to NP
formulation. Biodegradable NPs have been prepared using PLL-graft-polysaccharide
copolymers and poly(d,l-lactic acid) by using a solvent evaporation method or the
diafiltration method (6). NPs prepared using these copolymers had bifunctional
surfaces with positively charged amino groups of PLL and the polysaccharide moieties. This increased the adsorption capacity of NPs for polynucleotides and allowed
introduction of ligand (carbohydrate) moieties on the NP surface for ligandmediated recognition of specific receptors. The formulation resulted in NPs as small
as 60 nm in diameter, and showed excellent dispersive stability in phosphatebuffered saline. A preferential distribution of dextran moieties over PLL was
observed in the outer surface of NPs using a PLL-graft-dextran copolymer. Presence
of dextran chains on NP surface can be potentially useful to prevent nonspecific interactions with serum proteins and also for receptor-mediated targeting to specific cells.
Subsequent to intravenous injection of the NPs, the major problem is to avoid
their opsonization by plasma proteins and subsequent uptake by the phagocytic
cells. PEO/PEG has been used to coat the polymeric NPs to provide a protective
hydrophilic sheath, which prevents the rapid opsonization of the otherwise hydrophobic NPs by reticuloendothelial system (RES) and thus prolong the circulation
time of NPs in blood stream (21). The hydrophobic part of PEO/PEG polymers can
adsorb to NP surface, whereas the hydrophilic chains protrude towards the aqueous
medium. PEG coats on NP surface also provide an attractive opportunity to chemically conjugate active-targeting ligands to the NP surface (22,23). These coatings can
modify the biodistribution of NPs when injected into the systemic circulation.
However, it has been argued that some of these polymers can be easily displaced by
serum proteins, which can lead to aggregation of NPs (24). Thus, alternative
approaches of synthesizing copolymers of PLA/PLGA with PEG (25,26) and coencapsulation of PEG with plasmid DNA inside PLGA NPs have been tried (27).
Nanoparticles for Gene Delivery
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CHITOSAN NANOPARTICLES
Chitosan is a linear cationic polysaccharide, obtained by partial alkaline deacetylation of chitin, a polymer found in the exoskeleton of crustaceans. Chitosans are
biodegradable and biocompatible polymers and thus can be potentially safe and
nontoxic carriers for gene delivery. Moreover, derivatives of chitosan can be synthesized, with relative ease, to hydrophobically modify chitosan or to conjugate different
chemical entities for targeting chitosan NPs to specific tissues or organs.
NPs for gene delivery can be prepared by (i) complexing plasmid DNA with
chitosan (28) or (ii) by encapsulating DNA inside chitosan NPs (29). The high density of amino groups present in the glucosamine backbone of chitosan can be protonated and thus offers the opportunity to complex chitosans with negatively
charged DNA molecules, owing to the electrostatic interactions.
Self-assembling chitosan/DNA complexes, in the size range of 150 to 500 nm,
were first prepared by mixing a solution of chitosan with plasmid DNA (30). The
particle sizes of the complexes depend on the molecular weight of chitosan used,
but not on the buffer compositions. Chitosan, hydrophobically modified using
deoxycholic acid, can form spherical self-aggregates with a mean diameter of 160 nm
in aqueous media (28). Charge complexes formed between these self-aggregates of
modified chitosan and plasmid DNA have been shown to transfect COS-1 cells in
culture. The transfection efficiency of chitosan self-aggregate/DNA complexes was
reported to be more than that achieved with plasmid DNA alone, but lower than
liposomal formulations.
DNA can be encapsulated in chitosan NPs, prepared by a complex coacervation method, yielding particle sizes in the range of 200 to 500 nm. Mao et al. (29)
have studied the influence of various parameters of NP preparation on the size of
chitosan–DNA NPs. The size of the particles was optimized to 100 to 250 nm using
an N/P ratio of three to eight. The zeta potential of these particles was +12 to +18 mV
at a pH lower than 6.0 and nearly zero at pH 7.2. Chitosan NPs could partially protect DNA from nucleases. The transfection efficiency was found to be cell-dependent
and lower than that achieved with Lipofectamine–DNA complexes in HEK 293
cells. However, the presence of serum in the cell culture medium did not interfere
with the transfection using chitosan NPs. Chitosan–DNA NPs are also considered a
very appealing carrier choice for oral gene-delivery strategies, owing to the mucoadhesive properties of chitosan. Orally administered chitosan–DNA NPs can adhere
to the gastrointestinal epithelium and transfect epithelial or immune cells present in
the gut-associated lymphoid tissue (31).
POLY(β-AMINO ESTER)-BASED NANOPARTICLES
Poly(β-amino esters) constitute a relatively new class of synthetic biodegradable
cationic polymers with tertiary amines in their backbone. Unlike other cationic
polymers, these are hydrolytically degradable polymers and are generally less cytotoxic than the polycations such as PEI (32). Using MTT assay, these polymers and
their degradation products have been shown to be nontoxic relative to PEI in NIH
3T3 cell line.
These polymers have been shown to possess pH-dependent solubility, and
thus are suitable for the preparation of DNA NPs which can trigger the polymer
degradation and release of encapsulated DNA in the acidic pH in endosomal vesicles
inside cells (33). Owing to their polycationic nature, these polymers can condense
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Vasir and Labhasetwar
DNA into complexes of the order of 50 to 200 nm and thus can be used for gene
transfection.
POLY(ALKYLCYANOACRYLATE) NANOPARTICLES
Poly(alkylcyanoacrylate) (PACA) NPs were first prepared by Couvreur et al. in 1979
(34). PACA NPs have been used for intracellular delivery of various oligonucleotide
sequences. PACA NPs can be coated with cationic surfactants (cetyltrimethylammonium bromide) or polymers (DEAE-dextran) which can be used for complexing
oligonucleotides via electrostatic interactions (35). Alternatively, oligonucleotides
can be associated with the PACA NPs by covalently linking to a hydrophobic
molecule (cholesterol) which can then be anchored at the NP surface (36). In earlier
studies, it has been shown that the PACA NPs are taken up by the RES system of the
body and thus can be used for targeting DNA delivery to the RES organs such as
liver, spleen, lungs, and bone marrow (37).
MAGNETIC NANOPARTICLES
Magnetic NPs usually incorporate magnetic-responsive materials such as magnetite,
iron, nickel, cobalt, and so on. In contrast to the conventional polymeric NPs,
magnetic NPs can be directed and preferentially localized in the diseased tissue
upon application of an external localized magnetic field.
The magnetic NPs can be modified with different chemical compounds to
facilitate a change in the surface charge or particle size of the NPs. Magnetic NPs for
gene delivery involve associating naked plasmid DNA or DNA vectors (such
as DNA complexed with polycations or packaged in viral vectors) to the surface of
the NPs. This technique of gene transfection is commonly referred to as “magnetofection.” The magnetic field can improve the vector concentration near the cells to
be transfected and thus has been shown to reduce the vector dose and incubation
time with vector required to achieve high transfection efficiency in vitro (38).
The surface of superparamagnetic iron-oxide NPs coated with PEI has been used to
associate PEI–DNA complexes (38). The basic principle used for the formulation of
these magnetic NPs is that of salt-induced aggregation of colloidal particles.
PEI-coated magnetic NPs, PEI, and plasmid DNA are mixed in salt-containing solutions to form the magnetic vectors for gene transfection. PEI–DNA complexes
(polyplexes) associated with superparamagnetic NPs can be guided to accumulate
selectively in the target tissue, by the use of an external magnetic field focused at the
tissue. Magnetic-field-guided local transfection in the gastrointestinal tract and in
blood vessels has also been reported (38).
Another report of using magnetic NPs for gene delivery involves association
of magnetic (maghemite) NPs with the DNA vector hemagglutinating virus of
Japan-envelope (HVJ-E) (39). This can potentially enhance the gene transfection
efficiency of the vector by promoting its cell association, using magnetic force.
The magnetic NPs were coated either with protamine sulfate or heparin to produce
a cationically charged surface of the NPs. The DNA-carrying vector—HVJ-E—was
associated with the modified NPs and used for gene transfection. These magnetic
NPs promote gene transfection, mediated by cell fusion of the HVJ-E vectors.
Cationically charged magnetic NPs can be speculated to promote greater association of the vector with the cells, which then carry DNA into the cells under the influence of magnetic force, thus enhancing the transfection efficiency of the vectors.
Nanoparticles for Gene Delivery
301
Thus, the use of magnetic NPs in conjunction with DNA alone or DNA polyplexes
is an emerging gene transfection technique, which can effectively increase the
gene transfection efficiency of these vectors. However, it is important that such
magnetic NPs for gene transfection be water-based, biocompatible, nontoxic, and
nonimmunogenic.
SUMMARY
Biodegradable NPs constitute safer and versatile gene-delivery systems that allow
ligand conjugation for targeted gene therapy as well as modulation of the DNA
release to control the level and duration of gene expression. Although we have
achieved some success in overcoming the cellular barrier of DNA delivery, the
nuclear membrane still stands as a rate-limiting factor for efficient gene expression
using nonviral vectors (40). Incorporation of cationic peptides such as nuclear localizing signaling peptides, protein transduction domain from the transactivator of
transcription, and protamine have been shown to enable nuclear transport of DNA,
and hence conjugating these to a carrier system may prove to be an effective strategy
to enhance gene expression of nonviral systems (41). For in vivo applications, it is
important to restrict the gene transfection to the treated (target) tissue, and thus,
targeting of NPs can potentially increase the efficacy and reduce the toxicity of gene
therapy. Approaches such as decorating the surface of NPs with cell-specific ligands
and use of magnetic force to target the NPs to specific tissues or organs for targeted
gene delivery, though in early stages of development, constitute promising efforts at
improving gene delivery. To successfully develop such formulation approaches
for targeted gene delivery, it is essential to add to our basic understanding of the
cellular barriers and mechanisms of intracellular sorting of NPs, and ultimately
define rational formulation strategies to overcome these barriers (42).
ACKNOWLEDGMENTS
Grant support from the National Institutes of Health (R01 EB003975) and a
predoctoral fellowship to J.K.V. from the American Heart Association, Heartland
Affiliate (Award #0515489Z). Author (V.L.) acknowledges his former laboratory
members, Dr. Swayam Prabha and Dr. Sanjeeb K. Sahoo, whose data are cited in
this chapter.
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19
Gastrointestinal Applications of
Nanoparticulate Drug-Delivery Systems
Maria Rosa Gasco
Nanovector srl, Torino, Italy
EXCITING REASONS FOR STUDYING THE APPLICATION OF
NANOPARTICULATE SYSTEMS TO THE GASTROINTESTINAL TRACT
The uptake of nanoparticulate systems (NPSs) through the gastrointestinal tract
(GIT) is today a well-known and accepted phenomenon, and excellent reviews of
the intestinal uptake of particles have been published (1–5). NPS uptake from the
gut can provide an additional drug administration route; each system has its own
pharmacokinetic parameters and specific drug-carrying ability. The bioactive molecule is transported into the GIT by carriers whose chemicophysical characteristics
must be taken into account, although the chemicophysical and pharmacological
characteristics of the drug remain intact. This chapter will concentrate particularly
on the translocation of NPSs via the lymphatic system, and briefly consider the
hitherto less widely studied colonic targeting of NPSs.
Lymphatic Targeting
Peyer’s patches are the most important structural units of the gut associated with
lymphoid tissue (GALT); they are characterized by M-cells that overlie the lymphoid tissue and are specialized for endocytosis and transport into intraepithelial
spaces and adjacent lymphoid tissue. Nanoparticulates bind to the apical membrane
of M-cells, after which they are rapidly internalized and “shuttled” to the lymphocytes (3,5). The lymphatic absorption of a drug via the GALT has an advantage over
the portal blood route, because it avoids presystemic metabolism by the liver
(hepatic first-pass effect). Florence (6) has suggested a simplified diagram of
pre- and postabsorption processes (Fig. 1).
The requisites enabling NPS to be absorbed via the GALT not only concern the
loaded drug, but also and more specifically the carrier’s physical characteristics,
such as size, shape, specific surface, surface charge, chemical stability of both NPS
and loaded drug, potential interactions with gut contents, transit time through the
GIT, transport through the mucosa, adhesion to epithelial surfaces, and aggregation
of the particulates in contact with the fluid content of the gut. The transit and translocation of NPSs depend to a considerable extent on their mean diameter, surface
charge, and release characteristics (2,3). Some phenomena, such as aggregation, adsorption, and adhesion, can alter the zeta potential, hydrophilicity, and size of the NPS.
From the pharmaceutical standpoint, the measure of any feasible application
of a drug-delivery system is its efficacy, that is, its ability to exert a pharmacological
effect. Indeed, besides the different properties of NPS listed above, a critical aspect
is the loading capacity of the nanoparticulates. The higher the loading, the higher is
the bioavailability per particle absorbed (2). Another important issue is biocompatibility and biodegradability of the NPS components.
305
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Gasco
FIGURE 1 A simplified schematic of pre-absorption and post-absorption processes in nanoparticledependent drug delivery to gastrointestinal sited following oral administration, highlighting the variety
of processes involved in the journey of a nanoparticle from delivery to target. Source: From Ref. 6.
Lymph targeting of NPSs can permit (2,3,5): (i) oral delivery of labile molecules
protected by the carrier; (ii) oral delivery of poorly soluble molecules following
their nanosolubilization in NPSs; (iii) improved bioavailability of drugs with poor
absorption characteristics, due both to the large specific surface of NPSs and to their
increased residence time; (iv) oral delivery of vaccine antigens to gut-associated
lymphoid tissue; (v) translocation of antineoplastic drugs for treatment of lymphomas; (vi) delivery of diagnostics for the lymphatic system; (vii) sustained/controlled
drug release, particularly important for toxic drugs (e.g., antineoplastic drugs);
(viii) reduction of drug-related GI mucosa irritation; and (ix) avoidance of the
hepatic first-pass effect.
Colonic Targeting
Lymphatic tissue in the colon, rectum, and appendix vermiform is usually not found
in discrete patches, but is diffusely present in aggregate masses at irregular intervals.
Despite the fact that the colon and rectum contain the most abundant lymphoid
intestinal tissue, there is a lack of research examining lymphatic absorption from
these sites (2). The presence of M-cells (7) suggests that NPSs may also be absorbed
in this region of the GIT. Colon-specific drug delivery systems can be used to
improve drug bioavailability at the colon, particularly in the case of proteins and
peptides, targeting them to the less proteolytically active colon (8). Nanoparticulates
targeted to the colon have been studied with the aim of treating some types of colon
disease, but also to investigate transport and release of some labile molecules such
as peptides and proteins as some early studies showed.
Colonic targeting of nanoparticulates can permit therapeutic intervention in
pathological processes of the gut, such as ulcerative colitis or Crohn’s disease (2),
and possible targeting of labile molecules (such as peptides and small proteins) to
the colon (8).
DEVELOPMENT OF NANOPARTICULATE SYSTEMS
FOR LYMPHATIC TARGETING
The following NPSs are considered.
Gastrointestinal Applications of Nanoparticulate Drug-Delivery Systems
307
Dendrimers
Although the first reports on dendrimers were published more than two decades
ago, their potential biological applications have only been explored in the past few
years (2,9,10).
