Nanomaterials for Clinical Applications: Case Studies in Nanomedicines (Micro and Nano Technologies) [1 ed.] 012816705X, 9780128167052

Table of contents :
WARNING!!! DUMMY ENTRY
Nanomaterials for Clinical Applications
Nanomaterials for Clinical Applications
Copyright
Contents
List of contributors
one Solid lipid nanoparticles in dermaceuticals
1.1 General introduction
1.2 Why solid lipid nanoparticles?
1.2.1 Formulation aspects
1.2.2 Physiological aspects
1.3 Evolution of lipidic nanoparticles from solid lipid nanoparticles to nanostructured lipid carriers
1.4 Cosmetic and topical applications of solid lipid nanoparticles
1.5 Skin penetration with solid lipid nanoparticles
1.6 Mechanism of drug penetration with solid lipid nanoparticles
1.7 Incorporation into semisolid vehicle
1.8 Case studies of successful topical delivery with lipidic nanoparticles
1.8.1 Delivery of antimicrobials
1.9 Delivery of agents for other skin diseases
1.9.1 Solid lipid nanoparticles for cosmetic applications
1.10 Conclusions
References
two Cyclodextrin-based drug delivery systems
2.1 Cyclodextrins—structure, physiochemical properties, and toxicological profile
2.2 Cyclodextrin inclusion complexes—formation, stability, and application in drug delivery
2.3 Cyclodextrin-based products in clinical practice
2.3.1 Cyclodextrins as multifunctional excipients in dosage form design
2.3.1.1 Cyclodextrins in parenteral formulations
2.3.1.2 Cyclodextrins in ocular formulations
2.3.1.3 Cyclodextrins in nasal formulations
2.3.1.4 Cyclodextrins in oral formulations
2.3.1.5 Cyclodextrins in oromucosal formulations
2.3.1.6 Cyclodextrins in dermal formulations
2.3.2 Cyclodextrin as novel therapeutically active pharmaceutical ingredients
2.3.2.1 Sugammadex
2.3.3 Treatment of Niemann–Pick disease, type C1 disease with HPβCD
2.3.3.1 Cyclodextrins in control of obesity and hyperlipidemia
References
three Lipid vesicles for (trans)dermal administration
3.1 (Trans)dermal drug-delivery systems
3.1.1 Human skin barrier to xenobiotics
3.2 Lipid vesicles for breaching the skin barrier
3.2.1 Conventional liposomes
3.2.2 Transfersomes
3.2.3 Ethanol-based lipid vesicles
3.2.3.1 Ethosomes
3.2.3.2 Transethosomes
3.2.3.3 Other lipid vesicles
3.3 Liposomal formulation in clinics
3.3.1 Conventional liposomes
3.3.2 Transfersomes
3.3.3 Ethosomes
3.4 Final remarks
References
four Stimuli-responsive nanocarriers for drug delivery
4.1 Introduction
4.2 Types of stimuli
4.2.1 pH-responsive nanosystems
4.2.2 Thermoresponsive nanosystems
4.3 Development of chimeric stimuli-responsive liposomes with incorporated stimuli-responsive polymers
4.3.1 pH-responsive liposomes
4.3.2 Thermoresponsive liposomes
4.4 Thermotropic behavior of stimuli-responsive liposomes
4.4.1 Thermal analysis on pH-responsive liposomes
4.4.2 Thermal analysis of thermoresponsive liposomes
4.5 Physicochemical properties of stimuli-responsive liposomes
4.5.1 Physicochemical characterization of pH-responsive liposomes
4.5.2 Physicochemical characterization of thermoresponsive liposomes
4.6 Development of stimuli-responsive lyotropic liquid crystalline nanosystems
4.6.1 Stimuli-responsive lyotropic liquid crystalline nanosystems using polycation of PDMAEMA
4.6.2 pH-responsive cubosomes by using pH-sensitive polymer
4.6.3 pH-responsive liquid crystalline nanosystems using pH-responsive molecules
4.6.4 Thermoresponsive lipid-based liquid crystalline nanosystems
4.7 Conclusion and future directions
References
five Biodegradable nanomaterials
5.1 Introduction
5.2 Natural polymers
5.2.1 Polysaccharides
5.2.1.1 Chitosan
5.2.1.2 Hyaluronic acid
5.2.1.3 Alginate
5.2.1.4 Starch
5.2.1.5 Cellulose and cellulose derivatives
5.2.2 Proteins
5.2.2.1 Collagen
5.2.2.2 Gelatin
5.2.2.3 Albumin
5.2.3 Biopolymers of bacterial origin
5.2.3.1 Polyhydroxyalcanoates
5.2.3.2 Poly(γ-glutamic acid)
5.3 Synthetic polymers
5.3.1 Aliphatic polyesters
5.3.1.1 Polyglycolic acid
5.3.1.2 Polylactic acid
5.3.1.3 Polylactic-co-glycolic acid
5.3.1.4 Poly-ε-caprolactone
5.3.2 Poly-(orthoesters)
5.3.3 Polyanhydrides
5.3.4 Poly(alkyl cyanoacrylates)
5.3.5 Synthetic poly(amino acids)
5.3.6 Inorganic biodegradable polymers
5.3.6.1 Polyphosphazenes
5.3.6.2 Polyphosphates
5.4 Polymeric nanoparticles
5.4.1 Introduction
5.4.2 Properties—advantages of polymeric nanoparticles
5.4.3 Polymeric nanoparticles preparation
5.4.3.1 Emulsion—solvent evaporation
5.4.3.2 Coacervation
5.4.3.3 Nanoprecipitation, coprecipitation, and dialysis
5.4.3.4 Ionic gelation
5.4.3.5 Spray drying
5.4.4 Drug release mechanisms
5.4.5 Targeting
5.4.5.1 Passive targeting
5.4.5.2 Active targeting
5.4.5.3 Tumor targeting
5.5 Clinical applications of biodegradable nanoparticles
5.5.1 Introduction
5.5.2 Regulatory aspects
5.5.3 Properties of biodegradable nanoparticles (BNPs) which impact clinical use
5.5.4 Approved and investigational drugs with biodegradable polymeric nanoparticles of natural or synthetic origin
5.5.4.1 Anticancer drugs
5.5.4.2 Nanoparticles for oral delivery
5.5.4.3 Future trends: nanoparticles for vaccines and gene therapy
5.6 Future perspectives
Acknowledgments
References
Further reading
six Modulating the immune response with liposomal delivery
6.1 Introduction
6.1.1 Principles of lipid-based nanoparticles
6.1.2 Principles of the immune system
6.2 Liposomal immune modulation with small-molecule therapeutics
6.2.1 Principles of liposomal pharmacology based on Doxil
6.2.2 Immune modulation using small-molecule therapeutics
6.2.3 Future directions in liposomal immune modulation
6.3 Liposomal immune modulation with liposomal gene vectors
6.3.1 Principles of liposomal gene delivery
6.3.2 In vitro liposomal gene delivery
6.3.3 In vivo liposomal gene delivery
6.3.4 Immune modulation using liposomal gene vectors
6.3.5 Future directions for liposomal gene vectors
6.4 Immune stimulation with liposomal vaccines
6.4.1 Principles of Liposomal Vaccines
6.4.2 Liposomal adjuvants
6.4.3 Liposomal vaccine clinical trials
6.5 Conclusions and future directions
List of abbreviations
Acknowledgments
References
seven Recent advances in solid lipid nanoparticles formulation and clinical applications
7.1 Lipid nanoparticles
7.1.1 Solid lipid nanoparticles
7.1.2 Nanostructured lipid carriers
7.2 Formulation components
7.2.1 Lipids
7.2.1.1 Lipids polymorphic state
7.2.1.2 Types of lipids
7.2.1.3 Lipid proportion
7.2.1.4 Presence of a liquid lipid
7.2.2 Surfactants or emulsifiers
7.2.3 Other components
7.3 Preformulation studies
7.3.1 Solubility studies
7.3.2 Partitioning analysis
7.3.3 Compatibility between solid lipids and liquid lipids
7.4 Formulation procedures
7.4.1 High-energy methods
7.4.1.1 High-pressure homogenization
7.4.1.2 Emulsification–sonication technique
7.4.1.3 Supercritical fluid technology
7.4.1.4 Hot high-shear homogenization
7.4.2 Low-energy methods
7.4.2.1 Microemulsion technique
7.4.2.2 Double emulsion
7.4.2.3 Membrane contractor technique
7.4.2.4 Phase inversion technique
7.4.2.5 Coacervation
7.4.3 Organic solvent-based approaches
7.4.3.1 Solvent emulsification–evaporation method
7.4.3.2 Solvent emulsification–diffusion method
7.4.3.3 Solvent injection method
7.5 Characterization techniques
7.5.1 Particle size and size distribution
7.5.2 Surface charge
7.5.3 Morphology
7.5.4 Degree of crystallinity and polymorphism
7.5.5 Coexistence of different colloidal structures
7.5.6 Entrapment efficiency and drug loading
7.6 Drug incorporation models
7.6.1 Drug loading models of solid lipid nanoparticles
7.6.1.1 Homogeneous matrix
7.6.1.2 Drug-enriched shell
7.6.1.3 Drug-enriched core
7.6.2 Nanostructured lipid carriers drug loading models
7.6.2.1 Imperfect type
7.6.2.2 Amorphous type
7.6.2.3 Multiple oil-in-solid fat-in-water type
7.7 Administration routes
7.7.1 Topical administration
7.7.2 Parenteral administration
7.7.3 Oral administration
7.7.4 Pulmonary administration
7.7.5 Ocular administration
7.7.6 Intranasal delivery
7.8 Solid lipid nanoparticles and nanostructured lipid carriers case studies in humans for medical applications
7.8.1 Topical administration (skin and mucosa)
7.8.2 Oral administration
Self-assessment questions
References
eight Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants
8.1 Introduction
8.1.1 Terminology used to describe modified release systems
8.1.2 Advantages and limitations of modified release systems
8.1.3 Historical review of modified release systems
8.1.4 Formulation of modified release dosage forms
8.1.5 Classification of modified release systems according to the drug release rate mechanism
8.2 Mathematical models for drug release
8.2.1 Zero-order kinetics
8.2.2 First-order kinetics
8.2.3 Higuchi model
8.2.4 Hixson–Crowell model
8.2.5 Korsmeyer–Peppas model
8.2.6 Weibull model
8.2.7 Other release parameters
8.3 Release profiles comparison
8.4 Biopolymers in modified peroral drug delivery
8.5 Nanofibers in modified peroral drug delivery
8.6 Examples of liposomal-modified release formulations in clinical use
8.7 Conclusion
Self-assessment questions
References
nine Grafted polymethacrylate nanocarriers in drug delivery
9.1 Graft poly(meth)acrylates, including molecular brushes
9.2 Carriers based on poly(ethylene glycol) poly(meth)acrylate brushes
9.3 Carriers based on poly(ethylene glycol) grafted poly(meth)acrylates
9.4 Poly(ethylene glycol) and biodegradable polyester nonlinear amphiphilics
9.5 Other thermoresponsive graft polymethacrylate nanocarriers
9.6 Heterografted Janus-type carriers
9.7 Core–shell graft copolymers
9.8 Graft polymers containing disulfide linkers
9.9 Summary
References
Index

Citation preview

NANOMATERIALS FOR CLINICAL APPLICATIONS

NANOMATERIALS FOR CLINICAL APPLICATIONS Case Studies in Nanomedicines Edited by NATASSA PIPPA Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece Theoretical and Physical Chemistry Institute, National Hellenic Research Foundation, Athens, Greece

COSTAS DEMETZOS Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright © 2020 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-816705-2 For Information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisitions Editor: Simon Holt Editorial Project Manager: Emma Hayes Production Project Manager: Joy Christel Neumarin Honest Thangiah Cover Designer: Greg Harris Typeset by MPS Limited, Chennai, India

CONTENTS List of contributors

ix

1. Solid lipid nanoparticles in dermaceuticals

1

Indu Pal Kaur, Garima Sharma, Mandeep Singh, Mohhammad Ramzan, Joga Singh, Simarjot Kaur Sandhu and Jaspreet Singh Gulati 1.1 General introduction 1.2 Why solid lipid nanoparticles? 1.3 Evolution of lipidic nanoparticles from solid lipid nanoparticles to nanostructured lipid carriers 1.4 Cosmetic and topical applications of solid lipid nanoparticles 1.5 Skin penetration with solid lipid nanoparticles 1.6 Mechanism of drug penetration with solid lipid nanoparticles 1.7 Incorporation into semisolid vehicle 1.8 Case studies of successful topical delivery with lipidic nanoparticles 1.9 Delivery of agents for other skin diseases 1.10 Conclusions References

2. Cyclodextrin-based drug delivery systems Mario Jug 2.1 Cyclodextrins—structure, physiochemical properties, and toxicological profile 2.2 Cyclodextrin inclusion complexes—formation, stability, and application in drug delivery 2.3 Cyclodextrin-based products in clinical practice References