Dendrimer chemistry has been used to prepare monodisperse nanoparticles
(diameter below 50 nm) in order to investigate the relation between nanoparticle
diameter and uptake from the GIT (3). The study examined both dendrimers with
lipidic external character and a series of cationic dendrimers, and investigation of
dendrimer absorption in rats through Peyer’s patches and enterocytes showed their
preferential uptake through Peyer’s patches.
5-Fluorouracil (11) can be entrapped in polyamidoamine dendrimers that
have been coated with phospholipids; in vivo studies in albino rats have shown
phospholipid-coated dendrimers to be more effective orally than free drug.
Lymphatic uptake also increased, indicating absorption of the formulation via the
lymphatic route.
Polymeric Nanoparticulates (Nanoparticles and/or Nanocapsules)
Polymeric nanoparticulates are particles of diameter below 1 µm, prepared from
natural or synthetic polymers. Natural polymers (i.e., proteins or polysaccharides)
have not been widely used for this purpose, as they may vary in purity and often
require cross-linking, which could denature the embedded drug (5). Synthetic
biodegradable polymers have received much more attention; the most widely used
polymers have been poly(lactic acid), poly(glycolic acid), their copolymers
poly(lactide-co-glycolide acid) (PLGA) (12), and polyalkylcyanoacrylates (PACA)
(13). The bioactives are well protected in these polymers, which offer the advantage
of delivering drugs in a sustained fashion, avoiding repeated administration.
Nanoparticles were first developed for the parenteral route; more recently, they
have also been studied as oral delivery vehicles. The major interest is in the lymphatic uptake of nanoparticles by Peyer’s patches in the GALT (12). Microparticles
remain in Peyer’s patches, while nanoparticles are disseminated systemically (5).
Clearly, a wide variety of drugs can be delivered via the oral route using polymeric
nanoparticulate carriers. Much of the research has focused on the absorption
enhancement of peptides (14), proteins, and vaccine antigens.
Mucoadhesive ability can be conferred to NPSs by coating their surface with
microadhesive polymers such as chitosan or Carbopol®; action is more effective and
prolonged, confirming the usefulness of mucoadhesive properties (15). Insulin
entrapped in PACA nanospheres dispersed in an oily phase containing a surfactant
produces prolonged hypoglycemia (16). The fate of poly(isobutylcyanoacrylate)
nanocapsules carrying insulin administered to rats was monitored by fluorescence
and transmission electron microscopy (TEM) and evidenced intestinal absorption
through the epithelial mucosa (17). A luminescent polymer was used as a visible
tracer to monitor the oral fate of PLGA microspheres containing insulin (18). The
therapeutic effects of another peptide, octreotide, can be improved and prolonged
on incorporation into PACA nanocapsules (19).
Biodegradable PLGA nanoparticles containing salmon calcitonin (sCT), complexed with amphiphilic molecules, afford high loading efficiency; absorption was
good on oral administration to rats (20). Positively charged nanoparticles incorporating cyclosporin A (21) have been prepared by the emulsification solvent diffusion
method; the bioavailability in beagle dogs increased 1.2-fold over that of Neoral®.
308
Gasco
The same peptide was encapsulated by nanoprecipitation within nonbiodegradable
polymers, and different formulations were prepared; oral absorption of the peptide
nanoparticulates in rabbits was compared to that of Neoral® capsule. The relative
bioavailability of cyclosporin A from various formulations ranged from 20% to 35%
that of Neoral® (22).
Heparin has no oral bioavailability, and is thus generally administered by the
parenteral route. Heparin-loaded polymeric nanoparticles, prepared with biodegradable poly-ε-caprolactone and PLGA and nonbiodegradable positively charged
polymers, used alone or in combination, were administered orally to rabbits (23);
with each formulation anti-Xa activity and activated partial thromboplastin time
(aPTT) were detected. In particular, anti-Xa activity was detected for a longer period
than when a heparin solution was administered intravenously.
A delivery system allowing oral administration of vaccines would be ideal.
An antigen is taken up into the Peyer’s patches mostly through the M-cells, which
appear to be the main site for uptake of antigens after oral administration, and it is
generally accepted that limited doses of antigen are sufficient for mucous immunization. Oral delivery of antigens may be considered the most convenient means of
producing an IgA antibody response (24). The use of micro- and nanoparticles for
the oral delivery of antigens can be efficient if these systems are able to protect the
antigen molecule. Oral administration of antigens incorporated in particulates
induces a stronger antigen-specific immune response than do antigens in the watersoluble form, because incorporation in NPSs protects them from the low pH of the
stomach and from proteolytic enzymes (25,26).
The immune response has been studied after subcutaneous, oral, and intranasal
administration to mice of PLGA nanospheres loaded with bovine serum albumin
(BSA), evaluating some factors affecting the immune response, such as size, surface
hydrophobicity, zeta potential, adjuvants, or excipients used in the formulations (27).
Liposomes
Oral delivery of liposomes would be the simplest and most convenient route for
drug administration (28,29), but it is important to examine their stability in the acid
pH of the stomach environment and in the gastrointestinal medium. Stability in the
physiological conditions affecting oral drug delivery has been studied to determine
which components have the best chance of surviving through the GIT (30).
Liposomes coated with polyethylene glycol containing recombinant human
epidermal growth factor were administered orally to rats and the area under the
concentration–time curve (AUC) was evaluated and compared to that of the
solution. It increased 1.7- and 2.5-fold for phosphatidylcholine and dipalmitoylphosphatidylcholine liposomes, respectively (31).
Liposomes containing an extract of natural marine lipids containing large
amounts of N-3-polyunsaturated fatty acids (PUFA) were prepared with the aim of
increasing PUFA bioavailability. Thoracic lymph duct-cannulated rats were fed
intragastrically with these liposomes, and absorption of fatty acids was higher than
with fish oil (32).
The effect of liposomes coated with chitosan or Carbopol® on rats’ intestinal
absorption of calcitonin has been studied, testing negatively and positively charged
liposomes. The overall pharmacological efficacy of coated liposomes was more than
double that of noncoated liposomes (33). Absorption of sCT from the intestine is
poor because of its high molecular weight and rapid degradation by enzymes in the
Gastrointestinal Applications of Nanoparticulate Drug-Delivery Systems
309
GIT. Double liposomes (i.e., liposomes containing smaller liposomes) with different
charges were investigated as carriers of sCT, administered to rats, and compared to
the solution. The highest hypocalcemic effect was obtained with cationic charged
liposomes (34).
Self-(Micro)Emulsifying Drug-Delivery Systems
Self-(micro)emulsifying Drug-Delivery Systems (S(M)EDDSs) are isotropic mixtures
of oils, surfactants, solvents, and cosolvents/surfactants; they are used solely for the
purpose of improving oral absorption of highly lipophilic drugs. The principal
characteristic of these systems is their ability to form fine oil-in-water emulsion or
microemulsion upon mild stirring, following dilution by the aqueous phase (35,36).
A supersaturable (S-SEDDS) formulation of paclitaxel, developed using
hydroxypropylcellulose as precipitation inhibitor (37), in a pharmacokinetic study
on rats achieved fivefold oral bioavailability versus an orally dosed Taxol®
formulation.
An S(M)EDDS of paclitaxel was administered to rats with and without the
use of P-glycoprotein inhibitors (38). Compared to commercial Taxol®, the oral
bioavailability of paclitaxel S(M)EDDS increased at the various doses. The bioavailability was significantly improved for paclitaxel S(M)EDDS when cyclosporin A, as
inhibitor of P-glycoprotein, was coadministered. An S(M)EDDS system loaded with
simvastatin, orally administered to beagle dogs, increased bioavailability 1.5-fold
versus the conventional oral tablet (39).
Solid–Lipid Nanoparticles
Solid–lipid nanoparticles (SLNs) are in the submicron size range and usually consist
of biocompatible and biodegradable materials, such as triglycerides and fatty acids;
the production and properties of SLNs have been reviewed (40–42). Various preparation methods are possible: one depends on the application of mechanical highpressure homogenization to a melted lipid dispersed in an aqueous solution of
surfactants (40). When camptothecin-loaded SLNs, produced by high-pressure
homogenization, were administered perorally to rats, the concentration–time curves
in plasma and in organ tissues showed enhanced availability of the drug compared
to solution: AUC and mean residence time (MRT) increased significantly (43).
Piribedil SLNs, administered orally to rabbits, increased drug bioavailability more
than twofold compared to pure piribedil (44). SLNs loaded with all-trans retinoic
acid, as a poorly soluble model drug, can be prepared by high-pressure homogenization; administered orally to rats, they significantly enhanced trans retinoic acid
absorption (45). The bioavailability of clozapine formulated in SLNs prepared by the
hot homogenization method was determined after intravenous (IV) and intraduodenal administration and compared to those of a clozapine suspension. The area under
the curve (AUC) increased by about three times in the case of the IV route, and up to
4.5 times for the intraduodenal route (46). Clear solutions of cyclosporin A, a solid
triglyceride, a water-miscible organic solvent, and a mixture of surfactants and
emulsifiers were dispersed by mixing in water to produce a suspension of nanoparticles of sizes between 25 and 400 nm depending on the formulations. The oral
bioavailability was determined on humans; the best results were similar to the
Neoral® reference formulation and were obtained for the formulations forming
nanoparticles below 60 nm (47). Other preparation methods for SLNs are the solvent
emulsification technique (48) and the solvent diffusion method (49).
310
Gasco
SLNs can also be produced by dispersing warm microemulsions in cold water
(50,51). SLNs are the solidified droplets of microemulsions with a mean diameter of
80 to 200 nm; they can incorporate both hydrophilic and lipophilic drugs. Many
drugs have been incorporated in SLNs, and different administration routes have
been studied in laboratory animals. Stealth SLNs have been prepared for the purpose
of avoiding reticuloendothelial system recognition; after IV administration, SLNs
and stealth SLNs are able to cross the blood–brain barrier, increasing the MRT (to a
greater extent with stealth than with nonstealth SLNs) of the loaded drug compared
to solution (52). SLNs are taken up within a few minutes by neoplastic (53,54) and
nonneoplastic cell lines (55).
The gastrointestinal uptake and transport in the lymphatic circulation of drugunloaded SLNs was initially studied; labeled and unlabeled SLNs were administered duodenally to rats, at two different amounts in equal volumes. After a few
minutes, TEM evidenced SLNs in the lymph; the size of the particles in the lymph
(130–140 nm) was practically unchanged after administration. Labeled SLNs were
used to evaluate lymphatic uptake and the radioactivity data confirmed transport
of SLNs in lymph and blood. SLN concentration versus time increased out of
proportion to the increased amount administered, being very much higher with the
higher concentration of SLNs (56).
Successively, tobramycin was incorporated in SLNs as a model drug; it was
chosen because it is not absorbed at the gastrointestinal level and is therefore
still administered by the parenteral route. Tobramycin-loaded SLNs (Tobra-SLN)
were administered to rats into the duodenum and compared to Tobra-SLN
and tobramycin solution, both administered IV (57). The AUC of IV Tobra-SLN
was 1710 min µg mL−1, whereas that of Tobra-solution administered IV was
352 min µg mL−1. The AUC of Tobra-SLNs administered duodenally was over 100
times that of the solution and over 20 times that of Tobra-SLN, both administered IV.
Pharmacokinetic parameters showed a marked difference between Tobra-SLNs
administered duodenally and IV, in the former case providing a sufficiently high
level of the drug even after 24 hours. Most of the difference between the two administration routes is due to the transmucosal transport of SLNs to the lymph rather than
to the blood. Tobra-SLN administered duodenally act as a sustained-release system.
SLNs containing three different percentages (1.25%, 2.5%, and 5%) of tobramycin were prepared, and the same dose of Tobra of each of the three types
was administered intraduodenally to rats (58). Figure 2 shows the tobramycin
plasma concentrations versus time for the three types of SLNs. The pharmacokinetic
FIGURE 2 Tobramycin plasma
concentrations versus time ±
standard deviation after duodenal
administration of the three types
of Tobra-SLN. Abbreviation: SLN,
solid–lipid nanoparticle. Source:
From Ref. 58.
Formulation
Tobra-SLN 1.25
Tobra-SLN 2.50
Tobra-SLN 5.00
a“Apparent”
Cmax
(mg/L)
31.5b,e
± 1.9
28.5 ± 0.4
36.3h ± 1.2
MRT
(min)
2964c,f
± 858
1544 ± 304
439g ± 138
T1/2β
(min)
1901c,f
± 310
998 ± 21
291g ± 16
AUC
(min.mg/L)
74880d,f
± 2047
37469 ± 1516
13817h ± 707
Vd (L)
Cl (L/h)
0.094 ± 0.010
0.096 ± 0.009
0.073 ± 0.020
0.002d,f ± 0.0001
0.004 ± 0.0002
0.010i ± 0.0003
pharmacokinetic parameters are so named because they refer to the drug incorporated into SLN and not to the free drug; data expressed as mean ± SD; MRT, mean
residence time; T1/2β, elimination half-life; AUC, area under the curve of plasma concentration versus time, Cl, total body clearance; Vd, volume of distribution; statistical significances are;
SLN 1.25 versus SLN 2.50
bp < 0.05.
cp < 0.01.
dp < 0.001;
SLN 1.25 versus SLN 5.00
ep < 0.01.
fp < 0.001;
SLN 2.5 versus SLN 5.00
gp < 0.05.
hp < 0.01.
ip < 0.001.
Abbreviations: AUC, area under the curve; SLN, solid–lipid nanoparticle.
Source: From Ref. 57.
Gastrointestinal Applications of Nanoparticulate Drug-Delivery Systems
TABLE 1 Apparent Pharmacokinetic Parametersa
311
312
Gasco
parameters varied considerably with the percentage of Tobra (Table 1). This finding
is probably due to the differences among the three types of SLNs, in particular, the
number of SLNs administered, average diameter, total surface area, and drug content in each nanoparticle. The highest percentage of Tobra in SLNs corresponded to
the highest release rate, whereas the lowest percentage produced the most prolonged
release. Sustained drug release by the oral route could thus be obtained by appropriately mixing SLNs loaded with different percentages of the drug. Tobramycin
was still determined in lymph mesenteric nodes 21 hours after duodenal administration of 2.5% Tobra SLNs, as confirmed also by TEM. The amounts of Tobra in the
kidneys after Tobra-SLN administration, duodenally or IV, were lower than after IV
administration of the solution (59).
Idarubicin is an anthracycline administered by the IV or the oral route.
Its efficacy has been demonstrated in the treatment of different tumors. SLNs incorporating idarubicin (Ida-SLNs) (60) were administered IV and duodenally to rats in
a comparative study versus solution, the aim being to determine whether drug bioavailability can be improved. After duodenal administration of Ida-SLN and of Ida
solution, idarubicin and its main metabolite idarubicinol were determined over
time, as shown in Figure 3A and B; the AUC and elimination half-life of idarubicin
were, respectively, about 21 and 30 times higher than those with solution. Again the
AUC of Ida-SLNs administered intravenously was lower than when administered
duodenally. Pharmacokinetics and biodistribution of Ida-SLNs differ from those of
the solution. These changes could be exploited to reduce toxicity and increase clinical
efficacy of the drug.
FIGURE 3 (A) Idarubicin plasma
concentrations versus time after
duodenal administration of the two
formulations. (B) Idarubicinol plasma
concentrations versus time after
duodenal administration of the two
formulations.
Gastrointestinal Applications of Nanoparticulate Drug-Delivery Systems
313
COLONIC DRUG TARGETING USING NANOPARTICULATES
One way to target the colon is to incorporate drugs in appositely studied nanoparticulates. In one such study, coated calcium-alginate-gel beads were used to entrap
liposomes (61). Bee venom peptide was taken as a model drug, and investigated in
vitro and then in vivo: gamma-scintigraphy was used in a human study to determine
colonic arrival time, which was found to be four to five hours.
Bioadhesion of NPSs to inflamed colonic mucosa was shown to be sizedependent, using fluorescent particles, in a rat study taking nonulcerated tissue as
comparison (62). PLGA nanoparticles (63) carrying an antiinflammatory drug,
Rolipram, were thus studied for the treatment of inflammatory bowel disease.
Experimental colitis was induced in rats with trinitrobenzenesulfonic acid, which
produced significant damage of the intestinal tissue. Nanoparticles containing
Rolipram with a diameter between 200 and 500 nm were prepared; for better targeting, they combined efficient drug loading with an appropriate size range (64).
The particles were administered for five days and compared with the drug solution.
After five days, when drug treatment was discontinued, the drug solution group
underwent a severe relapse, whereas the nanoparticle group continued to show
reduced inflammation levels; the solution group had a higher adverse effect index
than the nanoparticle group.
SPECIFIC ADVANTAGES
Many adverse conditions in the GIT must be overcome to obtain the desired results
after administration of nanoparticulates. Both academia and industry have made
great efforts in this direction over the last 10 to 15 years, but more studies are
required to achieve the results that so many groups are working for. From the pharmaceutical standpoint, most nanoparticulates targeted to the lymph have increased
bioavailability versus the referee drug; this is particularly appreciable for labile
drugs and molecules with poor solubility. One factor requiring improvement for
some NPSs is incorporation efficiency, which must be increased in order to reach the
required dose. Studies are also necessary with regard to storage stability.
With regard to the different NPSs, the history of some of them, such as
dendrimers, is short and thus there are relatively few in vivo on the oral route.
Some polymeric biodegradable nanoparticles are already on the market, although
to date only for the parenteral route. Studies on the oral route have chiefly addressed
the administration of vaccine antigens, peptides, and small proteins. Provided that
the polymer is selected appropriately (type, molecular weight, kind of preparation),
the results obtained in vivo have been promising. Lipid-based systems have
achieved some of the desired results; some antiprotease drugs, ritonavir, and
saquinavir, carried by S(M)EDDS, are now on the market and provide better bioavailability than the referee drug (34). Some production methods of SLNs do not use
any solvent, their feasibility is relatively good, and their components are physiologically compatible.
SLNs loaded with a drug (Tobramycin or Idarubicin) and administered duodenally show better pharmacokinetic parameters than the same drug administered
IV as a solution. Tobra-SLN administered duodenally permitted absorption of
tobramycin by the GIT, interestingly, as this is a drug still only administered by the
parenteral route.
Interest in the oral administration of chemotherapeutics has been stimulated
by the discovery that oral fluoropyrimidines have at least equivalent efficacy, as
314
Gasco
well as the potential to reduce toxicity, compared to administration by the IV route;
using rational nanoparticulate design, several antineoplastic drugs could be developed for oral use (65). Although studies on nanoparticulates targeting the colon are
relatively few, some interesting results have been achieved using liposomes (61) and
PLGA nanoparticles (63).
EVALUATION AND FUTURE PERSPECTIVES
The main goal of administration of NPSs by the oral route is to lower the dose of the
drug (and consequently to diminish its toxicity) as well as to improve patient compliance and supply an easy administration route. Other aims may be to decrease the
fed/fasted variability and patient-to-patient variability. Efficient incorporation of
bioactive molecules in NPSs requires in-depth study. Factors that must be examined
in order to obtain the best formulation for the type of NPSs selected are: (i) sufficient
drug loading to achieve therapeutic levels; (ii) good translocation of NPSs to lymph
(i.e., small size, biocompatible, biodegradable components, chemicophysical
stability of carrier and drug, zeta potential, etc.); (iii) sustained/controlled drug
release from NPSs; and (iv) increased oral bioavailability to enhance efficacy.
Many steps have already been taken in these directions, but research must
continue, varying the surface properties of NPSs, improving their drug-loading
capacity, and testing new materials with little or no toxicity for use as carriers; this
last requisite is particularly important for drugs to be taken chronically. An innovative oral-delivery reformulation of a drug could extend its patent life; development
of new delivery systems for old, well-known molecules, whether natural or out of
patent, could lead to reduced side effects and to more efficient therapy, with evident
social benefits.
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20
Nanoparticles as Adjuvant-Vectors
for Vaccination
Socorro Espuelas
Department of Pharmacy and Pharmaceutical Technology, University of Navarra,
Pamplona, Spain
Carlos Gamazo
Department of Microbiology, University of Navarra, Pamplona, Spain
María José Blanco-Prieto and Juan M. Irache
Department of Pharmacy and Pharmaceutical Technology, University of Navarra,
Pamplona, Spain
INTRODUCTION
Looking for successful vaccines has become one of the driving forces in global
health. The history of vaccination is rich in many trials to treat numerous infectious
diseases. Vaccinations are responsible for approximately 25% of global mortality,
especially in children younger than five years (1). Live attenuated vaccines are still
by far the most utilized for their efficiency with respect to inactivated (i.e., bacterins)
and acellular or subunit ones, but is it necessary to run the risk of using live vaccinal
strains? Are there not other safer alternatives? Consequently, the need for better
vaccines clearly exists, but in spite of the progress made during the last few years,
“the ideal vaccine” against many severe infectious diseases still is not available.
The ideal live vaccine has to comply, at least, with the following requisites: (i)
innocuous for the vaccinated host, (ii) provide long-term protection with a single
vaccination, (iii) minimize the long-term production of antibodies, which may interfere with the current serodiagnosis tests of natural field infections, (iv) avoid the
contamination of food or any effect in the environment, (v) nonpathogenic for humans
or any other plant or animal host, and (vi) be biologically stable. In consequence, the
actual tendencies are oriented towards the development of new vaccines containing
perfectly characterized antigens, rigorously controlled during all the steps concerning
their preparation, and safe. However, the new vaccines of the new biotechnology era
suffer, in general, from immunogenicity, requiring the use of adjuvants (2,3).
The adjuvants have various functions, such as facilitate the passage through
cellular barriers and the stimulation of the cells of the immune system. The nanoparticles may be considered as adjuvants that can accomplish these requisites, being
able even to induce a response at the mucosa level given their capacity to interact
with it. We will review these carrier adjuvants in this chapter.
IMMUNOADJUVANTS
Adjuvants are compounds that enhance or modulate the immunogenicity of coadministered antigens. Specific antigen/adjuvant combinations preferentially induce
type 1 or type 2 cytokine responses. The Th1 subset secretes cytokines, including
interleukin-2 (IL-2) and interferon-γ (IFN-γ), to assist in cell-mediated immune
317
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Espuelas et al.
response. The Th2 subset assists preferentially in antibody immune responses after
secreting cytokines including interleukin-4 (IL-4). The mechanism of the adjuvant
action is in some cases, still unknown (4). In spite of large lists of compounds and
strategies described as adjuvants, the only FDA-approved adjuvant is alum, a general name for the aluminum-based mineral salt (5). It yields a reasonable antibody
response (Th2), but it does not induce a Th1 profile. Th1 immunity is essential for
protection against many infective organisms (e.g., intracellular parasites, including
virus and some prokaryotic and eukaryotic microorganisms) and even to limit allergenic processes (6). Moreover, aluminum adjuvants have shown limitations in their
applicability in vaccines based on small-sized peptides or antigen-expressing DNA
(7,8). Another limitation lies in the fact that aluminum-adsorbed vaccines are frostsensitive and thus not lyophilizable. Therefore, a large number of new adjuvants are
under investigation, such as the particulate-delivery systems.
It is well established that adjuvants can enhance the specific immune response
of the coadministered antigens by two major mechanisms (3,9):
1. They increase the antigen uptake by antigen-presenting cells (APCs) (“delivery”).
The adjuvants are directly engulfed by APCs or they form a depot of the antigens
that prolong their exposure and so the chance to be taken up by APCs.
2. Another group of adjuvants (“immunopotentiators”) directly activates innate
immune cells. To this category belong inflammatory stimuli (frequently associated
with vaccine administration), cytokines, CD40L, and pattern-associated molecular
pathogens (PAMPs: molecular patterns shared by multiple microorganisms, not
presented in mammalian cells, that activate immune cells through interaction
with pattern-recognition receptors). Lipopolysaccharide, murine, flagellin, and
other main structural components of bacterial cells are examples of PAMPs used
empirically for a long time (4).
Particulate-delivery systems belong to the first category of adjuvants (10,11),
although their activity as immunopotentiators cannot be completely discarded.
NANOPARTICLES AND ANTIGEN-PRESENTING CELLS
Numerous reports indicate that particulate antigens are naturally captured by APCs,
but also that size is a decisive factor for an adequate uptake. Dendritic cells (DCs), a
paradigm of APCs, are able to capture particles in vitro and in vivo but to a lesser
extent than macrophages and the optimal size is smaller than that described for
macrophages (in the range of 500 nm or below) (12,13). In addition, the surface
charge and hydrophobicity of nanoparticles also modify their uptake (14). Once
inoculated, the particle surface changes by the adsorption of endogenous proteins
presented in serum and in the interstitial fluid at the administration site. The more
hydrophobic the surface is, the easier the coating process by the dominant proteins
(opsonins as IgG or C3 complement factors) can be (15). Macrophages and DCs have
receptors for Fc fraction of IgG, C3 complement component, or the C-lectins family,
such as mannose receptors (16). Coating nanocarriers with these ligands has been
proven a good strategy for improving antigen delivery and targeting. Furthermore,
these ligands have influence in the profile of the elicited immune response (17).
THE FATE OF NANOPARTICLES AFTER ADMINISTRATION
Parenteral Immunization
Nanoparticles administered intramuscularly (IM) or subcutaneously (SC) are
expected to remain essentially at the site of administration and control the release of
Nanoparticles as Adjuvant-Vectors for Vaccination
319
the loaded antigen. Slowly degradable polymers [i.e., poly(lactide acid) (PLA),
poly(lactide-co-glycolide) (PLGA), poly(methylmethacrylate) (PMMA), poly(alkylc
yanoacrylates), or caprolactone polymers (PEC)] are suitable for this application as
they provide adequately prolonged antigen release. Nanoparticles of PMMA were
claimed to be biodegradable after SC or IM injection and shown to exhibit very
powerful adjuvant properties for a number of antigens. Thus, they were able to
improve the antibody response and the protection against influenza challenge in
comparison to the classical adjuvant aluminum hydroxide (18).
Desai and coworkers (19) have demonstrated adjuvant properties of PLGA
nanoparticles containing encapsulated staphylococcal enterotoxoid B. However, the
systemic antibody immune response elicited in the animals injected SC with these
nanoparticles was comparable to that obtained following injection of alum. The
same group (20) showed in a further study with tetanus toxoid (TT), that coinjecting
subcutaneously to rats TT-alum along with TT-loaded PLGA nanoparticles induces
a synergistic immune response. The combination induced a fourfold greater mean
serum anti-TT IgG response than a single injection of TT nanoparticles alone.
In another study, a single immunization of PLA nanoparticles entrapping
immunoreactive TT administered IM to rats elicited anti-TT antibodies that persisted for more than five months (21).
Recently, Schöll et al. (22,23) demonstrated the effectiveness of PLGA nanoparticles in immunotherapy of allergy to birch pollen (Bet-v1) in BALB/c mice. PLGA
nanoparticles loaded with Bet-v1, the major allergen, enhanced the right immunogenicity of the allergen and did not lead to novel sensitization of naïve animals. The
authors assumed that vaccination with PLGA nanoparticles can counterbalance an
ongoing Th2 response to Bet-v1, readdressing it to a Th1-protective response.
Mucosal Immunization
Although most vaccines traditionally have been administered by IM or SC injection,
mucosal administration of vaccines offers a number of important advantages,
including easier administration, reduced adverse effects, and the higher mucosal
immune response achieved. Another important advantage is that the vast majority
of infections are from a mucosal surface and protection at this port of entry level is
essential (gastrointestinal, respiratory, and urogenital tracts).
Oral Route
The most attractive, easiest, and most acceptable route for mucosal immunization is
the oral one. However, as a result of the acidity in the stomach, an extensive range
of digestive enzymes in the intestine, and a protective coating of mucus that limits
access to the mucosal epithelium, the oral delivery of vaccines remains a challenge
where it is difficult to achieve high and reproducible effects. In order to solve these
difficulties, nanoparticles may be useful to protect antigens and facilitate the interaction with components of the gut mucosa by a mechanism of bioadhesion.
In mice, oral immunization with PLGA nanoparticles has been shown to
induce potent mucosal and systemic immunity to entrapped antigens (24,25). For
instance, the encapsulation of Helycobacter pylori antigens in PLGA nanoparticles induced significantly specific mucosal IgA response as well as serum IgG1
and IgG2b responses (24). Another interesting study was based on the use of TT
associated with sulfobutylatedpoly(vinyl alcohol)-graft-poly(lactide-co-glycolide)
nanoparticles, that a given by peroral (PO) or intranasal (IN) route induced specific
IgG and IgA in mice in a reproducible manner (26).
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Espuelas et al.
Oral vaccination using chitosan nanoparticles as vector for Toxoplasma gondii
GRA 1 pDNA primed an antibody-mediated immune response, although not the
cellular protective one (25). Similarly, it was shown that mucosal immunity towards
Tat protein could be triggered in mice by oral or rectal route when this protein is
loaded in chitosan nanoparticles. Tat protein plays an essential role in HIV-1 replication and also participates in T-cell immunosuppression (27). In fact, sera from mice
immunized with chitosan nanoparticles inhibited the activity of Tat protein, which
is a prerequisite for the development of an anti-AIDS-protective vaccine (28). In
addition, this vaccine also induced cell-mediated immunity (28).
More recently, chitosan nanoparticles have also been proved effective as adjuvant immunotherapy (29,30). Thus, DNA encoding allergens from house mites were
complexed with chitosan and delivered orally, inducing a successfully specific Th1immune response (30). Finally, of special interest in immunotherapy of allergy and
treatment of diseases induced by a disbalance in the Th1/Th2 response such as
autoimmune disorders, is the capacity of the adjuvant to induce IL-10 (31,32). Thus,
pegylated nanoparticles containing ovalbumin were able to induce significant and
sustained levels of IL-10 for at least six weeks in mice (33).
In any case, it appears that “conventional” nanoparticles offer a limited ability
as adjuvants for oral vaccination. An interesting strategy to solve this problem may
be the use of nanoparticles coated with bacterial adhesins in order to obtain a similar distribution within the gut to that observed for the pathogen. In this context, we
have evaluated the properties of Gantrez nanoparticles coated with flagellin from
Salmonella enteritidis (34). In vivo, these carriers displayed a similar distribution
within the gut as the whole bacteria interacting strongly with the enterocytes and
Peyer’s patches (Fig. 1).