3. Lipid vesicles for (trans)dermal administration

1 2 3 5 11 12 13 13 15 22 22

29 29 31 33 60

71

Silvia Franzè, Umberto M. Musazzi and Francesco Cilurzo 3.1 (Trans)dermal drug-delivery systems 3.2 Lipid vesicles for breaching the skin barrier 3.3 Liposomal formulation in clinics 3.4 Final remarks References

71 74 88 92 92

4. Stimuli-responsive nanocarriers for drug delivery

99

Maria Chountoulesi, Nikolaos Naziris, Natassa Pippa, Stergios Pispas and Costas Demetzos 4.1 Introduction 99 4.2 Types of stimuli 100 4.3 Development of chimeric stimuli-responsive liposomes with incorporated stimuli-responsive polymers 103 v

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4.4 Thermotropic behavior of stimuli-responsive liposomes 4.5 Physicochemical properties of stimuli-responsive liposomes 4.6 Development of stimuli-responsive lyotropic liquid crystalline nanosystems 4.7 Conclusion and future directions References

5. Biodegradable nanomaterials Katerina Anagnostou, Minas Stylianakis, Sotiris Michaleas and Athanasios Skouras 5.1 Introduction 5.2 Natural polymers 5.3 Synthetic polymers 5.4 Polymeric nanoparticles 5.5 Clinical applications of biodegradable nanoparticles 5.6 Future perspectives Acknowledgments References Further reading

6. Modulating the immune response with liposomal delivery David Nardo, David Henson, Joe E. Springer and Vincent J. Venditto 6.1 Introduction 6.2 Liposomal immune modulation with small-molecule therapeutics 6.3 Liposomal immune modulation with liposomal gene vectors 6.4 Immune stimulation with liposomal vaccines 6.5 Conclusions and future directions List of abbreviations Acknowledgments References

7. Recent advances in solid lipid nanoparticles formulation and clinical applications Helena Rouco, Patricia Diaz-Rodriguez, Carmen Remuñán-López and Mariana Landin 7.1 Lipid nanoparticles 7.2 Formulation components 7.3 Preformulation studies 7.4 Formulation procedures 7.5 Characterization techniques 7.6 Drug incorporation models 7.7 Administration routes 7.8 Solid lipid nanoparticles and nanostructured lipid carriers case studies in humans for medical applications Self-assessment questions References

106 110 112 116 117

123 123 125 130 136 145 151 152 153 157

159 159 164 174 184 193 194 195 195

213 213 214 220 221 225 231 233 239 241 242

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Contents

8. Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants Marilena Vlachou and Angeliki Siamidi 8.1 Introduction 8.2 Mathematical models for drug release 8.3 Release profiles comparison 8.4 Biopolymers in modified peroral drug delivery 8.5 Nanofibers in modified peroral drug delivery 8.6 Examples of liposomal-modified release formulations in clinical use 8.7 Conclusion Self-assessment questions References

9. Grafted polymethacrylate nanocarriers in drug delivery Dorota Neugebauer 9.1 Graft poly(meth)acrylates, including molecular brushes 9.2 Carriers based on poly(ethylene glycol) poly(meth)acrylate brushes 9.3 Carriers based on poly(ethylene glycol) grafted poly(meth)acrylates 9.4 Poly(ethylene glycol) and biodegradable polyester nonlinear amphiphilics 9.5 Other thermoresponsive graft polymethacrylate nanocarriers 9.6 Heterografted Janus-type carriers 9.7 Core shell graft copolymers 9.8 Graft polymers containing disulfide linkers 9.9 Summary References Index

249 249 252 260 261 263 264 265 266 267

271 271 273 275 281 283 284 287 289 291 291 297

LIST OF CONTRIBUTORS Katerina Anagnostou Department of Electrical & Computer Engineering, Hellenic Mediterranean University Heraklion, Crete, Greece Maria Chountoulesi Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece Francesco Cilurzo Department of Pharmaceutical Sciences, University of Milan, Milan, Italy Costas Demetzos Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece Patricia Diaz-Rodriguez R+D Pharma Group (GI-1645), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain Silvia Franzè Department of Pharmaceutical Sciences, University of Milan, Milan, Italy Jaspreet Singh Gulati Hitech Formulations Pvt Ltd, Industrial Area 1, Chandigarh, India David Henson Department of Pharmaceutical Sciences, University of Kentucky, Lexington, KY, United States Mario Jug Faculty of Pharmacy and Biochemistry, University of Zagreb, Zagreb, Croatia Indu Pal Kaur University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Mariana Landin R+D Pharma Group (GI-1645), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain Sotiris Michaleas Department of Life Sciences, School of Sciences, European University Cyprus, Nicosia, Cyprus Umberto M. Musazzi Department of Pharmaceutical Sciences, University of Milan, Milan, Italy David Nardo Department of Pharmaceutical Sciences, University of Kentucky, Lexington, KY, United States Nikolaos Naziris Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece

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List of contributors

Dorota Neugebauer Silesian University of Technology, Faculty of Chemistry, Department of Physical Chemistry and Technology of Polymers, Gliwice, Poland Natassa Pippa Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece; Theoretical and Physical Chemistry Institute, National Hellenic Research Foundation, Athens, Greece Stergios Pispas Theoretical and Physical Chemistry Institute, National Hellenic Research Foundation, Athens, Greece Mohhammad Ramzan University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Carmen Remuñán-López NanoBiofar Group (GI-1643), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain Helena Rouco R+D Pharma Group (GI-1645), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain Simarjot Kaur Sandhu University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Garima Sharma University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Angeliki Siamidi Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece Joga Singh University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Mandeep Singh University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Athanasios Skouras Department of Electrical & Computer Engineering, Hellenic Mediterranean University Heraklion, Crete, Greece; Department of Life Sciences, School of Sciences, European University Cyprus, Nicosia, Cyprus Joe E. Springer Spinal Cord and Brain Injury Research Center, University of Kentucky, Lexington, KY, United States Minas Stylianakis Department of Electrical & Computer Engineering, Hellenic Mediterranean University Heraklion, Crete, Greece Vincent J. Venditto Department of Pharmaceutical Sciences, University of Kentucky, Lexington, KY, United States Marilena Vlachou Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece

CHAPTER ONE

Solid lipid nanoparticles in dermaceuticals Indu Pal Kaur1, Garima Sharma1, Mandeep Singh1, Mohhammad Ramzan1, Joga Singh1, Simarjot Kaur Sandhu1 and Jaspreet Singh Gulati2 1

University Institute of Pharmaceutical Sciences, Panjab University, Chandigarh, India Hitech Formulations Pvt Ltd, Industrial Area 1, Chandigarh, India

2

1.1 General introduction World Health Organization has included skin diseases as the most common noncommunicable diseases in hot and humid countries, including India; prevalence of these diseases is on the rise, world over. Skin being the largest organ that interfaces with the environment is exposed to a variety of physical [ultraviolet (UV) radiations], chemical, and microbial insults that affect its structure and function. Since significant part of skin is visible to others, any disfigurement of skin is often associated with social and psychological implications much beyond the actual disease symptoms. Global Burden of Disease survey reported skin and subcutaneous diseases as 18th leading cause of global disability-adjusted life years and 4th leading cause of nonfatal burden (Karimkhani et al., 2017). Years lived with disability from these diseases (36.4 million) are more than those caused by diabetes mellitus (29.5 million) and migraines (28.9 million) (Karimkhani et al., 2017). The global dermaceutical market (over the counter and prescription) is huge and evolving quickly indicating a global market of USD 91.40 billion in 2028 from USD 49.22 billion in 2018. The treatments available at present are unable to completely cure various diseases of the skin and meet the expectations of patients. This is attributed to either a lack of suitable treatment/agents or poor delivery of drug agent to the appropriate layer of the skin. The outermost part (15
20 μm) of the epidermis, namely stratum corneum (SC), is the major barrier to drug absorption into the skin. The resistant envelopes of SC corneocytes and keratin microfibrils are considered as bricks, and the lipidic layers found between these cells are called as mortar. This unique arrangement is responsible for basic skin permeation resistance and reduces the passage of molecules (larger than 500 Da) through skin (Erdogan, 2009). Although drugs may diffuse into the skin via hair follicles, sebaceous glands, or sweat glands, permeation through the multiple lipid Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00001-1

© 2020 Elsevier Inc. All rights reserved.

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bilayers of SC remains the main pathway (Ting et al., 2004) because the former comprise only a small area of the skin. The inability of drug molecules to penetrate the SC and reach the deeper dermis layer of the skin in sufficient concentration can usually result in recurrence of several skin diseases including infections that are often not limited to the SC. Small-sized carriers including liposomes, niosomes, aquasomes, transfersomes, elastosomes, microemulsions, nanoemulsions, self microemulsifying drug delivery system, self nanoemulsifying drug delivery system, and solid lipid nanoparticles (SLNs) are currently being explored extensively for their ability to permeate the SC and reach the lower skin layers including dermis and at times the subcutaneous tissue too. Both micrometer- and nanometer-range carriers were found effective in improving the delivery to skin. Since the drug is released gradually and over a prolonged period of time, irritancy or other side effects associated with the active ingredients, when applied in conventional formulations, are significantly reduced when incorporated into these systems, without compromising the efficacy (Castro and Ferreira, 2008). These systems not only mask the irritation and side effects of the selected agent to be delivered but also invariably improve its solubility and permeability.

1.2 Why solid lipid nanoparticles? In 1990 the lipidic nanoparticles were invented as an alternative to traditional drug carriers such as polymeric nanoparticles and liposomes. Many questions of ability to produce at industrial level, regulatory status of excipients, and nanotoxicity rose regarding the use of these conventional nanocarriers. Such questions were addressed with the development of SLNs as an alternative. Various advantages of SLNs over polymeric nanoparticles and liposomes are elaborated later.

1.2.1 Formulation aspects • • • •

SLNs can be prepared without employing organic solvents. Residues of these solvents have toxic effects (Kaur et al., 2014). High drug loading (B10% or more) can be achieved with SLNs versus a low drug loading of , 5% in case of polymeric nanoparticles (Singh et al., 2010). SLNs are reported to be stable for up to 3 years (Kakkar et al., 2011). This is an important advantage over other colloidal carrier systems. Remarkable scalability and reproducibility of important properties namely, particle size and encapsulation efficiency in large (Liu et al., 2007) batches using a

Solid lipid nanoparticles in dermaceuticals

cost-effective high-pressure homogenization technique as the preparation procedure is again an exclusive advantage with SLNs (Albanese et al., 2012). • SLNs can be sterilized by autoclaving. Other nano colloidal systems are sterilized by gamma irradiation, which is not only costly and a specialized technique but may possibly lead to the formation of free radicals and subsequently toxic reaction products (Kaur et al., 2014). • Quantification of SLN in creams is simplified as compared to other particles. Many cream bases do not exhibit a melting peak below 100 C, which means the content of SLN in a cream can be quantified by their melting peak determined by differential scanning calorimetry.

1.2.2 Physiological aspects • SLNs act as drug reservoirs in various skin layers (Vyas et al., 2014) by variety of uptake mechanisms like entering into a shunt such as hair follicle, accumulating between corneocytes, and intermingling with skin lipids, or by disintegrating and merging with lipidic layers (Toll et al., 2004; Bseiso et al., 2015). • Depending on the produced SLN type, controlled release of the active ingredients is possible. SLNs with a drug-enriched shell show burst release characteristics whereas SLNs with a drug-enriched core lead to sustained release (Wissing and Müller, 2003b). • SLNs act as occlusives, that is, they can increase the water content of the skin making it more hydrated and thus more permeable (Wissing et al., 2001). • SLNs show a UV-blocking potential, that is, they act as physical sunscreens on their own and can be combined with molecular sunscreens to achieve improved photoprotection (Wissing and Müller, 2003b). • The components used to formulate SLNs are safe as compared to polymeric nanoand microparticles which may cause systemic toxicity by impairment of the reticuloendothelial system due to slow degradation of its components up to 4 weeks (Cavalli et al., 2000).

1.3 Evolution of lipidic nanoparticles from solid lipid nanoparticles to nanostructured lipid carriers The most important parameters for evaluation of lipid nanoparticles are particle size and size distribution, zeta potential, polymorphism, degree of crystallinity, drug loading, entrapment efficiency, and drug release.

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The first generation of lipidic nanoparticles, that is, SLNs, necessarily comprise high-melting point lipid(s) which are heated at least once to melt and consecutively cooled. Latter results in the recrystallization of the lipid matrix leading to high possibility of polymorphism occurrence. Lipid particles crystallize in a higher energy modification (α or β0 ) which during storage transform to the low-energy, more-ordered modification (β). Drug molecules in SLNs, oriented between the fatty acid chains or glycerides, can be potentially expelled during transformation of the lipid from α to β form on storage. This happens due to the formation of more-ordered structure or reduced number of imperfections in the crystal lattice (Guimarães and Ré, 2011). Moreover because of their perfect crystalline structure, SLNs exhibit a low drug
loading efficiency (Ghasemiyeh and Mohammadi, 2018). To overcome these potential challenges faced by SLNs, the second-generation lipidic nanoparticles called nanostructured lipid carriers (NLCs) were introduced in 1999. Evolution of lipidic nanoparticles from emulsion to NLCs is shown in Fig. 1.1 (Guimarães and Ré, 2011). NLCs are composed of blends of solid and liquid lipids resulting in imperfections in the lattice which can accommodate a greater amount of the active ingredient. The less-ordered structure is due to inhibition of crystallization by liquid lipids. This enables a significant increase in loading capacity and also minimizes premature active ingredient expulsion. The structural comparison of SLN and NLC is depicted in Fig. 1.2

Traditional carriers 1950 s

1970 s

Innovative carriers 1991

1999/2000

Liquid lipid (oil)

Polymer (solid)

Solid lipids

Solid lipids + Liquid lipids (oil)

o/w emulsion

Polymeric nanoparticles

SLN 1st generation

NLC 2nd generation

Lipid nanoparticles

Figure 1.1 Evolution of lipid nanoparticle concept in comparison with the conventional technologies till the beginning of the 1990s. NLC, Nanostructured lipid carrier; SLN, solid lipid nanoparticle. Obtained with permission from Guimarães, K.L., Ré, M.I., 2011. Lipid nanoparticles as carriers for cosmetic ingredients: the first (SLN) and the second generation (NLC). In: Beck, R., Guterres, S., Pohlmann, A. (Eds.), Nanocosmetics and Nanomedicines, Springer, Berlin, Heidelberg, pp. 101
122.

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Solid lipid nanoparticles in dermaceuticals

SLN “Brick wall” structure

NLC Unstructure matrix

High drug load

Low drug load

Drug expulsion during storage

Long-term drug stability

Figure 1.2 “Symmetric brick wall” and “Welsh natural stone wall” model depicting difference between particle matrix structure of SLNs and NLCs, respectively. NLCs, Nanostructured lipid carriers; SLNs, solid lipid nanoparticles. Obtained with permission from Beloqui, A., Solinís, M.A., RodríguezGascón, A., Almeida, A.J., Préat, V., 2016. Nanostructured lipid carriers: promising drug delivery systems for future clinics. Nanomed. Nanotechnol. Biol. Med., 12, 143
161.

(Beloqui et al., 2016). Different patented and marketed products using SLN/NLC technology are given in Tables 1.1 and 1.2, respectively (Müller et al., 2007; Kaul et al., 2018), respectively.

1.4 Cosmetic and topical applications of solid lipid nanoparticles Occlusion: Epidermal layer of skin has 20% water content and play principal role as the barrier of topical absorption of foreign particles. Occlusion process can enhance hydration of SC layer which influence topical absorption. SLNs have the ability to form a hydrophobic monolayered film, which has attraction for the epidermal layer of skin. The occlusive nature of this film retards the water loss due to

Table 1.1 Important patents concerning the use of lipidic nanoparticles for topical dermal administration. Patent number Title Medical condition

Reference

US8715736B2

Skin inflammation

Singh and Patlolla (2009)

Hormone replacement therapy Atopic dermatitis

Viladot Petit et al. (2010)

Hair loss or acne

Cal and Frackowiak (2012)

Immuno booster and antiinflammatory

Jianqing Chen and Ping (2012)

Skin whitening

Jin et al. (2016)

Psoriasis and ichthyosis Moisturizing cream Skin care

Minmin and Haijun (2011) Omelyanchuk and Vilinskaya (2012) Sanz et al. (2009)

Hair loss

Padois et al. (2009)

ES2384060B1 RU2602171C2

EP2919756B1

CN102670484B

KR101860555B1

CN102342914A RU2491911C1 KR101810695B1

WO2010112749A1

Nanoparticle formulations for skin delivery Capsules lipid nanoparticles Composition containing lipid nanoparticles and corticosteroid or vitamin D derivative Solid lipid nanoparticles of roxithromycin for hair loss or acne Mannose-modified solid lipid nanoparticle plural gel and preparation method thereof Solid lipid nanoparticles composition for skin-whitening effect comprising MHY498 and preparation method thereof Calcipotriol solid lipid nanoparticle and preparation method of same Moisturizing cream with solid lipid nanoparticles Peptides used in the treatment and/ or care of skin, mucous membranes, and/or scalp and their use in cosmetic or pharmaceutical compositions Solid lipid nanoparticles encapsulating minoxidil and aqueous suspension containing same

Bastholm and Peterson (2012)

WO2017143421A1

WO2008041116A2

US6875438B2

US20180296493A1

JP2018521052A

Nanoscale system for the sustained release of active cosmetic and/or repellent substances Formulations of active principles incorporated in SLNs suitable for transdermal administration Preparations for topical administration of substances having antiandrogenic activity Lipid nanoparticle compositions and methods as carriers of cannabinoids in standardized precision-metered dosage forms Lipid and lipid nanoparticles formulations for delivery of nucleic acids

Insect repellent

De Paula et al. (2017)

Transdermal delivery of drugs with short half life Androgenic alopecia

Gasco (2006)

Antiinflammatory

Kaufman (2015)

Diseases related to a deficiency of a protein and enzymes

Du and Ansell (2016)

Kraemer et al. (2003)

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Table 1.2 Marketed products containing lipidic nanoparticles (Müller et al., 2007; Kaul et al., 2018). Marketed product

Active ingredients

Intended use

Cutanova Cream Nano Repair Q10

Q10, polypeptide, hibiscus extract, ginger extract, ketosugar Q10, polypeptide, mafane extract Q10, TiO2, polypeptide, ursolic acid, oleanolic acid, sunflower seed extract Kukuinut oil, Monoi Tiare Tahiti, pseudopeptide, milk, extract from coconut, wild indigo, noni extract Kukuinut oil, Monoi Tiare Tahiti, pseudopeptide, milk extract from coconut, wild indigo, noni extract Kukuinut oil, Monoi Tiare Tahiti, pseudopeptide, milk extract from coconut, wild indigo, noni extract Kukuinut oil, Monoi Tiare Tahiti, pseudopeptide, milk extract from coconut, wild indigo, noni extract Black currant seed oil containing omega 3 and 6 unsaturated fatty acids Coenzyme Q10 and black currant seed oil Coenzyme Q10, omega 3 and unsaturated fatty acids

Antiaging, smoothes fine lines, promotes restructuring Antiaging, antiwrinkle

Intensive Serum NanoRepair Q10 Cutanova Cream NanoVital Q10 SURMER Crème Legère Nano-Protection

SURMER Crème Riche Nano-Restructurante

SURMER Elixir du Beauté Nano-Vitalisant

SURMER Masque Crème Nano-Hydratant

NanoLipid Restore CLR

Nanolipid Q10 CLR IOPE Super Vital cream, serum, eye cream, extra moist softener, extra moist emulsion NLC Deep Effect Eye Serum NLC Deep Effect Repair Cream NLC Deep Effect Reconstruction Cream

NanoLipid Repair CLR

Coenzyme Q10, highly active oligosaccharides Q10, TiO2, highly active oligosaccharides Q10, acetyl hexapeptide-3, micronized plant collagen, high, active oligosaccharides in polysaccharide matrix Black currant seed oil and manuka oil

Antiaging

Skin-protecting serum

Intensely hydrating

Antiaging and moisturizing

Skin hydration

Skin hydration

Antiaging Antiaging, moisturizer under eye wrinkles, face lift Under eye wrinkles Antiaging Antiaging

Skin damage repair

(continued)

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Solid lipid nanoparticles in dermaceuticals

Table 1.2 (Continued) Marketed product

Active ingredients

Intended use

NLC Deep Effect Reconstruction SerumRegeneration screme Intensiv Swiss Cellular White Illuminating Eye Essence

Macadamia ternifolia seed oil, avocado oil, urea, black currant seed oil

Skin rejuvenation

Glycoproteins, Panax ginseng root extract, Equisetum arvense extract, Camellia sinensis leaf extract, Viola tricolor extract Glycoproteins, Panax ginseng root extract, Equisetum arvense extract, Camellia sinensis leaf extract, Viola tricolor extract Kukuinut oil, Monoi Tiare Tahiti, pseudopeptide, protein Olea europaea oil, panthenol, Acacia senegal, tocopheryl acetate Olea europaea oil, Prunus amygdalus dulcis oil, hydrolized milk protein, tocopheryl acetate, Rhodiola rosea root extract, caffeine Aloe barbadensis leaf juice, artemia extract, tocopheryl acetate, benzophenone-3, butyl methoxydibenzoylmethane (Parsol, 1789), ethylhexyl methoxycinnamate, ethylhexyl salicylate, homosalate Ethylhexyl glycerin, benzyl salicylate, citronellol, tocopherol, geraniol, alphaisomethyl ionone, coumarin, citral, ascorbyl palmitate, benzyl benzoate, ascorbic acid, citric acid, farnesol Sparkling notes of mandarin, rose, and vanilla Sparkling mandarin, may rose, sensual vanilla, and intoxicating passion fruit notes mixed with peony

Under eye dark circle lightening

Swiss Cellular White Intensive Ampoules

SURMER Creme Contour Des Yeux NanoRemodelante Olivenöl Anti Falten Pflegekonzentrat Olivenöl Augenpflegebalsam

Celazome MAX Sun Protection Factor (SPF) 29

Allure Body Cream

Allure Parfum Bottle Allure Eau Parfum Spray

Skin lightening

Antiwrinkle

Antiaging

Antiaging

Protect from sun, remove fine lines

Body moisturizer

Perfume Perfume

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evaporation. Experimental verification of moisture barrier properties has demonstrated the different degree of occlusion, depending on the size of the applied particles. It was further observed that maximum occlusivity was reached with SLNs having low-melting lipids, high crystallinity, and low-particle size (Müller et al., 2002). In a study, nanosized SLN and NLC systems were developed and showed a similar occlusion factor of 36%
39% with a reduction in transepidermal water loss of 34.3% 6 14.8% and 26.2% 6 6.5%, respectively. The marker (nile red) however showed that NLCs penetrate deeper into the SC as compared to SLNs (López-García and Ganem-Rondero, 2015). UV-blocking effect: The capability of SLNs to scatter and reflect the UV radiations makes them successful UV blockers. The matrix of SLNs measured higher UV absorption as compared to sunscreen of oil in water (o/w) nanoemulsion. Titanium dioxide is a commonly used UV blocker at molecular level. However, it exhibits significant side effects like photoallergies and phototoxicity (Wissing and Müller, 2003b). A synergistic effect was obtained when SLN were combined with sunscreen formulations. The amount of sunscreen Active Pharmaceutical Ingredients (API) could be decreased if combined with SLNs, thus minimizing the adverse reactions associated with these sunscreen molecules. Following are some examples of such combinations: 1. SLNs were found to act as excellent drug transport systems for oxybenzone, and the adverse effects like skin rashes, redness, and irritation were reduced significantly (Manea et al., 2014). 2. Stability of UV blocker agents was increased by incorporation into SLNs as carriers. Photodegradation of bis-ethylhexylphenolmethoxy-phenyltriazine was decreased considerably by incorporation into SLNs (Lee et al., 2007). SLNs encapsulating UV protector molecules also showed a better SPF factor and photostability (Lacatusu et al., 2011). 3. The SLN formulation with green tea prepared by high-pressure homogenization technique exhibited a higher photoprotective effect (Bose et al., 2013), good antioxidant activity, and better stability at room temperature. Adhesiveness: The application of formulation containing submicron-sized SLN (B200 nm) on dry horny layer shows good adhesiveness. SLNs form a film of tightly packed round particles, which under the applied force during application formed an intelligible film (Fig. 1.3). Such type of lipid film can help restore damaged skin or a broken lipid film on the skin surface. In addition to this, it can also have an occlusive effect (Wissing and Muller, 2003a). pH control and osmotic effect: Skin surface usually exhibits slightly acidic pH (pH 5.0
7.0). It is observed that a significant change in pH by application of any formulation can lead to irritation and redness of the skin. Strongly acidic and alkaline application will primarily act as deteriorating agents. SLN dispersions can be formulated or buffered at the skin optimum pH thus making them optimal for dermal application. Some

11

Solid lipid nanoparticles in dermaceuticals

2µm

200 nm

Section: H2O evaporation

Skin Top view:

Large pores Small ''capillary pores''

Fusion:

Application and capillary forces

Figure 1.3 Model of film formation on the skin for lipid 2-mm particles and lipid 200-nm particles shown as section (upper) and from the top (middle), and a new model of fusion of the nanoparticles to a poreless film (lower). Obtained with permission from Müller, R.H., Radtke, M., Wissing, S.A., 2002. Solid lipid nanoparticles (SLN) and nanostructured lipid carriers (NLC) in cosmetic and dermatological preparations. Adv. Drug. Delivery Rev. 54, S131
S155.

considerations are also given to osmotic effect of the topical formulation. Change in isotonicity can lead to irritation. NLCs and SLNs show remarkable isotonicity and good osmotic effect following their application on skin (Souto and Müller, 2008) Improved chemical stability: Solid matrix of SLNs is stable at room temperature as well as under physiological conditions. The solid core of SLNs better accommodates APIs which are prone to hydrolysis and oxidation, protecting them against chemical degradation from water and oxygen, for example, SLN dispersions of tocopherol, retinol, and coenzyme Q10 are chemically more stable as compared to the corresponding aqueous dispersions of free or unencapsulated agents (Dingler et al., 1999). It may be noted that tretinoin incorporated into liposomes was found prone to photodegradation (Brisaert et al., 2001), while retinol encapsulated within SLNs was chemically stable (Volkhard and Gohla, 2001). SLNs per se also exhibited high physical stability during long-term storage (Müller et al., 2007).