Nasal Route
In contrast to oral administration, nasally administered vaccines have to be transported over a very small distance and are not exposed to low pH values and degrading enzymes. In addition, the nasal mucosa shows a relatively good permeability
for proteins, compared to the gastrointestinal tract. Furthermore, nasal delivery of
antigens allows a direct access to the nasal-associated lymphoid tissue (NALT), the
major site of antigen uptake and presentation, whereas oral delivery requires diffusion of the antigen carriers to extensively widespread lymphoid aggregates.
Various authors have studied the nasal mucosa as a potential route of vaccination against tetanus, as nanoparticles have been demonstrated to be a useful vector
FIGURE 1 (See color insert.) Visualization
of flagellin-coated nanoparticles (red dots)
in the follicle-associated epithelium of
Peyer’s patches by fluorescence. Flagellin
of S. enteritidis was used to coat Gantrez
nanoparticles (labeled with rhodamine B
isothiocyanate) and, thus, obtained adjuvant
vectors able to spread within the gut in a
way similar to the whole bacteria. Source:
From Ref. 34.
321
Nanoparticles as Adjuvant-Vectors for Vaccination
system for TT. The intranasal application of that formulation resulted in significantly increased IgG and IgA levels compared to nontreated animals (26,35,36). For
other clinically relevant antigens, it was found that, in general, intranasal vaccination requires fewer administrations at lower dosing levels to stimulate immune
responses than orally. This fact can be explained by a lower efficient particle uptake
from the gut and a more destructive gastrointestinal environment for proteins (26).
Table 1 summarizes these experimental works.
TABLE 1 Some of Studies Using Nanoparticles as Vaccine Adjuvants for
Nasal Administration
Formulation
Antigen
SB-PVAL-gPLGA NP
Tetanus toxoid
Species
Mice
Results
Smallest particles induced the
most significant antibody
responses
Tetanus toxoid
Rat
Intranasal administration of
Chitosan NP,
chitosan NP formulation
CS-PLGA NP,
provided high and longPLA–PEG NP
lasting immune response
PLA NP, PLA– Tetanus toxoid
Mice
Pegylation increases the
PEG NP
stability of resulting
nanoparticles and elicited
higher antibody levels
Chitosan NP
Tetanus toxoid
Mice
6 months postadministration,
the IgA response was
significantly higher than for
a control vaccine
PLA–PEG NP
Tetanus toxoid
Rat
A decrease in the size
improves the transport
across the nasal mucosa
PLGA NP
Influenza virus
Rat
Smaller nanoparticles
antigens
provided better immunization than larger ones
Mice
Encapsulation of antigen in
PLGA NP,
Bovine paraPLGA NP provided higher
PMMA NP
influenza type 3
levels of specific antibodies
proteins
than adsorption in
PMMA NP
Polystyrene NP Concanavalin AMice,
Vaccines provided high
inactivated HIV-1
macaque
immune responses and
provided significant
protection against intravaginal challenge
Mice
Induction of protective
Calcium
Herpes simplex
immunity against intravagiphosphate NP
virus type-2
nal challenge
glycoprotein
Mice
Immunization induced
Chitosan NP
pDNA expressing
protection against challenge
either hemagglutiwith influenza A virus
nin HA or nuclear
protein of the
influenza A virus
References
(28)
(37)
(36)
(38)
(35)
(39)
(40)
(41,42)
(43)
(44)
Abbreviations: CS-PLGA NP, chitosan-coated poly(lactic acid–glycolic acid) nanoparticles; PLA NP, poly(lactic
acid) nanoparticles; PLA–PEG, poly(ethylene glycol)-coated poly(lactic acid) nanoparticles; PLGA, poly(lactideco-glycolide) nanoparticles; PMMA NP, polymethylmethacrylate nanoparticles; SB-PVAL-g-PLGA NP, sulfobutylated
poly(vinyl alcohol)-graft-poly(lactide-co-glycolide) nanoparticles.
322
Espuelas et al.
In contrast, pegylation appears to be a good strategy to enhance the ability of
nanoparticles as adjuvants for nasal vaccination. In fact, IgG levels elicited by TT
loaded in PLA–PEG nanoparticles were significantly higher and longer-lasting than
those corresponding to both the fluid vaccine and conventional nanoparticles (45).
Moreover, this potential of pegylated nanoparticles was also demonstrated for a
model DNA nasal vaccine, where a single nasal dose of DNA PEG-nanoparticles led
to a significant antibody response to the encoded protein (37,38). All of these results
concerning pegylated nanoparticles can be directly related to the ability of PEG
coating to facilitate the internalization of nanoparticles through a given epithelium
(33,37,46). It appears that the PEG coating has a role in stabilizing nanoparticles in
mucosal fluids, facilitating the transport of the nanoencapsulated antigen and,
hence, eliciting a high and long-lasting immune response.
Comparing PEG-nanoparticles with chitosan-coated nanoparticles, it was
observed that PEG-nanoparticles were more efficient than chitosan-coated nanoparticles in facilitating the transport of the associated antigen (TT) through the mucosa.
The explanation was found in the different interaction mechanisms of both types
of modified nanoparticles with the nasal epithelium (35). In particular, PEGnanoparticles did not interact with the mucin, whereas those coated with chitosan
were designed to stick to the mucus layer (47,48), impeding the transport across
the nasal epithelium. Again, modification of the physicochemical characteristics of
nanoparticles may facilitate the diffusion of nanoparticles through the protective
mucus layer and their interaction with the NALT. In this context, it has been
described that a decrease on the size of PLGA nanoparticles resulted in an increase
in their ability to reach the NALT and to potentiate the immune response (39). In
another interesting work, two different types of nanoparticles were evaluated
as adjuvants for nasal immunization against bovine parainfluenza type 3 virus. In
this work, viral proteins were either encapsulated in PLGA nanoparticles or
adsorbed in PMMA nanoparticles (44). Mice immunized with a single dose of PLGA
nanoparticles developed higher levels of virus-specific antibody than mice immunized with the PMMA vaccine or with the viral proteins alone. In any case, these
results can be due to the method of association between antigen and nanoparticle
(encapsulation vs. adsorption) or to the intrinsic ability of the polymer to act as
adjuvant (PLGA vs. PMMA) (44).
Miyake and coworkers described the ability of polystyrene nanoparticles
coated with Concanavalin A to efficiently capture HIV-1 particles (40). After intranasal immunization, this formulation induced a high vaginal IgA antibody response
in mice. In macaques, after a series of six intranasal immunizations with this formulation and intravaginal challenge, only 33% of animals became infected. Similar
results were obtained with a derived vaccine obtained by mixing ConA-NP with
inactivated HIV-1 particles (41).
In mice, nasal immunization with herpes simplex virus type-2 glycoprotein in
calcium phosphate nanoparticles induced protective immunity against an intravaginal challenge (42). Also in mice, mucosal immunization with chitosan–DNA
nanoparticles induced protection against challenge with influenza A virus (43).
Recently, a chitosan-influenza vaccine was found effective in humans when administered by the nasal route (49).
Overall, it is reasonable to consider that the nasal route seems to be more
promising for vaccination in mice than the oral one for nanoparticle vaccine administration (26).
Nanoparticles as Adjuvant-Vectors for Vaccination
323
Other Mucosal Routes
Few recent studies have focused on pulmonary immunization and the involvement
of the pulmonary immune system in eliciting protective immune responses against
inhaled pathogens. In this context, the pulmonary administration of pDNA (encoding protective epitopes from Mycobacterium tuberculosis) loaded in chitosan nanoparticles induced increasing levels of interferon-gamma compared to either a control
formulation or the more frequently used intramuscular immunization route (50).
In contrast, calcium phosphate (CP) nanoparticles, containing HSV-2 viral
proteins, have been proved effective in enhancing the mucosal immune response
against the virus. In mice, immunized intravaginal immunization with HSV-2
loaded in CP nanoparticles induced high levels of HSV-specific mucosal IgA and
IgG in vaginal lavage fluids. Furthermore, mice vaccinated with this formulation
were protected against challenge, with higher survival rates and less clinical infection than unvaccinated controls (42).
In summary, nanoparticles containing entrapped or adsorbed antigens are
being investigated as vaccine adjuvant alternatives to the currently used alum with
the objective of developing better vaccine adjuvants and minimizing the frequency
of immunization. However, little success has been proved by the oral administration of antigens or allergens in nanoparticles. Accumulated experimental evidence
suggests that simple encapsulation of vaccines into nanoparticles is unlikely to
result in the successful development of oral vaccines, and improvements in the
current technology are clearly needed. In contrast, the nasal route seemed to be
more promising for vaccination than the oral one for nanoparticle vaccine-based
administration.
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21
Transdermal Applications of
Nanoparticulates
Jongwon Shim
Nanotechnology Research Team, Skin Research Institute, R&D Center,
Amorpacific Corporation, Kyounggi, South Korea
INTRODUCTION
As the principal purpose of using nanoparticulates lies on increasing bioavailability
of drugs, the more accurate definition of transdermal applications should be the
applications of transdermal drug delivery with nanoparticulates. The pharmaceutical industries have applied these nanoparticulates to deliver drugs requiring acute
treatment or to formulate the nonsoluble drugs into the practical formula enabling
the in vivo applications. Of course, there are some topical applications of nanoparticulates targeted to the skin organ, but most of the pharmaceutical applications
with nanoparticulates have been developed to give systemic effects to the body by
penetrating the full thickness of skin layers (1,2). For these purposes, the drugs
encapsulated by nanoparticulates have to penetrate various regions of the body.
Many mechanical and electrical delivery systems for nanoparticulates are available
using injection, patch, and electrophoresis methods other than the simple topical
administration of spreading the drug onto skin surface (3). However, the use of
nanoparticulates in the cosmetic industry has led to expectation of the efficacy of
drugs on the skin itself in most cases. It is used to make the skin tones elegant, to
prevent the formation of melanin by UV light, to degrade callous layers, to remove
wrinkles, and to moisturize skin. The typical difference of a nanoparticulate system
applied in the cosmetic industry from that of the pharmaceutical industry is that the
major delivery method is very limited to the topical application of spreading onto
skin surfaces, and it also serves the emotional function of provided beauty and
psychological satisfaction.
But it is certain that the current major trend of interdisciplinary research is
conducted without any limitations, and the same trend is applied in all industries.
From the above, the application of nanoparticulates has been discussed by using
two industry categories for convenience, but it is true that such classification seems
to get more obscure as time goes by. The newly emerging cosmeceutical products in
the cosmetic industry propose more enhanced efficacy than previous products.
These products are available not only for skin whitening, hair removal, hair growth,
antiwrinkle, anticellulite, and acne treatments, but also for atopic dermatitis and
cosmetic supplementary foods for skin health. The same trend is going on in pharmaceutical industry in the form of noninvasive treatment in the dermatological area
for cosmetic purposes. Namely, the treatment methods such as some laser therapy
for the pore-tightening, chemical peeling, and ion-treatment methods have been
performed to enhance the skin condition or texture and to provide a semipermanent
makeup by penetrating colorants in epidermal layers. In the past, only the efficacy
of a product was appreciated in pharmaceutical product development but, now,
the development of pharmaceutical products that were only differentiated by color,
327
328
Shim
FIGURE 1 (See color insert.) Cryo-sectioned images of the fluorescent labeled nanoparticulate
with porcine epidermal skin layer after in vitro skin permeation test. (A) Stratum corneum outer layer
adsorption of 200 nm-sized polysturene nanoparticulate. (B) Percutaneous absorption of 40 nmsized polystyrene nanoparticulate. (C) Epidermal layer penetration of the modified nanoparticulate.
Abbreviations: EP, epidermis; SC, stratum corneum.
fragrance, taste, and texture are currently expected to meet the consumer’s complex
demands. Accordingly, it seems not desirable to classify the nanoparticulate system
for transdermal delivery in cosmetics by terms used to describe its transdermal
delivery of drugs. So, the current chapter is intended to describe the transdermal
applications of nanoparticulate by the depth of skin to which nanoparticulates are
delivered: stratum corneum adsorption, percutaneous absorption, and whole skin
penetration (Fig. 1).
CLARIFICATIONS IN THIS CHAPTER
Permeation
The terminology in skin delivery indicates the phenomenon of molecular movement of active ingredients into skin. As in the case of delivering some ingredients by
spreading onto a skin surface, these molecules are generally dissolved in some type
of liquid-phase vehicles or they should have fluidity by themselves, and this mass
transfer follows Fick’s diffusion law. The required driving force for the diffusion is
the thermodynamic energy of the concentration gradient (4).
Penetration
The term “penetration” indicates the skin delivery of particulates by colloidal movement, and this particulate is homogeneously dispersed in some medium. The
Transdermal Applications of Nanoparticulates
329
status of particulates dispersed in a medium could be dependent on the particulate
properties including its size and mass. As if particulates belong to the region that is
not affected by gravity like molecules, they seem to diffuse by thermodynamic
energy and the diffusion is in a direction that can reduce the overall entropy value
of the system.
Nanoparticulates
Nanoparticulates indicate various types of subsidiary concepts of particulates that
can be represented as capsules, aggregates, powders, crystals, micelles, emulsions,
complex, and vesicles. It is usually used to designate a particulate whose range size
reaches a submicron level. However, the current chapter uses the terminology not
only to indicate particulates with submicron size, but also to indicate the system
showing a colloidal dispersion within appropriate vehicles.
STRATUM CORNEUM ADSORPTION
The stratum corneum is the location where the transdermal delivery is initiated,
and as in the most molecules, it is also poweful barrier to nanoparticulates (5).
Physically considered, skin has irregular surfaces showing elevation differences
between surfaces, which makes it difficult for large materials greater than the elevation differences to have enough contact area on the skin surfaces, and an external
force can easily detach them (6). Primarily, there are application areas that utilize the
small sizes of nanoparticulates to achieve even dispersion and strong adhesion onto
the skin surface (7). These nanoparticulate applications are widely developed in the
cosmetic industry to be used as organic or inorganic pigments to manufacture cosmetics for color makeup, and there are inorganic nanoparticulates proposed for UV
light protection. However, it is never desirable for such inorganic nano-sized
materials to be penetrated under the stratum corneum, but rather it is recommended
they be removed naturally by the turn-over period of stratum corneum or by
washing (8–10).
In the aspect of drug delivery, it is most ideal for drugs dissolved in a liquidphase vehicle at the molecular level to be evenly spread onto skin surfaces to achieve
the high transdermal delivery efficiency. But if such ideal solvents or vehicles do not
exist or these molecules are repulsive to the skin surface, it is advantageous to use a
nanoparticulate system that has high affinity for skin. So, the use of solid–lipid
nanoparticles or a polymeric nanoparticulate that was modified to have a surface to
be easily adsorbed by skin surface is appropriate (11,12). By using such materials,
the nanoparticulate encapsulating the drugs can be evenly spread onto skin surfaces, and after the medium is either evaporated or absorbed, the strong adhesion
force between nanoparticulates produces membrane-like substrates, which function
as a type of patch with high drug concentration (11). On the basis of this, the drug
concentration in the skin surface is rapidly elevated and the strong driving force of
the concentration gradient stimulates the drug delivery. For such a mechanism, it is
essential to assume that the drug encapsulated by the nanoparticulates should be
released into skin easily. The appropriate kind of drugs for such a nanoparticulate
application is determined by the characteristics of molecular weight or hydrophilic
property of the drugs, which have a moderate range of skin absorption availability,
but no affinity with skin surface. The drugs for commonly used pharmaceutical patch
products could be the appropriate level of drugs that can apply in this type of
nanoparticulate applications.