1.5 Skin penetration with solid lipid nanoparticles For local as well as systemic effects, skin is considered to be the important site for drug application where SC is the main penetration barrier. Modern techniques

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usually aim at disrupting or bypassing the complex skin structure for drug delivery. Intercellular route is the most common route followed by the molecules to penetrate through the skin. SC is composed of morphologically different cells which carry diverse functions of this second largest organ of the body. It is 6
10 μm in thickness with cell and lipid layer alternating with each other. It is made up of B14
18 cellular layers and the barrier nature is attributed to the presence of 75%
80% proteins. In addition, lipids (5%
20%) and unidentified elements (5%
8%) also contribute toward hindrance for the drug molecules to pass through the skin (Wertz, 2018). Corneocytes are the uppermost cells of the epidermis and are observed to be stacked in the form of pillars when seen microscopically.

1.6 Mechanism of drug penetration with solid lipid nanoparticles As highlighted earlier intercellular pathway, that is, movement of drug molecules between the corneocytes of the SC is the most preferred route. Another important pathway for transport is the intrafollicular pathway also called shunt or appendageal pathway where penetration into the skin occurs through the sweat glands or the hair follicular route as seen in Fig. 1.4 (Palmer and DeLouise, 2016). Lipidic nanoparticles getting attached to the skin surface have the ability to conduct the exchange of lipid between SC (composed of high concentration of lipids) and

Figure 1.4 Diagrammatic representation of skin penetration with solid lipid nanoparticles.

Solid lipid nanoparticles in dermaceuticals

nanocarriers. Richness of SC in epidermal lipids and the nano size of SLNs trigger a precise interaction between the two, thus providing an increased penetrating power, occlusivity, and concentration of encapsulated drug in the dermal region of the skin. Skin occlusivity provided by the SLNs also increases skin hydration and thus the skin penetration (Korting et al., 2007). Nanoparticles of size greater than 100 nm do not perfuse the SC, majorly due to the rigidity and dimensions of the barrier layer. However, particles of B200 nm size provide an occlusive protective layer that in turn enhances penetration of the skin. Sebaceous glands, designated to secrete sebum, are associated with the hair follicles inside the skin. Sebaceous secretions are rich in lipids which provide a perfect environment for lipid nanoparticles to dissolve and release the encapsulated active. Sebum is a mixture of waxes and triglycerides which are also used in preparing SLNs/NLCs (ones prepared with biocompatible lipids). Former accelerate the absorption of the drug through these glands. Thus this route is particularly favorable for lipid nanoparticles and is exploited in antiacne therapy (Ranpise et al., 2014).

1.7 Incorporation into semisolid vehicle SLN dispersions are usually free flowing and can be easily incorporated into dermal carrier systems like gels and creams to formulate a dosage form of desired consistency to be applied topically. Control release, targeting the viable epidermis, and tissue compatibility are the characteristics particularly achievable with the gel systems. Gels are also preferred due to ease in manipulation for swelling level required in the final formulation (Housiny et al., 2018). Carbopol is mainly used as gelling agent, polymers of which crosslink together to form a microgel structure that make it ideal for dermatological purposes. These structures get adhered to the skin, increasing the contact time (Deshkar et al., 2018)

1.8 Case studies of successful topical delivery with lipidic nanoparticles 1.8.1 Delivery of antimicrobials 1. Gide et al. prepared SLNs with acyclovir (ACV-SLNs) in their matrix and further incorporated ACV-SLNs into semisolid gel. Study showed that amount of acyclovir from ACV-SLNs in the lower epidermal layer was two times of that achieved with commercial acyclovir gel (Gide et al., 2013).

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2. Aqueous dispersions of ketoconazole-loaded SLNs and NLCs with Compritol 888 ATO as the lipid were developed (Souto and Müller, 2005). Stability study revealed that SLN dispersion was physically stable and no significant change in particle size was observed; however NLCs also protected the drug from degradation but increase in size was observed upon storage. 3. Sanna et al. prepared SLNs loaded with econazole nitrate (ECN) for administration across the skin. They concluded from skin permeation study that controlled release of drug across SC was governed by lipid portion in SLNs. In vivo results showed that SLNs enhanced the penetration of ECN to deeper skin layers after 3 hours of administration (Sanna et al., 2007). 4. Miconazole (MN)-loaded SLNs (MN-SLN) when incorporated into a hydrogel showed a controlled release of MN over a 24-hour time period. About 10-fold increase in retention was noted with MN-SLN as compared to free MN drug suspension and MN gel. In vivo studies indicated that the hydrogel of MN-SLN provided sustained topical effect and treated the fungal infections at a faster rate ( Jain et al., 2010). 5. Topical gel of fluconazole (FLZ)-loaded SLNs were developed and evaluated clinically for pityriasis vesicolor. Results showed a significant improvement (P , .05) in therapeutic response with FLZ-loaded SLNs, as 90% patients showed complete eradication and 10% showed significant improvement in comparison to marketed cream (Housiny et al., 2018). 6. Stability study with clotrimazole-loaded SLN and NLC showed that these systems retained their colloidal phase after 3 months of storage at temperature of 4 C, 20 C, and 40 C. Release study showed controlled release of clotrimazole from both SLNs and NLCs over a period of 10 hours (Souto et al., 2004). 7. Voriconazole (VCZ) lipid-based nanoparticles (LNP) developed for the treatment of aspergillosis could improve the solubility of VCZ. The antifungal study revealed that optimized formulation of VCZ-LNP stopped fungus reproduction (Füredi et al., 2017). 8. Griseofulvin-loaded SLNs (GF-SLNs) enhanced the rate of dissolution of GF due to the decrease in mean particle size (165 nm). Release rate of GF showed sustained release (cumulative 63.53%) over a period of 12 hours (Anurak et al., 2011). 9. Amphotericin B
loaded SLNs for topical application with small size (111.1 6 2.2 nm) and higher entrapment efficiency (93.8% 6 1.8%) are reported. Freeze-dried SLNs showed two times higher permeation and higher zone of inhibition against Trichophyton rubrum as compared to free drug dispersion. The formulation was stable at refrigerator and room temperature over the period of 3 months (Butani et al., 2016).

Solid lipid nanoparticles in dermaceuticals

1.9 Delivery of agents for other skin diseases 10. Recombinant human epidermal growth factor
loaded SLN and NLC preparations showed the reduction of healing time of chronic wounds (Gainza et al., 2014, 2015). 11. A hydrogel formulation with astragaloside IV (a herb widely used in Chinese medicines)-loaded SLN for improved wound healing and inhibiting scar formation were developed using Carbopol 934 as the gelling agent. The formulation showed a sustained release and enhanced skin residence time. The in vitro studies of astragaloside IV
loaded SLN showed an increase in drug uptake by fibroblasts and enhanced migration and proliferation of keratinocytes to the wound (Chen et al., 2013). 12. An o/w cream containing tretinoin-loaded SLN for the treatment of dermatologic affections, such as acne, psoriasis, and ichthyosis is reported. The SLNs were incorporated, with a 50:50 ratio, in a previously prepared o/w cream. The formulation showed no aggregation or phase separation even after storage for 24 months. The ex vivo studies, performed in rat skin, showed that the drug release was lower for the semisolid formulation compared to the drug-loaded SLN dispersions and o/w cream (Nasrollahi et al., 2013). 13. The potential of using SLNs to improve the dermal delivery of spironolactone for treatment of acne was explored. Ex vivo permeation studies in rat skin showed a 1.6-fold higher drug penetration from SLNs in comparison to the free drug (Kelidari et al., 2015). 14. Carbopol 934 hydrogels with adapalene-loaded SLNs revealed the presence of high drug amount in epidermis over dermis, meaning a reduced systemic absorption with decreasing side effects, in ex vivo studies using rat skin ( Jain et al., 2014). 15. Hydroquinone-loaded SLNs were developed with an aim to overcome the drawbacks associated with the topical use of this compound (for hyperpigmentation), such as fast oxidation, insufficient penetration due to hydrophilic nature, and adverse effects related to systemic absorption. The SLN dispersion was transformed into a nanoemulgel by addition of Carbopol 934 to the aqueous phase (Ghanbarzadeh et al., 2015). 16. Ex vivo studies on resveratrol (RSV)-loaded SLN, carried out on pig skin, showed a drug cumulative amount of up to 45% after 24 hours. The developed SLN also showed good tyrosinase inhibition activity in comparison to control solution. Developed formulation also showed no cytotoxic effects in keratinocytes at concentrations up to 250 μg/mL (Rigon et al., 2016). 17. Nanoemulgel-based SLN for skin targeting of RSV were suggested as an innovative topical treatment for irritant contact dermatitis with minor side effects (Rigon et al., 2016).

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18. Human clinical evaluation of progesterone-loaded SLN and NLC-based xanthan gum hydrogel using tape stripping experiments to observe drug penetration across SC showed penetration within 6 hours after topical use of the formulations (Esposito et al., 2017). 19. Idebenone, a synthetic derivative of ubiquinone, shows a potent antioxidant activity that could be beneficial in the treatment of skin oxidative damages. The feasibility of targeting idebenone into the upper layers of the skin by topical application of SLNs was evaluated. SLN of idebenone were prepared by the phase inversion temperature method using cetyl palmitate as solid lipid and three different nonionic surfactants: ceteth-20, isoceteth-20, and oleth-20. All idebenone-loaded SLN showed a mean particle size in the range of 30
49 nm and a single peak in size distribution (Montenegro et al., 2012). 20. Naproxen-loaded SLNs were prepared by the ultrasonication method to improve its skin permeation and also to investigate the influence of hydrophilic
lipophilic balance (HLB) changes on nanoparticle properties. The average particle size and polydispersity index (PDI) of SLNs increased from 94.257 6 4.852 to 143.90 6 2.685 nm and from 0.293 6 0.037 to 0.525 6 0.038, respectively, when there was an increase in lipid concentration from 0.3% to 1.2%. The results also showed that a reduction in the HLB resulted in an increase in the PDI, particle size, zeta potential, and entrapment efficiency. The amount of naproxen detected in the receptor chamber at all the sampling times for the reference formulation (naproxen solution containing all surfactants at pH 7.4) was higher than that of the SLN formulation indicating that higher concentration of naproxen is retained in the skin layer with less systemic absorption (Akbari et al., 2016). 21. Auraptene is a bioactive antioxidant coumarin with valuable pharmacological properties; however, poor water solubility is a substantial issue for the topical application of auraptene. Therefore auraptene was incorporated into SLNs for enhancing its antiinflammatory effect. Auraptene-SLNs were compared to conventional cream containing auraptene for evaluating both the permeation rate and the antiinflammatory effect of the nanoparticles through in vitro and in vivo studies. The in vitro permeation studies exhibited that Auraptene-SLNs significantly enhanced cutaneous uptake of auraptene and skin targeting. The antiinflammatory and histopathological studies exhibited no significant differences between Auraptene-SLNs and indomethacin (Daneshmand et al., 2018). 22. Systemically administered methotrexate is indicated in psoriasis therapy. Its topical administration may be an option to overcome various side effects associated with its systemic delivery. Lipid nanoparticles were developed for codelivery of methotrexate with etanercept using lipid nanoparticles, after incorporating into a carbopol hydrogel. SLNs were evaluated for their potential of delivering the drug into the skin with reduced transdermal permeation. Permeation studies using pig ear as

Solid lipid nanoparticles in dermaceuticals

23.

24.

25.

26.

model revealed enhanced skin deposition of the applied methotrexate when incorporated within SLNs in relation to free drug (Ferreira et al., 2017). SLNs based on trehalose monooleate, loaded with cyclosporin-A for potential treatment of psoriasis, were prepared. Trehalose was esterified with oleic acid to obtain a more lipophilic compound suitable as a lipid matrix for the formulation of a new type of SLNs. With the aim to verify the localization of SLNs in the skin layers, investigations using confocal microscope and stripping tape test were carried out. The dermatological formulation, based on SLNs loaded with cyclosporin-A, was prepared and tested through Franz diffusion cells. The obtained results indicated the possibility of using these nanoparticles as vehicle of cyclosporin-A for topical treatment of psoriasis with reduced side effects (Trombino et al., 2019). A rapid one-step method was used to formulate the corticosteroidal drug, diflucortolone valerate into topical semisolid SLN formulations using a high-shear homogenization combined with sonication, using different types of solid lipids (e.g., Geleol, Precirol ATO5, Tristearin, and Compritol 888 ATO) and Poloxamer 407 as a surfactant. Diflucortolone valerate SLN formulations possessed average particle size ranging from 203.71 6 5.61 to 1421.00 6 16.32 nm with a narrow size distribution and possessed shear thinning behavior. Use of lipid-based surfactants (Labrasol or Labrafil) was found to significantly increase diflucortolone valerate encapsulation efficiency (up to 45.79% 6 4.40%) (AbdelSalama Fatma et al., 2016). Neem oil as a natural agent was incorporated into SLNs prepared by double emulsification method using different combinations of lecithin and Tween 80. The average particle size of neem oil-loaded SLNs decreased with increasing concentration of surfactant. SLNs of 221.6 6 2.0 nm with a PDI of 0.948 6 0.04 were obtained at higher concentration of lipid and surfactant. High entrapment of 82.10% revealed the ability of SLNs to incorporate a high quantity of neem oil actives (Vijayan et al., 2013). Dacarbazine-laden nanoparticle (DZNP) and dacarbazine-laden nanocream (DZNC) topical delivery system were prepared for the treatment of melanoma. DZNP was used to prepare a cream formulation for topical drug delivery for melanoma. DZNP and DZNC were evaluated for morphology, drug load capacity, efficiency of nanoencapsulation and size of particle, and zeta potential. Nanoencapsulation efficiency and drug-loading capacity were 67.4% 6 3.5% and 6.73 mg/10 mg, respectively. IC50 of DZNP was 0.19 mg/mL while it was 0.63 mg/mL for DZNC. It can be concluded that DZNP and its cream can be effectively used as a topical formulation for the treatment of melanoma (Hafeez and Kazmi, 2017).

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27. Paclitaxel-loaded SLNs were prepared using high-speed homogenization and ultrasonication technique. The optimized SLN were loaded in carbopol gel. The prepared gels were evaluated for in vivo anticancer efficacy and histopathological study. The histopathological study established the efficacy of SLNs in the treatment of skin cancer (Bharadwaj et al., 2016). 28. Varying ratios of lecithin and poloxamer 188 were used to produce shellenriched nanoparticles of 5-fluorouracil (5-FU) by enabling the formation of reversed micelles within the region of the SLN. The localization of 5-FU within the shell region of the SLN was confirmed using 5-FU nanogold particles as a tracer. SLN were introduced within sodium carboxy methyl cellulose hydrogel and then applied onto the skin of mice-bearing Ehrlich’s ascites carcinoma. The mice were treated with the gel twice daily for 6 weeks. SLN-treated mice exhibited reduced inflammatory reactions with reduced degrees of keratosis, in addition to reduced symptoms of angiogenesis compared to free 5-FU-treated mice (Khallaf et al., 2016). 29. Topical application of curcumin is known to relieve irritant contact dermatitis and hyperpigmentation, but its topical delivery is a challenge due to its low solubility. Encapsulation of curcumin into SLNs makes it agreeable to topical dosing due to small size of SLNs which further promotes its penetration into the skin. SLNs were prepared using Precirol ATO5 and Tween 80 by probe ultrasonication method. Further it was incorporated into Carbopol gel and investigated for ex vivo skin permeation, skin deposition, and skin irritation studies. In vitro tyrosinase inhibition assay indicated that the formulated gel showed a potential in skin depigmentation. The gel proficiently suppressed ear swelling and reduced skin water content in the BALB/c mouse. Thus the SLN gel was found to be safe and effective substrate in comparison to conventional vehicles for treatment of irritant contact dermatitis and pigmentation (Shrotriya et al., 2018). 30. Sesamol-loaded SLNs were prepared using microemulsification technique with particle size of 127.9 nm (PI: 0.256) and entrapment efficiency of 88.21%. Topical application of sesamol SLN in a cream base indicated significant retention in the skin with minimal flux across skin. Efficacy was confirmed by the in vivo skin retention and ex vivo skin permeation studies. In vivo anticancer studies performed using 12-O-Tetradecanoylphorbol-13-acetate-induced and benzo(a)pyrene initiated tumor production (ROS mediated) in mouse epidermis showed normalization (in histological studies) of skin cancers post their induction, upon treatment with sesamol SLN (Geetha et al., 2015). 31. In another study, green tea (Camellia sinensis) leaves extract
loaded SLNs showed improved penetration of epigallocatechin through the SC. SLNs were prepared by solvent emulsification method using glyceryl monostearate as the lipid (2%) and poloxamer 188 (4%) as a surfactant (Dzulhi et al., 2018).

Solid lipid nanoparticles in dermaceuticals

32. Triptolide (TPL), a diterpene lactone widely used to treat inflammation, autoimmune diseases, malignancy, and depression, was formulated into NLCs and SLNs using microemulsion method. Histopathological studies confirmed that TPL nanoparticles (TPL-NPs) permeated into skin through change in SC structure. Both the interactions between mixed surfactants (Tween 80 and Transcutol HP) and skin and lipid exchange between lecithin and SC increased the skin intercellular space which in turn loosened SC dense structure. Besides, TPL-NPs could inhibit or prevent evaporation, which improved skin moisture and hydration. Compared with TPL-SLN, TPL-NLC showed stronger interaction with skin, which may be related to the higher drug loading % (Gu et al., 2018). 33. Adapalene (Ada) SLNs were prepared and reported by hot melt homogenization method followed by its gelling with Carbopol 980 NF, Carbopol Ultrez 10 NF, and pemulen TR-1 (all approved by FDA for human use). Drug release from free-Ada suspension demonstrated B98% release in 8 hours, while B38% and B30% adapalene was released from Ada-SLNs and Ada-SLNs gel, respectively, in 48 hours. The Ada-SLNs followed characteristic initial burst followed by sustained release which may be ascribed to adapalene adhering to the surface of the lipid and adapalene encapsulated in solid matrix of lipid, respectively (Harde et al., 2015). All grades of carbopol produced stable gel without inducing agglomeration of SLNs. However, authors selected Carbopol Ultrez 10 on the basis of high viscosity and high spreadability as preferred gelling agent. Carbopol Ultrez 10 has the ability to form interconnected network of polymer chains, producing higher viscosities. Additionally publication cited the properties like ease of dispersibility in comparison to conventional gelling agents (e.g., acacia, hydroxyl propyl methyl cellulose, cellulose) and low thermosensitivity in aqueous system making Carbopol Ultrez 10 a suitable gelling agent. 34. Clinical studies with acitretin (Act) NLCs, prepared by solvent diffusion technique, demonstrated significant improvement in therapeutic response and reduction in local side effects indicating its effectiveness in the topical treatment of psoriasis. Act-loaded NLC gel was prepared by dispersing required quantity of Carbopol 934P (1% w/w) in small quantity of distilled water and allowing it to hydrate for 4
5 hours. Propylene glycol (10% w/w) and glycerol (30% w/w) were subsequently added to the aqueous dispersion. In vitro skin deposition studies showed significantly higher (P , .05) deposition of Act from Act-NLC gel (P , .05) in the skin than plain Act gel. Clinical findings indicated that Act-NLC gel significantly improved therapeutic index in terms of psoriasis area and severity index score, physician’s assessment, psoriasis symptom assessment, and safety and tolerability assessment ratings (Agrawal et al., 2010).

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1.9.1 Solid lipid nanoparticles for cosmetic applications 35. Enhancement of long-term chemical stability of ascorbyl palmitate (AP) was achieved after incorporation into NLCs and by addition of antioxidants coupled with nitrogen gas flushing. The particle size and zeta potential was found to be lower than 350 nm and 230 mV, respectively. The percentage of drug remaining at both 4 C and room temperature (25 C) was higher than 85% during 90 days of storage (Teeranachaideekul et al., 2007). 36. SLNs for all-trans retinol (AR) were formulated to improve the stability of AR. The mean particle diameter and zeta potential of the smallest SLNs were 96 nm and 228 mV, respectively. The loading of AR in optimized SLNs decelerated the degradation of AR compared with AR solution dissolved in methanol. It was further observed that coloading of antioxidants greatly enhanced the stability of AR loaded in SLNs, compared with those loaded in SLNs without antioxidant. The photostability (at 12 hours) of AR in SLNs was enhanced (B43%) than that in methanol solution (B11%) ( Jee et al., 2006). 37. Wissing et al. used SLNs for incorporation of perfumes into cosmetic products. Comparing the release of the perfume from an emulsion and SLN, after 6 hours, it was observed that 100% of the perfume was released from the emulsion but only 75% was released from the SLN. They concluded that the release of perfume was prolonged on incorporation into SLNs as compared to when present in an emulsion (Wissing et al., 2000). 38. UV radiation attenuators were loaded in nanolipidic systems with the lipid matrix composed of carnauba wax and decyl oleate. The particle size of the encapsulated species (barium sulfate, strontium carbonate, and titanium dioxide) was between 239 and 749.9 nm with PDI ranging from 0.100 to 0.425. Surface charge values of up to 240.8 mV proved the stability of dispersions. The formulations were of ideal viscosities (1.40
20.5 mPa s). Upon encapsulation of titanium dioxide a significant increase in SPF of up to 50 was observed (Villalobos-Hernandez and Müller-Goymann, 2005). 39. The ability of crystalline cetyl palmitate SLN to reflect and scatter UV radiation without any molecular sunscreens, leading to protection against sunlight, was investigated. In vitro assay revealed two to three times potency of placebo cetyl palmitate SLN formulation in absorbing UV radiation in comparison to conventional emulsion. It was concluded that the incorporation of sunscreens into SLNs lead to a synergistic photoprotection (higher than the additive). Further scanning electron microscopy proved formation of a dense film on the skin following fusion of particles and evaporation of water (Wissing and Müller, 2001). 40. The influence of the crystallinity of lipid nanoparticles on their occlusive properties was investigated in another study. SLN dispersions with different crystallinity

Solid lipid nanoparticles in dermaceuticals

41.

42.

43.

44.

45.

indices were produced, physicochemically characterized, and their occlusion factor was determined after 6, 24, and 48 hours. The study was based on the in vitro occlusion test by de Vringer. It was found that the occlusion factor depends strongly on the degree of crystallinity of the lipid matrix, that is, this effect is proportional and the desired degree of occlusivity can be achieved by choosing suitable lipids (Wissing and Muller, 2003a). The structure of loaded SLN was elucidated to study the incorporation of coenzyme Q10. A pattern of spin diffusion between protons of the lipid and protons of the coenzyme Q10 was observed, which indicated that the majority (60%) of the coenzyme Q10 is homogeneously mixed with the solid lipid, the residual amount, that is, 40% forms a separate solid phase associated to the particles. Another smaller fraction forms separate domains on the nanometer scale (Wissing et al., 2004). Pople and Singh (2006) reported SLN to be a promising particulate carrier having controlled drug release, improved skin hydration, and potential to localize Vitamin A palmitate in the skin with no skin irritation. The mean particle size was found to be 350 nm. Vitamin A palmitate nanoparticulate dispersion and gel revealed prolonged drug release (24 hours), owing to embedment of drug into solid lipid core. In vitro penetration studies signify two times higher drug concentration in skin with lipid nanoparticle gel as compared with blank gel. In vivo skin hydration studies showed increase in thickness of the SC with enhanced skin hydration. Formulation was nonirritant with no erythema or edema having primary irritation index of 0.00. Vitamin E (α tocopherol) SLNs were produced for an improved topical delivery. The developed Vitamin E-SLN dispersion showed a mean particle size of 292 nm. The in vitro release profile of vitamin E from SLN dispersion and its gel showed prolonged drug release for 12 hours. The in vitro penetration studies revealed about two times higher drug concentration in skin in case of SLNenriched gel as compared with other conventional gels (Subbaiah et al., 2011). In comparison to RSV-loaded SLNs, the corresponding NLCs penetrated deeper into the skin. Moreover RSV-loaded NLC with smaller particle size and higher drug loading appear to be superior to SLN for dermal applications (Gokce et al., 2012) Coenzyme Q10 (Q10) SLNs of average particle size 142.4 nm were prepared. Q10 entrapment efficiency was 89% and the production yield was 94%. Trolox equivalent antioxidant capacity analysis showed that antioxidant potential of Q10 can be protected in SLNs. Diffusion studies from rat abdominal skin revealed that the delivery of Q10 was doubled in SLN incorporating gels (B40 μg/cm2), in comparison with gels prepared with only Q10. It was concluded that Q10-SLNloaded gels efficiently deliver Q10 into the skin without losing its antioxidant properties (Korkm et al., 2013).

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46. The stabilizing effect of carrier systems for AP was investigated using microemulsions (ME), liposomes, and SLNs. Analysis of chemical stability by high performance liquid chromatography showed that AP is most resistant against oxidation in nonhydrogenated soybean lecithin liposomes, followed by SLN, water in oil (w/o) and o/w microemulsions, and hydrogenated soybean lecithin liposomes. It was further observed that the encapsulation of AP in SLN core leads to greater stability. It was concluded that the location of the sensitive group of the drug molecule in a carrier system is crucial for its stability (Kristl et al., 2003).