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Shim
PERCUTANEOUS ABSORPTION
In cases when the drug molecules can penetrate stratum corneum, and molecules
showing rapid molecular dispersion in water phase could be dispersed into whole
skin without many problems due to the high moisture level with low cell density
under the epidermal layer (14). So, the penetration of drugs or particulates into stratum corneum becomes the most important task in the transdermal delivery, and
currently two approaches are available:
1. The use of materials that can disturb or remove the barrier function of stratum
corneum or the mixed use of particulates with mechanical devices or with
electrical devices.
2. The noninvasive applications of special particulates that can enhance the
penetration without damaging the barrier function of skin.
The first method could be reclassified into passive and active approaches. The
passive method delivers the drug molecules through simple diffusion by disturbing
the barrier function of the stratum corneum with permeation enhancers that can
perturb, extract, and exchange the molecules which comprise the lipid bilayer of
stratum corneum, with solvents that can solubilize lipids, or with surface-active
agents appropriate to skin use (15,16). As the most representative formulations in
cosmetic products, emulsions and micelles can be the particulate systems that apply
to the first method. As they are comprised by oils and surface-active agents having
affinity to skin, the effective perturbation of stratum corneum is available when
being spread onto skin, which will facilitate easier drug delivery down into the
epidermal layer. The drug property that fits this delivery system can be water-soluble or oil-soluble ingredients, or they can be micellized by using surfactants and
appropriate solvents. Furthermore, if the droplet size of the emulsion particulate
can be uniformly made to nanoscale, this emulsion of nanoparticulates can enhance
drug permeation and systemic efficiency of drugs (17,18). However, because of
the difficulty of isolating drugs from external environments for these emulsion or
micelle particulates, the ability to prevent the drug degradation can be hardly
expected. So, if the drug to be captured is unstable in oil or water phase, it is not the
appropriate system to be used. For such drugs, various stabilizers and antioxidants
should be blended to acquire enough stability to achieve effective drug delivery
during the period of use. In addition, it is hard to achieve the enhancement of drug
delivery under the epidermal layers by the active transport of particulates because
these two types of particulates release drug as they contact skin surface by the
absorption of surfactants or oils onto skin surface, or by losing their particulate
shape due to the surface energy change (19). Meanwhile, the active approach enables the powder type of drug particulates to be directly transported by applying
instantaneously strong pressure of fluid onto skin surface. Also, the use of surfacecharged particulates and the use of iontophoresis or electrophoresis to enhance the
penetration of drugs are also available. Although these types of methods could
enhance the drug’s skin delivery very effectively by minimizing the damage on skin
stratum corneum, it is absolutely necessary to use specially manufactured instruments for the operations, and the additional cost and the usability depreciation have
to be considered (20,21). Contrary to the emulsion or micelle system, the second
method is a drug-delivery method without causing skin damage by using nanoparticulate which does not lose its original property of particulate while penetrating
the stratum corneum of the skin surface. The special types of liposomes and
Transdermal Applications of Nanoparticulates
331
polymeric nanoparticulates could be suggested as its most representative examples.
As a special type of liposome, there is lipid vesicle system that could be represented
as Transfersome. It has been known as the vesicle membrane that has ultradeformability and can overcome the instability of typical liposomes in skin surface,
which can penetrate skin appendages and even the intracellular space of skin
smaller than the size of vesicle itself (22,23). As in the case of polymeric nanoparticulate dispersed in a water phase, recent studies have revealed that most of the
particulates penetrate through the skin appendages and shunts that have been
known to be the pathways for large hydrophilic molecules (24,25). Especially the
hair follicles were regarded as the major pathway, and transport efficiency seems to
be determined by the average particulate size. So, the smaller size of nanoparticulate could transport the drug more easily under the epidermal layer (26–28) (Fig. 2).
As the polymeric nanoparticulates have different absorption characteristics due to
the material property of polymers, the drug encapsulation and release behavior
within the nanoparticulate should be expected to be different from that of microparticulate systems. Previously, the skin absorption tendency of the drug by nanoparticulate was enhanced or delayed depending upon the drug types, questioning the
true efficacy of the nanoparticulate (29–31), but it is more appropriate to interpret
the problem by the way that the chemical property or surface characteristics like the
partition coefficient of captured molecules cannot be completely covered by the
nanoparticulate encapsulation. But yet, there are no proven results regarding this
matter, and further studies to resolve the problem should be made. If any of these
methods has been selected to deliver the drug to the epidermal layer, the suitable
drugs should be active materials with similar functions as keratinocytes and melanocytes in cosmetic products. Namely, these active materials can degrade melanin or
inhibit the formation of it from melanocytes. As a pharmaceutical product, the
steroidal drug of hydrocortisone could be suggested as its example, which alleviates
the atopic symptom by the action on Langerhans cells existing in the epidermal
layer. It also shows its effectiveness on erythema, psoriasis, and dermatorrhea.
However, the long-term use of this drug can causes severe side effects. So, the delicate design has to be made to maximize its efficacy by using an effective delivery
system and to minimize the side effects.
FIGURE 2 (See color insert.) CLSM images of the nanoparticulate penetration with a guinea pig
skin (cryo-sectioned). Fluorescent-labelled nanoparticulate (green region), nuclei-labeled DAPI
(blue region). (A) 40 nm size of nanoparticles, (B) 130 nm size of nanoparticles (10 pieces of
z-direction sectioning image of cross sectioned tissue are merged). Nanoparticulate: Polycaprolactone-polyethyleneglycol block copolymer aggregates. Fluorescent: Rubrene. Source: From Ref. 27.
332
Shim
FIGURE 3 (See color insert.) Images of the liposome nanoparticulate with fluorescent by in vitro
permeation test with a guinea pig skin. (A) RITC saturated solution (B) RITC with the modified
liposome. (Red region), hydrophilic fluorescent dye (RITC); (blue region), DAPI for nuclei; basal
layer (white arrows).
As mentioned previously, if all material delivery and diffusion can be easily
achieved after penetrating the stratum corneum, it seems to be not required for
nanoparticulates to penetrate the epidermal layer. However, some kind of drugs
could not penetrate the epidermal permeability barrier (EPB) which has the basal
layer (dermal–epidermal junction layer). EPB is thought to act as the barrier to
water loss, intestinal fluid, and infection of microorganisms (32). It is supposed
that the high cell density and abundant amounts of materials comprising the extracellular matrix and tight junction between cells in this layer might affect drugpermeation behavior. So, the particulate that can penetrate the whole epidermal
layer should be used when delivering drug to the dermal layer (Fig. 3). As for the
cosmetics to be used on the dermal layer level, the antiwrinkle application to
enhance the collagen synthesis and the anticellulite application to be used to
destroy and to remove the irregularly presenting subfats under the dermis could
be suggested as appropriate products for use, and applications for deep wound
healing could be suggested for effective use in pharmaceutical products. Although
it is differentiated by the depth of wound and the degree of its seriousness, if a part
of skin was injured, the injured part is initially blocked by blood coagulation to
prevent blood and body fluids loss, and rapid epidermal cell divisions are carried
out to recover the injured skin surface. But the recovery for the injured dermis
takes more time to progress. At that time, the insufficiency or over-multiplication
of dermal cells can cause the scar in the skin. So, if the nanoparticulates are directly
delivered down to the dermal layer to release the drug in concentrative manner,
the resulting rapid multiplication of fibroblast cells will increase cell density, and
can stimulate the normal synthesis of collagen to have the wrinkle-free skin
promised by cosmetic products; the application of it seems appropriate for the
restoration of scar-free skin as promoted in pharmaceutical products.
WHOLE SKIN PENETRATION
The particulate system designed for whole skin penetration has the purpose of
enhancing the overall bioavailability of drug molecules which are required to be
delivered to the target organs in spite of low permeability due to large molecular
weight and low partition coefficient on skin. For the nanoparticulate systems,
Transdermal Applications of Nanoparticulates
333
needed to penetrate the skin, nearly identical particulate characteristics are required
to deliver the drugs to dermal or epidermal layer as just discussed above.
Considered from the aspect of drugs to be absorbed, material partitioning into
skin can be inhibited by corneocytes and by various compositions of lipids and
fatty acids that fill up the gap. That is why the particulate system is primarily
formulated to penetrate the stratum corneum, but once penetration below stratum
corneum is accomplished, free diffusion is allowed by the abundant amount of
humoral fluid in skin tissue that can be delivered into the whole body through the
circulation system. Although the nanoparticulate system may succeed in penetrating the stratum corneum, the efficiency of delivering the drug into the circulation
system can be downgraded if the drug cannot permeate the epidermal and dermal
layer or cannot avoid various types of existing defense systems in skin. Otherwise,
if the drugs have lipophilic property and low molecular weight, the initial partitioning onto lipid compositions in the outermost layer of skin can be easily achieved,
but delivery under the stratum corneum layer with the relatively high moisture
content in epidermal and dermal layers will get more difficult for the mass transfer
of the molecules. So, these molecules will be clusterized due to the decreased
solubility in the environment, and eventually the mobility of these molecules could
be nearly disabled by diffusion.
Even if nanoparticulates can penetrate whole skin layer and can avoid various
immune systems presenting in the skin organ after dispersion, they have to face the
highly developed and more complex defense system in the body such as the blood–
brain barrier, Langerhans cells and T-cells in liver, Peyer’s patches in the gut-associated lymphoid tissue, and so on (33,34). So, a high delivery efficiency could not be
expected unless a proper counterplan for this is prepared. As in the case of percutaneous absorption, the particulate for skin penetration should not be destroyed by
external environments between the early steps of skin penetration and the delivery
into the circulation system, and must have material properties that can suppress
drug release and avoid degradation of the encapsulated drugs. Although there is
much unclarified information, if the particulate types of micelles, emulsions, and
generally manufactured liposome vesicles were applied by spreading onto skin surfaces, it is difficult to maintain their original particulate shape due to the absorption
behavior between skin surface and molecules comprising the particulates and by
the difference of surface energy (35). In some cases of their practical uses, it was
reported that the particulates were destroyed by the salts existing on exudates or
sweat ducts in the skin surface. There are many cases that cannot maintain the
stability of the system due to the rapid evaporation of water or other solvents in
the skin surface, and if the product requires spreading action by users, the external
physical stress of spreading may cause the structural destruction of the particulate to occur. So, these systems are not an appropriate system because the particulate
cannot be maintained until it reaches the destination due to burst-releasing the drug
premature (Fig. 4).
To avoid such problems, the most easily applied method is injection administration. The material is delivered down to the dermal layer in the skin barrier by
using a mechanical device. This method is not strictly restricted by particulate size,
and it also allows direct injection into blood vessels to have immediate efficiency of
the drug. Additionally, it can avoid the particulate disruption caused by environmental change and by the primary protection mechanisms existing in skin. However,
the injection administration method is stressful to the user and its use is difficult,
which seems disadvantageous in long-term and repeated use. The method has
334
Shim
FIGURE 4 (See color insert.) Fluorescent images of in vivo permeation study with Albino Hartley
guinea pig. (A) O/W emulsion formulation. (B) Modified polymeric nanoparticulate.
been improved to create the patch formulation method, but it also has some disadvantages due to the action of barrier function in skin by swelling and by the aspect
of having relatively low delivery efficiency. As for the system to be used easily, to
minimize skin damage and to deliver a higher concentration of drug to the circulation system, the use of particulate that can penetrate the whole skin layer such as
the polymeric particulate, other solid type nanoparticulates, or special lipid vesicle
particulate seems appropriate. Ethosome™ has been known as a liposome system
that can maintain high ethanol content, which can effectively disturb the skin lipids
without losing its particulate property to enhance the drug-delivery efficiency
(36,37). In addition, the previously mentioned Transfersome has shown high
efficiency of drug delivery for whole body application in addition to the function
of delivering drugs to skin layers. However, it is still difficult to find a commercialized polymeric nanoparticulate for external skin use as well as for whole body
functions of a drug. Its limited delivery pathway other than hair follicles make
it difficult to increase the delivery efficiency, its inability to capture hydrophilic
drugs, and the very limited amount of applicable materials are thought to be the
appropriate reasons.
THE LIMITATIONS OF NANOPARTICULATE SYSTEM
FOR TRANSDERMAL APPLICATIONS
First, there are limitation of acquiring stable materials. Most of the nanoparticulates and their debris that penetrate deep into skin layers might cause an immune
response. So, materials that can be disintegrated by appropriate mechanisms
to be absorbed into body have to be used. However, very limited numbers of
materials exist for a nanoparticulate system, which can be degraded within skin
to be absorbed naturally into the body. This could be suggested as the major
topic that researchers speculate, the conventionally used U.S. Food and Drug
Administration (FDA)-approved materials of poly(lactic acid) (PLA), poly(glycolic
acid) (PGA), and polydioxanone (PDS) that are using lecithin or biocompatible
polymeric biomaterials; these have nearly reached their maximum capacity to
enlarge their field of application. Similar degradable polyester-type polymers of
poly(anhydride), polyphosphonate, polyamide and polyiminocarbonate, polycaprolactone, polyphosphazene, and poly(phosphate ester) are available, but
Transdermal Applications of Nanoparticulates
335
cannot be used because their official approvals have not been granted. For these
reasons, many investigators currently make every effort to find and to apply
materials that can be as safe as lecithin in the human body and materials having
clear degradation mechanisms that can be synthesized and purified such as PLA
(38,39). Although these nanoparticulates will be formulated from safe materials
for human body, consideration for a different aspect of the material not shown in
the bulk phase might be necessary when these materials are formulated in the
form of nanoparticulates and when these nanoparticulates are actually used by
encapsulating drugs. As an example, some kind of nanoparticulate system is proposed for skin penetration, and the enhanced skin absorption rate has the risk of
causing strong irritation or immune response due to the overstimulation of the
skin immune system by the drug. So, the nanoparticulate system for transdermal
delivery should be investigated in the appropriate concentration ranges to represent its safety and efficacy from the following four aspects: drugs, raw materials
comprising nanoparticulate, drug-encapsulated nanoparticulate, and the empty
nanoparticulate.
Secondly, a limited delivery pathway could be suggested. In spite of the
excellent skin absorption rate of general skin products of external use, there are
still some problems remaining to be solved. Especially for the nanoparticulates to
penetrate lower than the stratum corneum, use injured skin parts, hair follicles,
and auxiliary skin systems as the pathway to move. But the area of the pathways
provided by these auxiliary skin systems make up a very small portion compared
to the total skin area. The skin absorption of the drug using the limited pathways
essentially requires a larger range of administration and repeated use to reach the
desired drug concentration in a body compared to the oral and injection administration methods.