1.10 Conclusions Progress of nanomedicine over the last decade has provided ample platforms for launch of novel technologies for improved delivery of drugs to various body parts including skin. Researchers have lately devoted to develop SLN systems for skin application that offer advantages in terms of better drug loading and encapsulation efficiency, colloidal stability, adhesion, film formation, and occlusion in contrast to liposomes and polymeric nanoparticles. SLNs provide an inimitable prospect for the development of highly competent and low-toxicity treatment modalities for skin as it offers easy entry and deeper penetration into the skin. The incorporation of SLNs into a semisolid base further enhances stability, application, and acceptance of the final product. This chapter described several case studies that successfully employed SLNs/ NLCs for dermaceutical products. Despite the established success of SLN systems with some such products in the market, the wider application of SLNs is still limited as some challenges still remain to be overcome. On the other hand, due to the high safety threshold of the lipids and emulsifiers used for SLN preparation (the ingredients are usually accepted by FDA or are of GRAS status), these systems are invariably found to be safe in in vitro cytotoxicity assay and in vivo toxicity tests. However, an important consideration for these lipid-based colloid systems is their transformation from lab-scale to industrial production and final product development.

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CHAPTER TWO

Cyclodextrin-based drug delivery systems Mario Jug Faculty of Pharmacy and Biochemistry, University of Zagreb, Zagreb, Croatia

2.1 Cyclodextrins—structure, physiochemical properties, and toxicological profile Cyclodextrins (CDs) are structurally related cyclic oligosaccharides produced by the enzymatic degradation of starch. Naturally occurring αCD, βCD, and γCD are composed of 6, 7, or 8 α-1,4-linked D-glucopyranose units, respectively. CDs with less than 6 glucopyranose units cannot be formed due to steric reasons, whereas large CDs with more than 8 units has been identified, however, are not relevant in drug formulation and delivery. Because of the chair conformation of D-glucopyranose units, CDs present a structure of a hollow truncated cone (Fig. 2.1) ( Jansook et al., 2018; Larsen, 2002). Hydroxyl groups are oriented to the exterior of the molecule, with secondary ones directed toward narrower edge of the cone and primary on the wider edge, giving the hydrophilic character to exterior of CD molecule. The central cavity of the CD is lined with skeletal carbons and ethereal oxygens that gives it lipophilic character, with polarity resembling to that of diluted ethanol (Del Valle, 2004; Duchêne and Bochot, 2016; Kurkov and Loftsson, 2013). Natural CDs and particularly βCD are less soluble in water than liner maltodextrins of comparable molecular weight due to relatively strong inter- and intramolecular hydrogen bonding between vicinal C2 and C3 hydroxyl groups. This diminishes their ability to form hydrogen bonds with the surrounding water molecules and leads to formation of highly stabile crystal lattice. Substitution of any of hydroxyl group in the CD molecule disrupts the hydrogen bonds formation and causes a dramatic increase in the CD solubility ( Jansook et al., 2018; Saokham et al., 2018). Furthermore, modified CDs are a complex mixture of different stereoisomers with an amorphous structure, that additionally contributes to their aqueous solubility. This lead to the synthesis of numerous CD derivatives with increased solubility, improved functionality, and more acceptable toxicological profile (Kurkov and Loftsson, 2013). The list of most important CDs in the field of drug delivery are given in Table 2.1. Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00002-3

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Figure 2.1 Chemical structure and schematic representation of β-cyclodextrin. Table 2.1 Some physiochemical properties of pharmaceutically most relevant cyclodextrins (Saokham et al., 2018). Surface tension Log Derivative Average number of Solubility at (mM/m) PW/O substituents per CD 25 C (mg/ mL) molecule

α-Cyclodextrin β-Cyclodextrin Randomly methylated β-cyclodextrin Dimethyl-β-cyclodextrin 2-Hydroxypropylβ-cyclodextrin Sulfobutyletherβ-cyclodextrin sodium salt γ-Cyclodextrin 2-Hydroxypropylγ-cyclodextrin

13 14

9.7 13.6

145 18.5 . 500

71 71 62

12 16 2.8 10.5

570 . 1200

6 11

57.5 54.1 54.8 57.5

6.2 6.9

. 1200

10

71

3.0 5.4

232 800

17 13

71 71

After oral administration, γCD is readily digested in small intestine mainly by pancreatic α-amylases, whereas αCD, βCD, and hydrophilic CD derivatives are digested by bacterial microflora present in colon ( Jansook et al., 2018; Saokham and Loftsson, 2017). High molecular weight, presence of numerous hydrogen-bond donors and acceptors in the structure and pronounced hydrophilicity lead to low permeability of CDs across biological membranes. Because of that, CDs in general show limited oral bioavailability in animals and humans, ranging from 0.1% to 3%, which all makes them practically nontoxic when administered orally. The total daily dose of αCD, βCD, γCD and 2-hydroxypropyl-β-cyclodextrin (HPβCD) taken orally may

Cyclodextrin-based drug delivery systems

reach 6, 0.5, 10, and 8 g/day, respectively. The only exception of this is randomly methylated β-cyclodextrin (RAMEB), which is somewhat more lipophilic, thus having an oral bioavailability of 12% in rats. Both αCD and βCD are not suitable for parenteral administration due to renal toxicity, whereas among CD derivatives, methylated CDs as surface-active ones cause significant hemolysis (Muankaew and Loftsson, 2018). HPβCD, sulfobutylether-β-cyclodextrin (SBEβCD) sodium salt, γCD, and 2-hydroxypropyl-γ-cyclodextrin (HPγCD) are considered safe even for parenteral administration, displaying low volume of distribution and rapid clearance in intact form by glomerular filtration. Hydrophilic CDs like HPβCD and SBEβCD are generally nonirritant and safe for topical application, whereas RAMEB is only well tolerated in low concentrations (B5%) (Duchêne and Bochot, 2016). The regulatory status of CDs is developing continuously. Natural αCD, βCD, and γCD are introduced on the generally regarded as safe list of Food and Drug Administration (FDA) for the use as food additives and along with HPβCD have monographs in European Pharmacopeia and United States Pharmacopeia and the National Formulary (USP/NF), whereas SBEβCD has monograph only in USP/NF. βCD is approved in European Union as food additive (E459) with an acceptable daily intake of 5 mg/kg/day (Kurkov and Loftsson, 2013; European Medicines Agency, 2017). Nowadays, CDs are considered as excipients and not as part of the drug substance, thus βCD, HPβCD, SBEβCD, γCD, and HPγCD are introduced into FDA’s list of inactive pharmaceutical ingredient (“Inactive Ingredient Search for Approved Drug Products,” 2014; Jansook et al., 2018).

2.2 Cyclodextrin inclusion complexes—formation, stability, and application in drug delivery Inclusion complex formation occurs spontaneously in the aqueous solution driven by the displacement of energy rich water molecules from the lipophilic CD central cavity caused by the entrapment of poorly soluble drug molecule or more often, sterically compatible lipophilic functional moiety of the drug (Fig. 2.2). During the inclusion complex formation, a whole set of intermolecular interactions, including hydrogen bonding, van der Waals forces, and hydrophobic interactions as well as ionic interactions (in case of charged drugs and CD derivatives), sterically stabilize the complex formed (Crini, 2014; Kurkov and Loftsson, 2013). No covalent bonds are formed or broken during the inclusion complex formation. This is a reversible process characterized by dynamic equilibrium where drug/CD complexes are continuously being formed and dissociated (Fig. 2.2). Such equilibrium is described by inclusion complex stability constant (Ks), with values in range from 2 M21 (fenofibrate) up to

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Figure 2.2 Mechanism of the inclusion complex formation in the solution. There is a rapid equilibrium between free and drug in the complex defined by the stability constant (Ks). Formed complexes could further associate forming aggregates ranging from nano- to microparticulate ones.

40,000 M21 (telmisartan) (Kurkov et al., 2012). In most cases, one drug molecule is included in one CD molecule thus presenting 1:1 molar stoichiometry, but in some cases, higher order complexes were detected, where one drug molecule in included into two or more CDs (Saokham et al., 2018). Natural CDs and especially γCD have tendency to self-assemble in concentrationdependent manner, forming aggregates in an aqueous solution (Loftsson and Stefánsson, 2017; Saokham et al., 2016). Aggregation is further enhanced by inclusion complex formation and then chemically modified CDs, which alone show low affinity for aggregation, self-assemble upon inclusion complex formation with lipophilic drugs (Fig. 2.2), due to surfactant-type structure of such complexes (Messner et al., 2010). The formation of aggregates should be carefully considered, especially when developing liquid formulations containing high CD concentration, as they could affect the physiochemical properties an in vivo performance of the product. Inclusion complex formation could also occur in the solid state, caused by mechanochemical activation during grinding in high-energy mills, but exact mechanism behind this solid state interactions remains unclear ( Jug and Mura, 2018). The release of the drug from the inclusion complex occurs mainly by simple media dilution (i.e., after intravenous or oral application), but other mechanisms like drug to protein binding, competitive drug displacement from CD cavity caused by bile salt complexation in the gut, and direct drug partitioning from the complex to the tissue can further enhance the release of the included drug molecule. All these contribute to rapid and complete release of the drug upon parenteral and oral administration. However, in topical applications where media dilution is limited, CDs can hamper drug release and absorption, thus the use of excessive CD concentration in such formulations should be avoided ( Jansook et al., 2018).

Cyclodextrin-based drug delivery systems

From a pharmaceutical standpoint, CD complexation represents a valuable multifunctional technology able to increase solubility, dissolution rate, chemical stability, and bioavailability of different drugs through the inclusion complex formation, thereby increasing their therapeutic potential and efficiency. Moreover, CD complexation could reduce or prevent irritation and other side effects of the drugs, enhance drug permeation across the biological membranes, prevent drug-to-drug and drug-to-excipient interactions, favorably modify organoleptic characteristics of included molecule, and even convert liquid and volatile compounds to technologically more acceptable free-flowing powders (Arima et al., 2012; Jansook et al., 2018; Kurkov and Loftsson, 2013; Loftsson and Brewster, 2011, 2010; Marques, 2010; Popielec and Loftsson, 2017). Such multifunctionality makes CDs appealing excipients for the use in the development of pharmaceuticals, including the reformulation of the existing drug product and even providing alternative application route for a given drug (Muankaew and Loftsson, 2018). Nowadays the focus of scientific research is directed toward the application of CD in development of different advanced drug delivery formulations, ranging from monolithic to micro- and nanoparticulates, with the aim to target and control the drug release in accordance with the therapeutic needs (Adeoye and Cabral-Marques, 2017; Gref and Duchêne, 2012; Salústio et al., 2011; Zhang and Ma, 2013). In some cases, addition of a third component, such as polymers, organic acids, metal ions, or lipids could additionally enhance CD efficiency through a ternary complex formation (Kurkov and Loftsson, 2013). At the moment, CD technology has been applied in the development of at least 49 pharmaceutical products available worldwide while numerous products are in advanced phases of clinical testing and will be in focus of this chapter (Wen et al., 2015).

2.3 Cyclodextrin-based products in clinical practice 2.3.1 Cyclodextrins as multifunctional excipients in dosage form design A recent comparative analysis showed that βCD is the most commonly used derivative in nonparenteral marketed products, due to its accessibility and low price. Substituted CDs are presented in approximately one-third of the products on the market and include mostly HPβCD, HPγCD, and SBEβCD, providing an excellent platform for the development of injectable formulation with improved efficiency (Kurkov and Loftsson, 2013). In addition to injectables, the application of CDs is particularly interesting in development of ocular and nasal formulations, where the drug dose must be dissolved in very limited volume of the aqueous solution ( Jansook et al., 2018). CDs as biocompatible compounds could provide efficient

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drug solubilization and chemical stabilization thereby replacing potentially irritant or toxic excipients like cosolvents, surfactants, and antioxidants (Crini, 2014; Loftsson, 2017; Popielec and Loftsson, 2017). 2.3.1.1 Cyclodextrins in parenteral formulations Currently available parenteral CD-based formulations in the clinical practice were developed using HPβCD and SBEβCD. Both derivatives have well-documented parenteral safety, presenting relatively small volume of distribution after parenteral administration (VD  0.2 L/kg) and are rapidly cleared by glomerular filtration with t / ranging from 84 to 114 minutes, appearing in the urine unmetabolized (Luke et al., 2010; Stella and He, 2008). The small VD and short t / indicate that those CDs are mainly located in the blood plasma and are cleared at the rate that is close to that of glomerular filtration without tubular resorption (Loftsson et al., 2016). When drug/ CD aqueous solutions are administered parenterally, they rapidly mix with blood plasma, leading to the rapid drug release from the complex caused by dilution and competitive displacement due to binding of the drug to the plasma proteins, while free CDs are excreted by glomerular filtration. Because of that, HPβCD and SBEβCD that are predominantly used in the development of parenteral formulations, have in general little effect on the pharmacokinetic of complexed drug, at least case of complexes with Ks lower than 105 M21 (Kurkov et al., 2010). As elimination of CDs upon parenteral administration strongly depends on renal clearance, the accumulation of HPβCD and SBEβCD could occur in patients with renal insufficiency. A study comparing the pharmacokinetic of SBEβCD in patients with normal and moderate impairment, showed that subject with moderate renal insufficiency had higher total exposure to SBEβCD, but no signs of SBEβCD toxicity were observed (Hoover et al., 2018; Luke et al., 2010). Furthermore, the pharmacokinetics of voriconazole solubilized with SBEβCD was unaffected by the degree of renal impairment in treated subject and most patients with moderate renal insufficiency were able to tolerate 7 days of intravenous therapy with such formulation (Abel et al., 2008). A study in patients with end-stage renal failure revealed that SBEβCD could be extensively and rapidly eliminated by hemodialysis, with t / similar to that in healthy individuals, whereas voriconazole, as active compound of the formulation, was poorly eliminated. Despite the rapid elimination of SBEβCD by hemodialysis using high-flux membranes, the SBEβCD exposure remained still considerably higher after repeated doses, exceeding that observed in patients with normal renal function by factor 6.2 (Hafner et al., 2010). Whether such high exposure to SBEβCD could be associated with clinically relevant toxicity in unknown and because of that, intravenous voriconazole is not recommended in patients with creatinine clearance ,50 mL/min to avoid potentially toxic accumulation of SBEβCD, used as a carrier to solubilize this poorly soluble drug (Luke et al., 2010). However, a recent study demonstrated that continuous 1

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venovenous hemofiltration could effectively remove SBEβCD, without any clinically significant clearance of voriconazole, providing a possibility to use standard recommended dosages of intravenous voriconazole to treat susceptible fungal pathogens in critically ill patients. However, higher drug doses may be required for less susceptible pathogens (Kiser et al., 2015). Nowadays voriconazole 200 mg sterile powder for intravenous solution is available as a product developed with SBEβCD (VFEND I.V. for Injection, Pfizer, USA) and the one developed using HPβCD (Panpharma, UK) (VFEND voriconazole for injection, for intravenous use, n.d.; Voriconazole 200mg Powder for Solution for Infusion - Summary of Product Characteristics (SmPC) (eMC), n.d.). Injectable formulation of itraconazole was developed using 40% HPβCD aqueous solution and became clinically available in 1999 as Sporanox (Janssen Pharmaceutica, Belgium), providing the advance of achieving adequate drug levels in blood more rapidly and with less patient-to-patient variability than those obtained by orally administered preparation of the drug. Intravenous HPβCD itraconazole solution can be safely administered to infants beyond 6 months, children, and adolescents using a weightnormalized approach to dosing (Abdel-Rahman et al., 2007). However, due to HPβCD accumulation in renally impaired population, the administration of itraconazole intravenous preparation was limited in patients with a creatinine clearance of # 30 mL/min. Hemodialysis significantly removes the HPβCD carrier, without any significant effect on pharmacokinetics of itraconazole and its active metabolite hydroxyitraconazole, supporting a predialysis administration strategy for intravenous itraconazole in renally impaired patients (Mohr et al., 2004). A posaconazole intravenous formulation was recently developed as an aqueous solution containing SBEβCD as the solubilizer, present in similar concentration as in voriconazole i.v. formulation. Currently available formulations are only for oral administration (Noxafil, posaconazole oral suspension and tablet, Merck & Co., USA) therefore such new SBEβCD-based formulation provided the unmet need for an intravenous formulation that would allow the treatment of patients with the risk of invasive fungal infection, that are not able to take the drug orally. Posaconazole intravenous solution was well tolerated in subjects at high risk for invasive fungal infection and pharmacokinetic analysis showed that posaconazole at the dose of 300 mg administered intravenously can achieve targeted plasma concentration of the drug that are associated with positive therapeutic outcome (Maertens et al., 2014). Based on that, intravenous posaconazole 300 mg once daily was selected as the dose to use in the phase 3 study for further investigation in a larger, more diverse patient population. Such dosage regimen was well tolerated and resulted in adequate steady-state systemic exposure in patients at risk of invasive fungal infection, with 94% of patient’s population achieving the prespecified exposure target of cavg $ 500 and # 2500 ng/mL. Posaconazole plasma levels decreased after patients transitioned from intravenous

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treatment to dosing with posaconazole oral suspension, regardless of the dose of suspension or regimen, however, the clinical significance of this effect is still to be evaluated (Cornely et al., 2017). Diclofenac is one of the most prescribed NSAIDs worldwide (Scavone et al., 2016). The HPβCD complexation of diclofenac increased drug solubility approximately 7.5 times, enabling the development of an aqueous drug solution that is stable at the room temperature and is suitable for subcutaneous, intramuscular, and intravenous administration. The previously available diclofenac parenteral formulation, as Voltarol Ampoules (Novartis, UK), was developed using polyethylene glycol and benzoyl alcohol as cosolvent and contained 75 mg of the drug in 3 mL of the formulation. Such solution could be readily administered as intramuscular injection or should be diluted to 50 or 100 mL with either sodium chloride solution (0.9%) or glucose solution (5%), buffered with sodium bicarbonate and then administered as 30-minute intravenous infusion. The benefit of HPβCD-based formulation is in the fact that it contains the same drug dose in reduced volume of the formulation (75 mg in 2 or 1 mL) that could be readily administered as a fast-intravenous bolus injection, providing the clinical advantage in producing analgesia in significantly shorter time frame with the same duration as observed for Voltarol (Mermelstein et al., 2013). Another benefit of HPβCD parenteral solution over conventional formulation is in significantly lower incidence of thrombophlebitis (1.2% vs 6.8%, respectively) (Colucci et al., 2009). The bioavailability of HPβCD diclofenac and Voltarol are equivalent, regardless the administration route (intramuscular or intravenous). As diclofenac HPβCD solution is administered as an intravenous bolus injection, it results in 3.8 times higher cmax attained in 10 times shorter time period (0.05 and 0.5 hours) than that after Voltarol 30-minute intravenous infusion. This difference may contribute to the clinical observation of a more rapid onset of analgesia for HPβCD-diclofenac than Voltarol. Furthermore, pharmacokinetic parameters after single and multiple doses were comparable. The lack of accumulation and linear pharmacokinetics of HPβCD diclofenac were also demonstrated, which could potentially provide added benefits in patients with complex analgesic regimens or receiving multimodal analgesia (Mermelstein et al., 2013). Intravenous HPβCD diclofenac was not associated with increased risk of treatment-emergent cardiovascular adverse effects when given for less than 5 days postoperatively to abdominal/pelvic and orthopedic surgery patients (Gan et al., 2016). Furthermore, the relatively brief exposure of HPβCD diclofenac, typical for parenterally administered analgesics for acute postoperative pain management, in patients with normal renal function is not associated with added renal safety risks over placebo (Daniels et al., 2016). Mild-to-moderate renal or mild hepatic insufficiency did not significantly affect the exposure to or elimination of diclofenac following intravenous administration of a single dose of HPβCD diclofenac solution. Although in patients with mild-to-moderate renal impairment, the HPβCD clearance is decreased,

Cyclodextrin-based drug delivery systems

the product contains approximately 24 times lower concentrations of HPβCD than voriconazole intravenous solution, therefore, the HPβCD plasma concentrations achieved in this patient group was far below levels associated with adverse effects. This indicates that HPβCD diclofenac may be administered to patients with mild or moderate renal insufficiency or mild hepatic impairment at the usual dose and schedule without a need for dose reduction (Hamilton et al., 2018). In the United States, HPβCD diclofenac (Dyloject, Pfizer, USA) is approved in adults for the management of mild-to-moderate pain, and as monotherapy or in combination with opioid analgesics for the management of moderate to severe pain. The recommended dose is 37.5 mg, administered as an intravenous bolus over 15 seconds every 6 hours as needed, whereas the total daily dose should not exceed 150 mg (Hoy, 2016). In European countries, HPβCD diclofenac is available as Akis (Flynn Pharma, Ireland), containing 75 mg of the drug per 1 mL, suitable for intramuscular and subcutaneous administration in treatment of acute forms of pain. The same formulation could be administered as direct intravenous bolus injection, to treat postoperative pain in hospital settings. Another EU approved HPβCD diclofenac parenteral formulation is Dicloin solution for injection (IBSA, Slovakia), available in 25, 50, and 75 mg/1 mL doses that is indicated only for subcutaneous or intramuscular administration (Keating et al., 2016). A recent clinical report indicated that Akis/Dicloin could be administered as a local submucosal injection prior to lower third molar surgery, following achievement of local anesthesia. Administered in doses ranging from 5 to 50 mg, the formulation produced significantly superior effects compared with placebo in preventing pain during the 6-hour postsurgical observation period in terms of pain intensity levels, time until pain onset, and time to first rescue medicine intake. Interestingly, the analgesic effect of 5-mg dose appeared superior to that obtained with 12.6- and 25-mg drug doses, providing the novel concept of low-dose local analgesia (Gorecki et al., 2018). CD complexation enabled the development of parenteral formulations of numerous drugs, that previously were not available in such type of formulation. The inclusion complexation of progesterone with HPβCD enabled the development of stable pharmaceutical product, suitable for intramuscular or subcutaneous administration, indicated as luteal phase supporting during in vivo fertilization in women who are unable to use vaginal preparation. Nowadays several registered HPβCD progesterone products are available (Lubion, Pleyris, and Prolutex), representing valid alternative to existing formulation, enabling easy self-administration and good tolerability with pharmacokinetic profile, clinical efficiency and safety profile comparable to that of locally acting vaginal products, like progesterone vaginal cream or progesterone vaginal insert (Scavone et al., 2016). Increased aqueous solubility of ziprazosine mesylate obtained through inclusion complexation with SBEβCD enabled the development of the drug parenteral formulation (Geodon for injection, Pfizer, USA). By that, ziprazosine becomes a

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first atypical antipsychotic to be clinically available in both intramuscular and oral formulation. The intramuscular formulation is present as a lyophilized powder to be reconstituted with sterile water and is indicated to treat occurrences of severe agitation in patients with schizophrenia, that may pose a safety problem for patients and its surroundings. Once an acute episode is controlled, patients can be safely moved to maintenance therapy with oral ziprasidone (as monohydrochloride) (Preskorn, 2005). Carbamazepine, an oral antiepileptic drug prescribed worldwide as a first-line treatment for partial seizures, is insoluble in water and currently there are no intravenous replacement formulation available. Complexation with SBEβCD enhanced the drug solubility, permitting the preparation of an intravenous carbamazepine formulation. Intravenous SBEβCD carbamazepine solution, administered over 30 minutes at the dose equivalent to 70% of the usual oral dose, appeared bioequivalent to oral carbamazepine in patients with normal renal function. When such infusion were administered every 6 hours, the carbamazepine plasma concentration were maintained in the therapeutically efficient range (Tolbert et al., 2015). Intravenous SBEβCD carbamazepine administered as multiple 30- or 15-minutes infusions every 6 hours or as a single rapid infusion were well tolerated. Infusion site reactions were mild, and the seizure control was generally unchanged when patients were switched between oral and intravenous carbamazepine. The new drug application for this product is currently under review by the FDA (Lee et al., 2015). SBEβCD complexation of the neuroactive steroid allopregnanolone (Brexanolone, SAGE Therapeutics, USA) enabled the development of an intravenous solution able to produce rapid therapeutic actions in women suffering from postpartum depression, that has potential to improve treatment options for women with this disorder (MeltzerBrody et al., 2018). Solubilization by SBEβCD was also used to develop intravenous topiramate formulation that could be used in the treatment of neonatal seizures or to provide bridge therapy to patient whom oral topiramate was interrupted. Furthermore, the use of such injectable, stable formulation given simultaneously to patients on maintenance topiramate therapy enabled for the first time determination of oral topiramate bioavailability (Clark et al., 2013). Letermovir is a novel antiviral agent in clinical development aimed for prophylactic treatment against human citromegalovirus in immunocompromised transplant patient. To aid the administration of letermovir, a novel intravenous drug formulation has been developed using HPβCD. Such formulation would enable the start of prophylaxis immediately after transplantation. Intravenous HPβCD letermovir solution was well tolerated in tested subjects and determined pharmacokinetic parameters supported a once-daily dosing regimen (Erb-Zohar et al., 2017) In some cases, the application of CD drug solubilization enabled the development of parenteral drug formulations without the use of potentially irritant cosolvents and surface-active agents. Hypotension is the dose-limiting adverse event in intravenous amiodarone and is attributed to the polysorbate 80 and benzyl alcohol, used as