Thirdly, there is an absence of proper stabilization functions for unstable active
ingredients. Because of the relatively thin wall thickness of nanoparticulate compared to microparticulate, it is beyond the capability of nanoparticulate to prevent
the material degradation by reactive oxygen species, UV light, and heat. Practically,
the active ingredients of cosmetic products are easily degradable by oxygen and
heat (40). If a nanoparticulate system is applied to enhance bioavailability of actives
to the skin, it is often necessary to add various additives, like antioxidants and antifading agents, or highly air-tight containers should be used.
THE FUTURE OF NANOPARTICULATE APPLICATIONS
FOR TRANSDERMAL DELIVERY
Simply considered, the research can overcome the limitation of nanoparticulate
systems as described in this chapter. As mentioned above, the advantages of
nanoparticulate applications should be maximized and the disadvantages should
be overcome. The ultimate nanoparticulate applications for transdermal delivery
that will be developed in the future can deliver nearly all types of drugs including
macromolecules into target sites by simply spreading onto the skin surface to provide concentrative and long-term release of the drug. To enlarge the applicable
scopes of nanoparticulate for transdermal delivery, it is absolutely necessary to
develop new nanoparticulate systems to enhance the skin absorption and not be
affected by the active ingredients to be encapsulated. Above all, the material diversification of nanoparticulate by the development of material science should be
solved initially, and the analysis methods to analyze the chemical and physical
336
Shim
properties of nanoparticulate should be developed along with the development of
analytical instruments which will enable the above goals.
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Index
Abraxane®
parenteral formulations
characterization and manufacture, 44
clinical trials, 44–45
pharmacokinetics, 44
vs. Taxol®, 44
Absorptive-mediated transcytosis, 266
Acetylsalicylic acid
SEM, 125
Acrylates, 274–276
derivatives, 275
Active tumor targeting
metallic nanoparticle drug-delivery
systems, 149–150
Actively targeted therapies
drug-delivery system nanoengineering, 106
Adjuvant-vector vaccination, 317–323
Adriamycin®, 278
Aerosol flow reactor method, of
multicomponent drug nano- and
microparticle synthesis, 101–126
experimental, 114–117
instrumentation and characterization,
116–117
materials, 114
particle production, 114
precursor solutions, 114
results, 117
nanostructured drug microparticles, 121–125
peptide coating particles, 121–124
reactor drug crystallization, 121–124
polymeric drug nanoparticles, 118
Agitator bead mill, 77
Albumin, 272
nanospheres, 16
Alendronate
liposomal, 240
Alginates, 14, 52
Alkylcyanoacrylates
monomer interfacial polymerization, 157
Alumina
nanospheres, 18–19
AmBisome®, 72
2-Aminochromone U-86893 (U86)
local delivery for restenosis, 246
Amorphous materials
electrospinning, 63–64
Amorphous model, 219
Amphiphilic block copolymers, 13
Amphotericin
liposomal, 91
Anesthetics, 40
parenteral formulations, 40–45
Angiogenic tumor blood vessels
tumor-targeting metallic nanoparticle
drug-delivery systems, 149–150
Antibody-coated nanospheres, 22
Anticancer parenteral formulations,
40–45
Anti-Flk antibody-coated 90Y
nanoparticles, 169
Antifungal parenteral formulations, 38–39
clinical study, 39–40
preclinical studies
with drug-resistant fungal
stain, 39
with drug-susceptible
fungal stain, 39
Antigen-presenting cells (APC), 318
Anti-HER2 antibody-targeted gold/silicon
nanoparticles, 169
Antimicrobial nanoemulsions, 181
Antitranspirants, 227
APC. See Antigen-presenting cells (APC)
Apolipoprotein E, 75
Arecaidine propargyl ester (APE), 275
Arterial localization
doxorubicin, 248
Balloon injury
carotid artery, 244
BBB. See Blood-brain barrier (BBB)
Beclomethasone dipropionate powders
production, 121, 122
SEM, 125
Bee venom peptide, 313
339
340
Berner type low-pressure impactor (BLPI), 115
Bioconjugate
nanoparticle-aptamer, 169
Biodegradable polylactide nanospheres, 18
Bisphosphonate, 250–252
Blood-brain barrier (BBB), 107, 197, 281–283
pathologic alterations, 283
BLPI. See Berner type low-pressure
impactor (BLPI)
Brain compartment, 284–286
Brain endothelial cells, 238–287
Brain tumors
BBB, 283
iron oxide nanoparticles, 180
Caelyx®, 284
Calcium carbonate nanoparticles, 14
Calcium oxide powder
antimicrobial properties, 180
Camptothecin, 22
Cancer. See Oncology
Carbohydrates
cells, 161
Carbon nanotubes, 20, 169
Cardiovascular disease
nanobiotechnology, 180
Carotid artery
balloon injury, 244
Carrageenans, 274
gelatin, 274
Cationic liposomes, 240–242
formulation strategies, 91–94
CD. See Cyclodextrins (CD)
Cells
chemical composition, 160–162
macromolecules, 161–162
small organic molecules, 161
structure, 159
Central nervous system (CNS), 185–186,
281–284
doxorubicin, 284
nanobiotechnology, 180
Ceramic nanoparticles, 169
CFC. See Chlorofluorocarbon (CFC)
Chemical vapor deposition (CVD), 164
Chitosan, 14, 52, 272
Chitosan-based nanoparticulate drug-delivery
systems, 14
Chitosan nanoparticles
gene delivery, 299
Chlorofluorocarbon (CFC)
Montreal Protocol, 76
Index
Cholesteryl myristate (CM), 130–131
supercooled smectic nanoparticles
properties and phase behavior, 130–132
Clinical nanodiagnostics, 178
CLIO paramagnetic nanoparticles, 169
Clodronate
liposomal, 240
Clofazimine nanosuspension
vs. liposomal form, 75
liposomal nanosuspension, 75
Clotrimazole
release profiles, 223
CM. See Cholesteryl myristate (CM)
CMC. See Critical micelle concentration (CMC)
CNS. See Central nervous system (CNS)
Collagen, 14
Colonic targeting
NPDDS, 305, 306–312
Computer-controlled molecular tools, 181
Controlled-release systems
drug-delivery system nanoengineering, 99–102
Cosmetics
lipid nanoparticles, 213–229
NLC, 214, 229
Creams
NLC, 229
Critical micelle concentration (CMC), 9
Cryopreparation method
DNA microencapsulation, 293
CVD. See Chemical vapor deposition (CVD)
Cyclodextrins (CD), 75, 103
Cyclophosphamide-loaded
polybutylcyanoacrylate, 162
Cyclosporine A, 273
HPC, 12
Cypher, 236
CYT-6091, 151–154
Cytarabine
liposomal, 91
Cytoplasm, 159–160
Cytoskeleton, 160
Cytosol, 160
Dalargin, 195
Dantrolene
malignant hyperthermia, 41, 42
Dative binding, 142
Daunorubicin
liposomal, 91
DC. See Dendritic cells (DC)
Dendrimer-based drug-delivery systems, 14
lymphatic targeting, 254
341
Index
Dendrimers, 167–169, 199–201
drug discovery, 175–177
Dendritic cells (DC), 318
Deodorants, 227
Dermal lipid nanoparticles, 213–229
Dermal nanosuspensions, 76
Dexamethasone, 189, 199, 278
local delivery for restenosis, 243–245
Dextran nanospheres
magnetic plain, 19
Dextrans, 14
Differential scanning calorimetry (DSC), 56
Dioleoyl phosphatidylethanolamine
(DOPE), 92, 93
Diphtheria toxin A (DTA), 93
Dip-pen nanolithography, 164–165
Direct compress technology, 85
Dispatch catheter, 207
Dissocubes technology, 79
Dissolution velocity, 71
1,2-Distearoyl-3-phosphatidylethanolamine
(DSPE-PEG), 277
DL Bending Tester, 67–68
DNA
microencapsulation
cryopreparation method, 293
plasmid
PLGA, 291–298
polymer composition, 296
DOCSPER, 241
DOPE. See Dioleoyl phosphatidylethanolamine
(DOPE)
Double emulsion-solvent evaporation
method, 244
Doxil®, 284
Doxorubicin, 198
CNS, 282
liposomal, 91
local delivery for arterial localization, 247
LTSL, 101–102
Drug administration
common routes, 164
Drug-delivery system nanoengineering,
99–107
actively targeted therapies, 106–107
activity duration enhancement, 103–107
controlled- and triggered-release systems,
99–103
drug structural stability, 103–106
formulation processing strategies, 104
future directions, 107
passive tissue targeting, 106
[Drug-delivery system nanoengineering]
polymeric strategies for drug action
prolongation, 104–106
sustained and triggered releases
lipid-based, 101
physicochemical properties, 102–103
polymeric, 99–104
targeted therapies, 106–107
transport across tight endothelial junctions, 107
Drug-eluting stents, 235
Drug nanocrystals, 71–85
clinical applications,
combination technologies, 81–82
defined, 72
drug discovery, 175
final formulations, 84
high-pressure homogenization, 79–84
media milling processes, 76–78
microfluidizer technology, 79
milling, 83
nanoedge technology, 81–82
nanopure technology, 81
XP, 82–83
ocular delivery, 76
oral delivery, 74–75
parenteral administration, 75
particle size reduction techniques, 76–84
physicochemical properties, 73–74
piston-gap homogenization in water, 79–80
precipitation, 78–79
pulmonary drug delivery, 76
SEM, 85
Drug powders
nanoscale
antimicrobial properties, 180
production
micronsized, 113, 116
nano, 116, 117
DSC. See Differential scanning calorimetry (DSC)
DSPE-PEG. See 1,2-Distearoyl-3phosphatidylethanolamine (DSPE-PEG)
DTA. See Diphtheria toxin A (DTA)
Efflux pumps
protection, 59
Electrospinning, 61
amorphous materials, 64
Electrospun fibers
sustained-release
nanofiber-based drug delivery, 65
Electrospun nanofibers, 64–65
drug-delivery systems, 23
342
Emulsification-diffusion technique, 217
Emulsification-evaporation method, 217
Emulsion droplets
vs. lipid nanoparticles, 225
Endoplasmic reticulum, 160
Endosomolytic liposomes
formulation strategies, 91–92
Endothelial junctions, tight
transport across, 107
Enhanced permeability and retention (EPR)
effect, 167
Envelope-type nanodevice, 8
Enzymes, 196–197
Epidermal permeability barrier (EPB), 332
EPR. See Enhanced permeability and
retention (EPR) effect
Ethium bromide displacement assay, 201
Etomidate
supercooled smectic nanoparticles, 134
Eudragit® L100 particles, 114, 116
DSC thermograms, 120
physical properties, 121
Evaporative precipitation, 111
Face powders, 227
Facial makeup products, 227
Fibroblast growth factor (FGF), 235
5-Fluorouracil, 277
Fullerene molecules
drug discovery, 173–176
Fusigenic liposome hemagglutinating virus of
Japan-liposomes (Virosome), 241–242
Fusogenic liposomes
formulation strategies, 91–92
GALA, 92
GALT. See Gut associated with lymphatic
tissue (GALT)
Ganciclovir, 272
Gastrointestinal applications
NPDDS, 305–318
Gelatin, 14, 52
carrageenan, 272
nanospheres, 16
Gels
NLC, 214
Gene delivery
chitosan nanoparticles, 299
nanoparticles
formulation characteristics, 291–299
PACA, 300
polymeric nanoparticles
schematic representation, 292
Index
Gene expression
nanoparticle-mediated
formulation factors, 294
Gene therapy
NPDDS, 188–190
Gene transfection
NP, 294
Geometric number mean diameter
(GNMD), 116–123
Geometric standard deviation (GSD), 116–121
GFP. See Green fluorescent protein (GFP)
Gliadin, 53
Glial cells, 179
GNMD. See Geometric number mean
diameter (GNMD)
Gold nanoparticles
anti-HER2 antibody-targeted, 169
drug discovery, 175
molecular diagnostics, 174
tumor-targeting metallic nanoparticle
drug-delivery systems, 149–152
Gold nanoshells, 169
Gold nanospheres, 19
Golgi bodies, 160
Green fluorescent protein (GFP), 146, 147
GSD. See Geometric standard deviation (GSD)
Gut associated with lymphatic tissue
(GALT), 305
Hair dressings, 227
Heparin, 308
High-pressure homogenization, 81, 111,
216, 224
drug nanocrystals, 76–78
High-shear forces, 111
HPbetaCD. See Hydroxypropyl-betacyclodextrin (HPbetaCD)
HPC. See Hydroxypropyl cellulose (HPC)
Human immunodeficiency virus (HIV), 18
Human serum albumin (HSA), 9
Hydrocarbon ammonium compound
antimicrobial properties, 180
Hydrogel matrices, 14
Hydrogen-based nanoparticulate
drug-delivery systems, 12
Hydrophilic HCL salt drugs
flux enhancement, 103
Hydrophobic binding, 151
Hydrosol technology, 78
Hydrothermal process, 164
Hydroxypropyl-beta-cyclodextrin
(HPbetaCD)
thermogravimetric curves, 105
343
Index
Hydroxypropyl cellulose (HPC)
cyclosporine A, 12
Hyperthermia
malignant
dantrolene, 41
Ibuprofen
supercooled smectic nanoparticles, 135–136
ICAM-1. See Intracellular adhesion
molecule-1 (ICAM-1)
Idarubicin, 312–313
IDD-PTM, 79
IFP. See Intestinal fluid pressure (IFP)
IL-6. See Interleukin-6 (IL-6)
ILP. See Isolated limb protocol (ILP)
Immunization
mucosal, 319
nasal, 320–322
oral, 319–320
parenteral, 318–319
Immunoadjuvants, 317–318
Immunoliposomes, 192
Imperfect crystal model, 219
Infections
nanobiotechnology, 179–181
Injectable nanosuspension
monocyte phagocytic system, 40
in vivo distribution, as function
of particle size, 39
Insulin, 167
Interleukin-6 (IL-6), 241
Interpenetrating polymer networks (IPN)
PAA, 200
Intestinal fluid pressure (IFP), 147–148
Intestinal permeability, 71
Intracellular adhesion molecule-1
(ICAM-1), 241
Intramuscular drug administration, 164
Intravenous drug administration, 164
Intravenous parenteral formulations, 41
Ionic binding, 161
IPN. See Interpenetrating polymer
networks (IPN)
Iron nanoparticles
tumor-targeting metallic nanoparticle
drug-delivery systems, 149–150, 151
Iron oxide nanoparticles
brain tumors, 180
Isolated limb protocol (ILP), 153
Isopropylacrylamide
polyacrylic acid, 200
Itraconazole, 43
vs. Sporanox®, 43
Ketoconazole, 224
Ketoprofen, 121
physical properties, 122
Lactose
SEM, 124
Laser pyrolysis, 164
Leucine
SEM, 124
LH-RH-targeted silica-coated lipid
micelles, 169
Lipex extruder, 90
Lipid(s)
cells, 161
drug expulsion, 215
freeze dried nanocapsules, 7
liposome formulation strategies, 91–92
ocular applications, 271–279
Lipid-based colloidal nanodrug-delivery
systems, 6–7
liposomes
drug carriers, 91
formulation strategies, 91–95
Lipid-based drug-delivery system
nanoengineering
for sustained and triggered
releases, 99–103
nanoparticulate, 89–95
liposome preparation and
characterization, 89–91
Lipid carriers. See Nanostructured lipid
carriers (NLC)
Lipid-encapsulated perfluorocarbon
nanoemulsions, 169
Lipid micelles
LH-RH-targeted silica-coated, 169
Lipid nanoparticles
cosmetic, dermal and transdermal
applications, 212–229
vs. emulsion droplets, 225
liposome formulation strategies, 95
Lipofectamine, 241
Lipophilic drugs
liposome preparation, 89–90
Liposomal alendronate, 252
Liposomal amphotericin, 91
Liposomal clodronate, 252
Liposomal cytarabine, 91
Liposomal daunorubicin, 91
Liposomal doxorubicin, 90
Liposomal nanosuspension
clofazimine nanosuspension, 75
Liposomal verteporfin, 91
344
Liposome(s), 72, 168, 169
cationic, 240–242
cryo-TEM microscopy, 252
discovery, 163
formulation strategies, 91–95
cationic liposomes, 93–94
fusogenic and endosomolytic
liposomes, 92–93
lipid composition for increased stability
in vivo, 91–92
lipid nanoparticles, 94–95
pH-sensitive liposomes, 92
sterically stabilized liposomes, 92
targeted liposomes, 94
temperature-sensitive liposomes, 93
lipid-based colloidal nanodrug-delivery
systems, 6–7
lipid-based nanoparticulate drug-delivery
systems, 89
preparation and characterization, 89–90
lipophilic drugs, 90
low temperature thermosensitive
doxorubicin, 101, 102
nanocarriers, for restenosis, 23
purification, 90
restenosis
local drug delivery, 237
Liposome-based drug-delivery systems
lymphatic targeting, 305–306
Lipsticks, 227
L-leucine, 112–114
production, 114
SEM, 117
Low temperature thermosensitive
liposomes (LTSL)
doxorubicin, 102
Luciferase protein
HEK-293 cell line, 245
Lymphatic targeting
drug-delivery systems
dendrimer-based, 307
liposome-based, 308–309
NPDDS, 305
Lysosomes, 160
Macrophages, 249–253, 318
Magnesium oxide powder
antimicrobial properties, 180
Magnetic nanoparticles, 169
gene delivery, 300–301
molecular diagnostics, 173
Magnetic plain dextran nanospheres, 19
Index
Malignant hyperthermia
dantrolene, 41–42
Mannitol
malignant hyperthermia, 41
Mass median aerodynamic diameter
(MMAD), 17
Mechanical integrity testing
nanofiber-based drug delivery, 61–69
Media milling processes
drug nanocrystals, 75–78
Megace®, 84
Melamine nanospheres, 18
Metabolomics, 167
Metallic nanoparticle drug-delivery
systems
tumor-targeting, 150–154
Methylmethacrylate (MMA), 275
Micelles, 169
drug carriers
nanoparticulate polymeric, 9–10
LH-RH-targeted silica-coated lipid, 169
polymeric
drug solubilization, 10
research trends, 12
Miconazole
supercooled smectic nanoparticles, 135
Microencapsulation
DNA
cryopreparation method, 293
Microfluidics, 5–6
drug nanocrystals, 78–79
Micronization, 72
Micronsized drug powders
production, 113
Microparticles
preparation, 115
Mifepristone, 8
Milling
drug nanocrystals, 83–84
Mitochondria, 160
MLV. See Multilamellar vesicles (MLV)
MMA. See Methylmethacrylate (MMA)
MMAD. See Mass median aerodynamic
diameter (MMAD)
Molecular condensation, 165
Molecular diagnostics
nanobiotechnology, 173–175
Molecular nanotechnology, 162
Monoclonal antibodies, 147
Monocytes, 249–251
phagocytic system
injectable nanosuspension, 38
345
Index
Montreal Protocol
CFC, 76
Mucoadhesive, 197–198
Mucoadhesive nanoparticulate
drug-delivery systems, 197–198
Mucosal immunization, 319
Multifunctional nanoparticle
schematic representation, 142
Multilamellar vesicles (MLV), 89
Multiple model, 219
Nanobiotechnology
cardiovascular disease, 180
central nervous system disorders, 179–180
drug delivery, 176–177
drug discovery, 175–176
future prospects, 181–182
infections, 180–181
molecular diagnostics, 173–174
oncology, 179
personalized medicine, 181
therapeutic applications, 168–170
Nanobodies
drug discovery, 176
Nanocapsules, 169, 236
lipid
freeze dried, 7
polymeric
drug carriers, 17
Nanocarriers
restenosis, 235–256
Nanocontainer technology, 23
Nanocrystal. See also Drug nanocrystals
technology, 76–78
Nanocrystallites, 22
Nanodevices
envelope-type, 9
fabrication, 163–166
synthetic, 177
Nanodiagnostics, 173, 177–178
clinical, 177
Nano drug powders
production, 113–116
Nanoedge technology
drug nanocrystals, 81–82
Nanoemulsions
antimicrobial, 181
lipid-encapsulated perfluorocarbon, 169
Nanoendoscopy, 177–178
Nanoengineered prosthetics, 21
Nanofiber-based drug delivery, 61–66
automated fiber sizing, 66–67
[Nanofiber-based drug delivery]
dissolution enhancement for
immediate-release dosage, 63–64
large-scale manufacturing, 65–69
mechanical integrity testing, 67–69
sustained-release electrospun
fibers, 64–65
Nanofibers
electrospun, 61–63
Nanohybrids, 22–23
Nanolithography
dip-pen, 164–165
Nanomaterials
characteristics, 1–3
manufacture, 3–6
measurement, 1–3
Nanomedicine, 173–182, 177–182
twenty-first century, 178
Nanometer
length scale, 2
Nanometrology, 1–3
Nanoparticle(s). See also Solid-lipid
nanoparticles (SLN); Supercooled
smectic nanoparticles
affecting biological performance, 56–58
analytical characterization, 4
biocompatibility, 58
and biodegradability, 55
biological research, 169
bottom-up manufacturing, 4–5
calcium carbonate, 14
ceramic, 169, 169
characterization, 54–58
CLIO paramagnetic, 169
commercially available, 18–20
copolymer ratio, 55
crystallinity, 55
dissolution profile, 57
drug discovery, 175–176
drug loading and loading efficiency, 57
drug-polymer interactions, 56
future directions, 24
gene delivery
formulation characteristics, 291–301
gold
drug discovery, 175–176
molecular diagnostics, 173
hydrophobicity, 55
internalization into cells, 58
iron oxide
brain tumors, 180
lipid vs. emulsion droplets, 225
346
[Nanoparticle(s)]
magnetic, 169
gene delivery, 299–301
molecular diagnostics, 173
molecular weight, 54–55
natural biodegradable polymers, 52–53
neurosurgery, 190
nonbiodegradable polymers, 54
particle size, 56–57
physical properties, 54–56, 119
poly(beta-amino ester)-based gene
delivery, 299–300
polybutylcyanoacrylate, 288
polyester polysaccharide, 16
polylactide-co-glycolide
scanning electron microscopic images, 57
polymeric carriers, 51–53
polyvinyl caprolactone, 54
preparation, 115
protective
against pathogens, 21
rhodamine-containing, 238
route of administration, 56
silicon
anti-HER2 antibody-targeted, 169
solubility, 55
structure, 186
surface modification, 58
synthetic biodegradable polymers, 53–54
tagging, 20
targeted delivery, 59
therapeutic applications, 167–168
thermosensitive, 54
thiamine-targeted, 169
thiomer, 21
top-down manufacturing, 5–6
zeta potential, 57
Nanoparticle-aptamer bioconjugate, 169
Nanoparticle drug delivery systems
(NPDDS), 11–12, 185–193
absorptive-mediated transcytosis, 286
advantages, 186–187
alkylcyanoacrylate monomer interfacial
polymerization, 187
applications, 188–202
brain
compartment, 285–286
endothelial cells, 286–287
brain-targeted delivery, 284–285
cancer, 195–196
central nervous system, 281–288
colonic targeting, 306, 313
Index
[Nanoparticle drug delivery systems (NPDDS)]
drugs, 189–190
enzymes, 196–197
formulation applications, 190
gastrointestinal applications, 305–314
advantages, 313–314
evaluation, 314
future perspectives, 314
gene therapy, 199–202
hydrogen-based, 12–13
lymphatic targeting, 305–306
development, 306–308
manufacturing techniques, 187–188
mucoadhesive, 197–198
ocular applications, 193, 271–279
pharmaceutical applications, 185–203
polymer interfacial deposition, 187–188
proteins and peptides, 188–193
pulmonary treatment, 193–195
receptor-mediated transcytosis, 286–287
risks, 202–203
vascular thrombosis, 198–199
Nanoparticle-mediated gene expression
formulation factors, 294
Nanoparticulate
defined, 327–328
drug carriers
polymeric micelle, 9–12
protamine-based, 14–15, 197
drug-delivery systems
polymer-based, 12–18
Nanoprecipitation, 111
Nanoprobes
plaque, 180
Nanopure technology, 81–82
drug nanocrystals, 81
XP, 82–83
Nanoscale drug powders
antimicrobial properties, 190
Nanoscience, 1
Nanoshells
gold, 169
oncology, 179
Nanospheres, 169
albumin, 16–17
alumina, 19–20
antibody-coated, 22
biodegradable polylactide, 18–19
gelatin, 16–17
gold, 19
magnetic plain dextran, 19
melamine, 18
347
Index
[Nanospheres]
nanocarriers for restenosis, 237
plain polymethyl methacrylate, 18–19
polystyrene, 18
silica, 19
silver, 19
Nanostructured drug microparticles
aerosol flow reactor method
multicomponent drug synthesis, 113–114
peptide coating particles, 122–125
reactor drug crystallization, 121–122
Nanostructured lipid carriers (NLC), 214–216
chemical protection, 224
controlled-release properties, 222–223
cosmetic products, 213, 215, 226–229
creams, 221–222
entrapment efficiency, 222–223
gels, 220–221
hydration, 224–226
loading capacity, 222–223
morphology and structure, 218–219
occlusive effects, 224–226
penetration enhancement, 226
physical stability, 221–222
production, 220–221
skin effects, 221–226
transdermal products, 219–221, 227–229
types, 218–219
Nanostructured materials, 21
Nanostructured monoliths, 22
Nanosuspensions
clofazimine, 75
vs. liposomal form, 75
dermal, 76
injectable monocyte phagocytic system, 40
liposomal, 75
parenteral delivery, 33–46
combination of component technologies, 34
combined precipitation/homogenization, 36
formulation approaches, 34–36
formulation selection, 36
historical introduction, 33–34
homogenization, 35
injectable nanosuspension safety, 38–40
intrathecal delivery, 46
manufacturing methods, 34–36
molecular determinants of particle size,
36–37
parenteral formulation applications, 40–45
pharmacokinetic profiles, 37–38
pharmacokinetics, 44
precipitation, 35
[Nanosuspensions
parenteral delivery]
predecessor technology, 33
prediction of stability, 36–38
in vitro dissolution, 37
Taxol®, 34, 44
Nanotechnology, 1
development, 162–163
diverse and emerging trends, 20–23
Nanotherapeutic applications
biological requirements, 159–170
Nanotherapeutic devices, 166–168
Nanotubes
carbon, 20, 169
cellular manipulation, 21
Nanowires
silicon-based, 169
Naproxen
physical properties, 118
Nasal immunization, 319–322
Natural and synthetic polymers, 51–59
in vitro cell culture, 58–59
New chemical entities (NCE), 77
Nitric oxide eluting stents
nanofibers, 180
NK911, 180
NLC. See Nanostructured lipid carriers (NLC)
Nonviral vectors, 201–202
Normalization window, 148
NP
gene transfection, 293
poly(D,L-lactic acid)-based formulation
characteristics, 291–298
restenosis
local drug delivery, 237
nanocarriers, 237
polymeric gene delivery, 242–243
NPDDS. See Nanoparticle drug delivery
systems (NPDDS)
Nucleic acids, 161–162
Octylglucoside, 90
Ocular applications
drug nanocrystals, 76
lipids, 278
NPDDS, 193–194, 271–279
Oil-in-water emulsion, 111
Oligodeoxyribonucleotides (ODN), 93
Oncology, 197–198
nanobiotechnology, 176
photodynamic therapy, 14
Ophthalmic drugs, 193
348
Opsinization, 142
Optical tweezers, 3
Oral delivery
drug nanocrystals, 74
Oral drug administration, 164
Oral immunization, 319
Oswald ripening, 78
PAA. See Polyacrylic acid (PAA)
PACA. See Polyalkyl-cyanoacrylates (PACA)
Paclitaxel, 249
local delivery for arterial localization,
248–249
parenteral formulations, 40–45
characterization and manufacture, 44
supersaturable formulation, 309
Palm oil, 101
Parenteral Abraxane®
characterization and manufacture, 44
clinical trials, 44–45
pharmacokinetics, 44
Parenteral formulations
anticancer, 44–45
antifungal, 42–43
clinical study, 43–44
drug nanocrystals, 74
intravenous, 41–42
preclinical studies
drug-resistant fungal stain, 42–43
drug-susceptible fungal stain, 42
regional anesthetics, 39–41
Parenteral immunization, 318–319
Parenteral paclitaxel, 318–319
characterization and manufacture, 44
Parenteral Taxol®
clinical trials, 44–45
Particle formation
experimental parameters, 118
Passive tissue targeting
drug-delivery system nanoengineering, 106
Passive tumor targeting
metallic NDDS, 149
PBCA. See Polybutylcyanoacrylate (PBCA)
PC. See Phosphatidylcholines (PC)
P-channel MOSFET
fabrication, 165
PCI. See Percutaneous coronary intervention
(PCI)
PCNA. See Proliferating cell nuclear
antigen (PCNA)
PDGF. See Platelet-derived growth factor
(PDGF)
Index
Pectin, 272, 273
PECVD. See Plasma-enhanced vapor
deposition (PECVD)
PEG. See Polyethylene glycol (PEG)
PEI. See Polyethylenimine (PEI)
Penetration
defined, 327–328
PEO. See Polyethylene oxide (PEO)
Peptide coating particles
nanostructured drug microparticles,
121–122
Peptides
NPDDS, 188–193
Percutaneous adsorption, 328
Percutaneous coronary intervention (PCI), 235
restenosis, 180
Perfume, 228
Permeation
defined, 328
Peroxisomes, 160
Personalized medicine
nanobiotechnology, 181
Peyer’s patches, 305, 306, 320
Phagocytic system
monocytes
injectable nanosuspension, 40
Phase-inversion-based technique, 217
Phase separation, 111
Phosphatidylcholines (PC), 92
Phosphorotioated oligodeoxynucleotides
(PT-ODN), 239–240, 244
Photodynamic therapy
cancer, 14
Photosensitizers, 14
Photothermal tumor ablation, 179
PillCam™, 178
Piston-gap homogenization in water
drug nanocrystals, 79
Plain polymethyl methacrylate
nanospheres, 18–19
Plasma cell membrane, 159
Plasma-enhanced vapor deposition
(PECVD), 166
Plasmid DNA
PLGA, 297
Platelet-derived growth factor (PDGF),
235, 243–244, 263
Platelets, 144–145
PLG. See Polyglycolide (PLG)
PLGA. See Polylactide-co-glycolide (PLGA)
PLGA-PEG. See Polylactide-co-glycolide
polyethylene glycol (PLGA-PEG)
349
Index
PMA. See Polymethacrylate (PMA)
PMMA. See Polymethyl methacrylate
(PMMA)
Poly(D,L-lactic acid)
formulation characteristics, 277–278
Poly(D,L-lactide-co-glycolide), 277–278
Poly(epsilon-caprolacton), 276
Polyacrylic acid (PAA), 275
IPN, 200
isopropylacrylamide, 200
Polyalkyl-cyanoacrylates (PACA),
54, 274–277
gene delivery, 299
Polyanhydrides, 53
Poly(beta-amino ester)-based nanoparticles
gene delivery, 299
Poly(D,L-lactic acid)-based NP
formulation characteristics, 291–297
Polybutylcyanoacrylate (PBCA)
cyclophosphamide-loaded, 193
nanoparticles, 288
Polycarbonate membrane, 90
Poly-ε-caprolactones, 53
Polyester polysaccharide nanoparticles, 16
Polyethylene glycol (PEG), 143
Polyethylene oxide (PEO), 295
Polyethylenimine (PEI), 298
Polyglycolide (PLG), 237
Polylactide, 53
Polylactide-co-glycolide (PLGA),
53, 186–187, 297, 307
molecular weight, 295
nanoparticles, SEM, 57
plasmid DNA, 297
Polylactide-co-glycolide polyethylene glycol
(PLGA-PEG)
formulation, 196
Polymer(s)
natural origin, 272–274
Polymer-based nanoparticulate drug-delivery
systems, 12–18
Polymeric drug nanoparticles
aerosol flow reactor method, 114–116
Polymeric growth inhibitors, 78
Polymeric micelles
drug solubilization, 10–11
research trends, 12
Polymeric nanocapsular-based drug-delivery
systems
lymphatic targeting, 306–312
Polymeric nanocapsules
drug carriers, 17
Polymeric nanoparticulate-based
drug-delivery systems
aerosol flow reactor method, 113–114
gene delivery
schematic representation, 269
lymphatic targeting, 305
preparation, 56
Polymethacrylate (PMA), 53
Polymethyl methacrylate (PMMA), 53
Polystyrene nanospheres, 18
Polyvinyl caprolactone nanoparticles, 53
Poorly soluble drugs
solubilization, 71–85
Precipitation
drug nanocrystals, 78, 83–84
evaporative, 111
Progesterone
supercooled smectic nanoparticles, 135–136
Proliferating cell nuclear antigen (PCNA), 239
Propofol, 17
Prosthetics
nanoengineered, 21
Protamine-based nanoparticulate
drug carriers, 14–15
Protective nanoparticles
against pathogens, 21
Proteins
cells, 162
NPDDS, 188–193
Proticles, 14–15, 197
PT-ODN. See Phosphorotioated
oligodeoxynucleotides (PT-ODN)
Pullulan, 53, 202
Pulmonary drug delivery, 193–195
drug nanocrystals, 96
Quantum dots, 169
drug discovery, 175–176
molecular diagnostics, 173–174
Rapamune®, 78
Rapamycin, 236
Reactor drug crystallization
nanostructured drug microparticles, 121–122
Receptor-mediated transcytosis, 286–287
Regional anesthetics
parenteral formulations, 40–41
RES. See Reticuloendothelial system (RES)
Restenosis, 235
dexamethasone, 245–246
gene therapy, 239–240
liposomal gene delivery, 240–242
350
[Restenosis]
local delivery
formulation, 246
local drug delivery, 237–238, 243–246
nanocarriers, 235–256
polymeric NP gene delivery, 242–243
systemic drug delivery, 248–255
arterial localization, 248–249
for systemic target, 249–251
Reticuloendothelial system (RES), 9
Retinoic acid, 278
Rhodamine, 253
nanoparticles containing, 238
RH-targeted silica-coated lipid micelles, 169
Ribosomes, 159
Rolipram, 313
Rough endoplasmic reticulum, 159–160
Salting-out, 111
Scanning electron microscopy (SEM), 3
Scanning probe microscopy (SPM), 3
Self-diagnostics, 181
Self-(micro) emulsifying drug-delivery systems
lymphatic targeting, 309–312
SEM. See Scanning electron microscopy (SEM)
Shaving aids, 227
Silica-coated lipid micelles
LH-RH-targeted, 169
Silica nanospheres, 19
Silicon-based nanowires, 169
Silicone nanopore-membrane-based
drug-delivery system, 15
Silicon nanoparticles
anti-HER2 antibody-targeted, 169
Silver nanospheres, 19
Silver powder
antimicrobial properties, 181
Single-beam gradient trap, 3
Single-molecule detection, 21
SiRNA, 93
Skin
aging, 227
permeation test, 327, 332, 334
SLN. See Solid-lipid nanoparticles (SLN)
Small-scale cyclone, 116
SMC. See Smooth muscle cells (SMC)
Smectic nanoparticles
preparation methods, 132
Smooth endoplasmic reticulum, 160
Smooth muscle cells (SMC), 235
Sodium chloride
SEM, 125
Index
Sodium leucine
SEM, 125
Sol-gel process, 164–165
Solid-lipid nanoparticles (SLN), 5, 54, 188, 214
brain, 287–288
chemical protection, 224
controlled-release properties, 222–224
cosmetic products, 219–221
creams, 221
entrapment efficiency, 222–224
gels, 220–221
hydration, 224–226
loading capacity, 222–224
lymphatic targeting, 308
morphology and structure, 218–219
occlusive effects, 224–226
ocular applications, 278
penetration enhancement, 225
physical stability, 221–222
production, 216–218
research trends, 7–9
skin effects, 221–229
transdermal products, 219–220
types, 219
Solvent displacement method, 217
SPIO. See Superparamagnetic iron oxide (SPIO)
SPM. See Scanning probe microscopy (SPM);
Sulfopropylmethacrylate (SPM)
Sporanox®, 42–44, 72
vs. itraconazole, 43
parenteral administration, 75
Spray-drying, 84, 112–113
Starch, 14
Starch acetate, 272
Stem-cell-based therapies
nanobiotechnology, 178
Stents
drug-eluting, 235
Sterically stabilized liposomes
formulation strategies, 92
Stratum corneum adsorption, 328
Subcutaneous drug administration, 164
Sulfopropylmethacrylate (SPM), 275
Supercooled smectic nanoparticles, 129–138
CM nanoparticles
properties and phase behavior, 133–137
ultrastructure, 136, 138
etomidate, 135
ibuprofen, 135
matrix composition, 133–135
model drugs, 135–136
preparation methods, 132
351
Index
[Supercooled smectic nanoparticles]
stabilizer system, 132–133
thermotropic mesophases, 130
Superparamagnetic iron oxide (SPIO), 150
Sustained-released systems
drug-delivery system nanoengineering,
99–101, 102–103
modulation of drug environment, 102–103
physicochemical properties, 102–103
polymeric drug-delivery system, 99–101
Sustained-release electrospun fibers
nanofiber-based drug delivery, 64–65
Synthetic nanodevices, 177
TAM. See Tumor-associated macrophages (TAM)
Targeted drug delivery, 176
Targeted liposomes
formulation strategies, 94
Taxol®
vs. Abraxane®, 44
nanosuspensions, 33–34
parenteral formulations
clinical trials, 44–45
TAXUS®, 236
TEM. See Transmission electron microscopy
(TEM)
Temperature-sensitive liposomes
formulation strategies, 93
TGA. See Thermogravimetric analysis (TGA)
Thermal ablation
tumor-targeting metallic nanoparticle
drug-delivery systems, 152–153
Thermogravimetric analysis (TGA), 56
Thermosensitive nanoparticles, 54
Thermotropic calamitic mesophases
structure, 130
Thiamine-targeted nanoparticles, 169
Thiomer nanoparticles, 21
Tissue factor, 237
TNF. See Tumor necrosis factor (TNF)
TNT system, 169
Tobramycin, 310, 312–313
SLN, 278
Transcytosis
receptor-mediated, 286–287
Transdermal applications, 327–335
future, 335
limitations, 334–335
Transfersomes, 169
Transmission electron microscopy (TEM), 3
Triacylglycerols
three-dimensional structure, 215
TriCor®, 78
Triggered-released systems
drug-delivery system nanoengineering,
99–103
modulation of drug environment and
physicochemical properties
drug-delivery system nanoengineering,
102–103
polymeric drug-delivery system
drug-delivery system nanoengineering,
99–100
Tumor(s)
blood vessels, 145
photothermal ablation, 179
Tumor-associated macrophages (TAM), 145
Tumor necrosis factor (TNF), 148, 241
Tumor-targeting metallic nanoparticle
drug-delivery systems, 141
active vs. passive tumor targeting, 149
angiogenic tumor blood vessels, 146–147
chemistry, 151
cytokine-mediated reduction of interstitial
fluid pressure, 150
gold nanoparticles, 150
studies, 151–152
intratumor barriers, 149
iron nanoparticles, 150
studies, 152
multifunctional nanoparticles, 153–154
RES clearance, 142–144
safety, 150–151
thermal ablation, 152
tumor angiogenesis
and interstitial fluid pressure, 148
and vascularization, 144–145
tumor models, 149–150
vascular endothelial growth factor, 145–146
Tweezers
optical, 3
Tyrphostins
local delivery for restenosis, 243–245
U86. See 2-Aminochromone U-86893 (U86)
Ultra-small superparamagnetic iron oxide
(USPIO), 150
Ultraviolet blockers, 227, 228
Vaccination
adjuvant-vector, 317–323
Vascular endothelial growth factor (VEGF)
tumor-targeting metallic nanoparticle
drug-delivery systems, 145–146, 150
352
Vascular smooth muscle cells
(VSMC), 239
Vascular thrombosis, 198–199
Vasculogenesis, 146
Vasoactive intestinal peptide (VIP), 194
Vectors
nonviral, 201
VEGF. See Vascular endothelial growth factor
(VEGF)
Verteprofin
liposomal, 90
Vesicular vacuolar organelles (VVO), 145
VIP. See Vasoactive intestinal peptide (VIP)
Index
Virosome, 241–242
von Willebrand factor, 144
VVO. See Vesicular vacuolar organelles (VVO)
Water-in-oil-in-water double emulsion, 292
Water-soluble amphiphilic nanocarriers
freeze drying, 11
Wet milling, 111
Whole skin penetration, 332–334
Zeldox®, 72
Zeta potential
nanoparticles, 57
FIGURE 1.1 Length scale showing the nanometer in context. The length scale of interest for nanoscience and nanotechnologies is from
100 nm down to the atomic scale approximately 0.2 nm. Source: From Ref. 1.
FIGURE 3.2
Schematic diagram of specific receptor targeting of nanospheres using ligands.
FIGURE 4.6
Results of automated fiber size analysis.
FIGURE 5.5 Two nanosuspensions composed similarly, produced
by high-pressure homogenization,
conventional method (left ) versus
translucent nanosuspension (particle size well below 100 nm) resulting
from H96 technology. The red laser
beam is reflected by the tiny nanocrystals. Source : From Ref. 40.
FIGURE 11.1 Structure of a typical cell.
FIGURE 13.1 Different types of
nanoparticulate structures.
FIGURE 14.1 Basic types of
solid nanoparticles and nanostructured lipid center. Abbrevations:
NLC, nanostructured lipid carrier;
SLN, solid–lipid nanoparticle.
Source: From Ref. 44.
FIGURE 15.1 Confocal images of balloon-injured rat carotid arteries after intaluminal delivery of
rhodamine solution and rhodamine-containing nanoparticles. The arteries were harvested 90 minutes
(A and D), eight hours (B and E), and one day (C and F) after delivery. Source: From Ref. 108.
FIGURE 15.2 (A) Inhibition of neointimal formation 14 days after balloon injury to rat carotid artery
and intraluminal instillations with 20 µM (1nmole) antisense or scrambled partially phosphorothioated
oligodeoxynucleotides (PT-ODNs) encapsulated in PLGA nanoparticles (total of 50 µL suspension).
Photomicrographs of representative histological sections. Verchoeff’s elastin stain, magnification
×12.5. (B) Fluorescence micrographs of blank and PT-ODN-loaded PLGA nanoparticles. The PTODNs were covalently labeled with FITC prior encapsulation. Note fluorescence signal in over 90%
of the particles. Abbreviations: A, adventitia; L, lwnen; M, media; N, neointima; NP, nanoparticles;
SC-NP, scrambled NP; AS-NP, antisense NP. Source: From Refs. 198 and 201.
FIGURE 15.3 (C) Low- and high-power photomicrographs of hypercholesterolemic rabbit iliac
artery stents at 28 days (Verhoeff staining) of control (I and II) and of animals treated with LA 3 mg/kg
(III and IV) and 6 mg /kg (V and VI). Abbreviation: LA, liposomal alendronate. Source : From Refs. 269
and 292.
FIGURE 18.2 Formulation factors influencing nanoparticle-mediated gene expression.
Abbreviations: NPs, nanoparticles; PE, primary endosomes; RE, recycling endosomes.
FIGURE 20.1 Visualization of
flagellin-coated nanoparticles (red
dots) in the follicle-associated epithelium of Peyer’s patches by fluorescence. Flagellin of S. enteritidis
was used to coat Gantrez nanoparticles (labeled with rhodamine B
isothiocyanate) and, thus, obtaining adjuvant vectors able to spread
within the gut in a similar way than
the whole bacteria. Source: From
Ref. 34.
FIGURE 21.1 Cryo-sectioned images of the fluorescent labeled nanoparticulate with porcine
epidermal skin layer after in vitro skin permeation test. (A) Stratum comeum outer layer adsorption
of 200 nm-sized polysturene nanoparticulate. (B) Percutaneous absorption of 40 nm-sized polystyrene nanoparticulate. (C) Epidermal layer penetration of the modified nanoparticulate.
Abbreviations: EP, epidermis; SC, stratum comeum.
FIGURE 21.2 CLSM images of the nanoparticulate penetration with a guinea pig skin (cryosectioned). Fluorescent-labelled nanoparticulate (green region), nuclei-labeled DAPI (blue region).
(A) 40 nm size of nanoparticles, (B) 130 nm size of nanoparticles (10 pieces of z-direction sectioning
image of cross sectioned tissue are merged). Nanoparticulate: Polycaprolactone-polyethyleneglycol
block copolymer aggregates. Fluorescent: Rubrene. Source: From Ref. 27.
FIGURE 21.3 Images of the liposome nanoparticulate with fluorescent by in vitro permeation test
with a guinea pig skin. (A) RITC saturated solution (B) RITC with the modified liposome. (Red region),
hydrophilic fluorescent dye (RITC); (blue region), DAPI for nuclei; basal layer (white arrows).
FIGURE 21.4 Fluorescent images of in vivo permeation study with Albino Hartley guinea pig.
(A) O/W emulsion formulation. (B) Modified polymeric nanoparticulate.