Cyclodextrin-based drug delivery systems

cosolvents in the formulation. To minimize hypotension, the initial dose of amiodarone should be diluted with sterile 5% dextrose solution and slowly infused over 10 minutes. Cosolvent-free intravenous amiodarone solution was developed using SBEβCD. Clinical evaluation of such formulation demonstrated its bioequivalence with cosolvent-based formulation, without causing hypotension in healthy human subjects (Cushing et al., 2012, 2009). Such formulation is nowadays clinically available as Nexterone Injection for intravenous use (Baxter, USA). Alphaxalone solubilized with Cremophor EL (Althesin, Glaxo Laboratories, UK) was used in clinical practice for the induction and maintenance of anesthesia from 1972 to 1984 but was withdrawn from the marked because of hypersensitivity reactions caused by Cremophor EL. Nowadays, intravenous alphaxalone solution 10 mg/ mL was developed using SBEβCD (Phaxan, Drawbridge Pharmaceuticals, Australia). Such formulation caused fast-onset, short duration anesthesia with fast cognitive recovery like that obtained with propofol, but with less cardiovascular depression and no airway obstruction (Monagle et al., 2015). Autologous stem cell transplantation after high-dose melphalan conditioning is considered as a standard of care procedure for patients with multiple myeloma. Propylene glycol-free melphalan hydrochloride intravenous solution (Evomela, Spectrum Pharmaceuticals, USA) was developed using SBEβCD, to avoid the risk of propylene glycol toxicity and to increase drug chemical stability upon reconstitution. Evomela was confirmed to be equivalent to Alkeran (GlaxoSmithKline, USA), a polyethylene glycol-based formulation. Substitution of propylene glycol by SBEβCD in the formulation provided the mean to safely administer higher drug doses and eliminated the time constrain imposed on pharmacists and nursing staff when preparing and administering the drug, caused by chemical instability of the drug in polyethylene glycol-based formulation. Evomela allows administration of high-dose melphalan infusion 8 hours upon reconstitution, resulting in more reliable delivery of planned drug dose with acceptable safety and high efficiency (Hari et al., 2015). Camptothecin is a topoisomerase 1 inhibitor acting as a potent broad-spectrum anticancer agent. However, its clinical development and application was hampered by poor solubility (B4 μg/mL), insufficient stability at physiological conditions due to spontaneous hydrolysis of lactone ring and considerable toxicity. In an attempt to overcome these obstacles, a polymeric nanoparticle consisting of β-CD-poly(ethylene glycol) copolymer conjugated to camptothecin (CRLX101) has been developed. The backbone of this biocompatible and nontoxic copolymer is built of alternately repeating βCD and PEG units, whereas camptothecin is derivatized at the 20-OH position with glycine to form an ester linkage for covalent attachment to copolymer backbone. Inter- and intramolecular inclusion complex formation between camptothecin and βCD units lead to selfassembly of several polymer strands into nanoparticles of 20 40 nm in diameter. The PEG blocks contribute to neutral surface charge, improved solubility, and stealth

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properties of the formed nanoparticles, enabling their safe systemic administration, reduced immunogenicity, and avoidance of phagocytosis by the reticuloendothelial system. Consequently, CRLX101 exhibits extended plasma stability and prolonged circulation time, passively accumulating in the targeted tumor tissue through enhanced permeability and retention effect. The chemical linkage of camptothecin increases the drug solubility by three orders of magnitude and stabilizes the lactone ring, preventing its premature pH-mediated inactivation upon systemic administration. Furthermore, the drug is also shielded against metabolic enzymes inside the formed nanoparticles, being released in a controlled manner over a sustained period. The drug release is mediated through the cleavage of the glycine linker by both enzymatic and base-catalyzed hydrolysis of the ester bond, with t / in human plasma of 41 hours. The drug release leads to disassembly of the nanoparticles into individual polymer strands with less than 10 nm in diameter that are cleared from the body by glomerular filtration in the kidney (Davis, 2009; Svenson et al., 2011). In preclinical development, CRLX101 was shown to be highly effective in a variety of tumor tissues and multiple human tumor xenograft models that respond poorly to irinotecan or topotecan treatment. First-in-human phase 1/2a trial of CRLX101 enrolling patients with advanced solid malignancies demonstrated encouraging safety, pharmacokinetics, and efficiency. The maximum tolerated dose was determined at 15 mg/m2 biweekly, that was generally well tolerated with neutropenia and fatigue as the most common adverse effects. Stabile disease was observed in 64% of patients treated with such dose of which 34% had confirmed stable disease at subsequent evaluations (Weiss et al., 2013). The selective accumulation of CRLX101 in gastric tumors and not in adjacent non-neoplastic tissue was demonstrated in humans, whereas carbonic anhydrase IX and topoisomerase I expressions in tumor tissues after treatment provided clear evidence of biological activity of delivered camptothecin (Clark et al., 2016). Furthermore, CRLX101 demonstrates a promising antitumor activity in heavily pretreated or treatment-refractory solid tumor malignancies presumably by the inhibition of proliferation and angiogenesis correlating with tumor growth inhibition (Gaur et al., 2014). However, despite such promising efficiency, a randomized phase II trial of CRLX101 in combination with bevacizumab versus standard of care in patients with advanced renal cell carcinoma, although generally well tolerated, did not demonstrate actual improvement in progression free survival when compared with approved agents (3.7 vs 3.9 months, respectively) (Voss et al., 2017). Future clinical trials of CRLC101 will be focused on topoisomerase 1—sensitive tumors such as ovarian and gastrointestinal ones. 1

2

2.3.1.2 Cyclodextrins in ocular formulations Diclofenac sodium 0.1% ophthalmic solution is widely distributed in Europe and the United States under several brand names and in a variety of formulations, as efficacious topical ocular antiinflammatory agent. It presents a wide range of therapeutic

Cyclodextrin-based drug delivery systems

indications, including inhibition of preoperative miosis during cataract surgery as well as the treatment of postoperative inflammation and discomfort in cataract surgery, argon laser trabeculoplasty, strabismus surgery or after accidental nonpenetrating trauma. Available formulations in clinical practice are developed using various excipients that play a major role in the product safety profile. One of the available formulations is developed using HPγCD (Voltaren Ophta CD, Novartis, and Voltarol Ophtha, Thea Pharmaceuticals), providing stabile solution without the use of potentially irritant surfactants like macrogolglycerol ricinoleat and polyoxyl 35 castrol oil (Bodaghi, 2008). Furthermore, the use of HPγCD allowed formulation preservation with benzalkonium chloride instead of thiomersal. Clinical trials confirmed that CDbased formulation was as effective as standard one formulated with surfactants, in the treatment of inflammation after cataract intraocular lens surgery with an additional benefit in lower incidence of ocular discomfort upon application (Mester et al., 2002). Several years ago, a new 0.005% latanoprost ophthalmic solution containing CDs was developed and introduced in clinical practice in several countries in Latin America. The efficiency and safety of latanoprost in reducing intraocular pressure are well established, however, the drug is unstable in aqueous solution at temperatures above 4 C. A new 0.005% latanoprost ophthalmic solution containing CDs (GAAP Ofteno, Laboratorios Sophia) exhibits a comparable efficacy and safety profile as the standard formulation, however, improves the stability of the active agent and thus does not require refrigeration. The use of CDs allowed storage at temperatures up to 40 C for at least 3 months and storage at 25 C for at least 24 months without any significant loss in the content of the active agent (Gonzalez et al., 2007). Nowadays different novel CD derivatives are screened as potential carriers for latanoprost, and among them propylamino-βCD appeared to be the optimal, providing improved stability, solubility, and ocular tolerance of the drug, with lower incidence of ocular irritation on rabbit model than the commercial latanoprost formulation (Xalatan, Pfizer), having a great potential for further clinical development (Rodriguez-Aller et al., 2015). Topical corticosteroids are lipophilic compounds that are only soluble to a limited extent in aqueous eye drop formulation and are frequently administered as ophthalmic suspensions, such as 0.1% dexamethasone alcohol suspension, or as hydrophilic watersoluble prodrugs, such as 0.1% dexamethasone sodium phosphate, both showing low ocular availability. HPβCD at concentration of 5% and 10%, in the presence of 0.1% of Hypromellose, allowed the preparation of stabile dexamethasone ophthalmic solutions containing 0.3% and 0.7% drug dose. Application of such formulation resulted in significantly higher dexamethasone concentrations in aqueous humor than dexamethasone alcohol suspensions (Kristinsson et al., 1996). Furthermore, clinical trial on patients undergoing cataract surgery showed that 0.7% dexamethasone CD ophthalmic solution applied once daily is a more effective postoperative antiinflammatory

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medication than 0.1% dexamethasone sodium phosphate applied three times a day (Saari et al., 2006). There are some evidences that formation of drug/CD complex aggregates might have important role in enhancing the drug bioavailability (Crini, 2014). Furthermore, through stepwise optimization of the formulation, inclusion complex aggregation could be precisely controlled, providing preparations containing solid drug/CD microparticles, nanoparticles, or dissolved drug. Such systems release the drug in sustained manner, providing locally high drug concentrations in the aqueous tear fluid for several hours. This enhances drug transcorneal permeation and enables the efficient therapy of diseases affecting the posterior segment of the eye (Loftsson and Stefánsson, 2017). A very illustrative example in this regard is a recent study evaluating the safety and efficacy of topical 1.5% dexamethasone aqueous eye drops with CD microparticles for diabetic macular edema (Tanito et al., 2011). The aqueous dexamethasone ophthalmic microsuspension was prepared by autoclaving the drug suspension in the presence of γCD and poloxamer. The microsuspension was preserved using ethylenediaminetetraacetic acid and benzalkonium chloride, whereas product osmolarity was regulated through sodium chloride addition. In such formulation, approximately 85% of the drug was present as dexamethasone/γCD microparticles with a diameter between 1 and 3 μm, whereas the rest of the drug was dissolved in formulation. A prospective pilot study showed that such formulation, applied three times a day, can reduce central macular thickness for more than 10%, clearly demonstrating the potential of such topical formulation to deliver therapeutically significant drug doses to the human retina. However, it seems that the ability of CD micro/ nanoparticles to ensure high-local drug concentration in the tear fluid and efficiently deliver the drug to the posterior segment of the eye is highly dependent of the physiochemical properties of the drug in question. In the γCD nanoparticle drug delivery platform, the kinetics of dexamethasone and dorzolamide in the tear film is completely different. Dexamethasone complexed with γCD nanoparticles has up to 30-fold higher concentration and longer retention time in tear fluid than commercial dexamethasone eye drops (Maxidex, Novartis Pharma), whereas in case of dorzolamide formulated with γCD nanoparticles, such effect was not observed. In this case, the overall concentration dorzolamide was about two times higher than after Trusopt (Laboratories Merck Sharp & Dohme—Chibret) administration (Jõhannesson et al., 2014). Regardless to that, prospective randomized single masked crossover trial over 24 hours showed that 3% dorzolamide γCD nanoparticle ophthalmic formulation given once a day has the same efficiency in reducing intraocular pressure as standard Trusopt formulation, administered three times a day. Further benefit of γCD nanoparticle ophthalmic formulation is in better tolerability and safety profile than that of Trusopt (Gudmundsdottir et al., 2014). Topical 1.5% dexamethasone γCD nanoparticle ophthalmic formulation, due to its ability to enhance the transcorneal permeation of the drug, significantly improved visual acuity and decrease macular thickness in

Cyclodextrin-based drug delivery systems

patients with diabetic macular edema in a way comparable with that obtained by subtenon triamcinolone administration, clearly demonstrating the efficiency of such formulation in delivering the drug to the posterior segment of the eye (Ohira et al., 2015). The presented examples from clinical practice clearly reveal that CDs-based ophthalmic formulation can decrease frequency of drug administration, increase ocular tolerance of the applied drug, and even efficiently treat the disease of the posterior segment of the eye through topical drug administration, all improving patient compliance and consequently leading to more efficient treatment of ocular diseases. 2.3.1.3 Cyclodextrins in nasal formulations In 1990 the nasal application of CDs gained notable attention, as they emerged as very efficient nasal absorption enhancers of different drugs, including small lipophilic molecules like estradiol, morphine, dihydroergotamine and buserelin or even peptides and proteins such as calcitonin and insulin (Merkus et al., 1999). In this regard, dimethylβ-cyclodextrin appeared to be particularly effective; however, its high efficiency observed on animal models was not confirmed in humans. Regardless to that, even nowadays CDs-based formulations are extensively investigated for their potential to target the central nervous system or to provide enhanced systemic drug delivery following nasal administration ( Jacob and Nair, 2018). Although the results obtained on animal models are encouraging, their efficiency still needs to be confirmed by clinical trials on humans. To authors best knowledge, Aerodiol (Servier Laboratories, France) was the first and only CD-based nasal formulation in clinical practice. This aqueous 17β-estradiol nasal formulation was developed using RAMEB, providing more than 1000-fold increase in drug solubility, thereby allowing the incorporation of therapeutically relevant drug doses in a volume suitable for nasal administration. Nasal Aerodiol administration in dose of 300 μg/day provided pulsed estrogen therapy that was well tolerated and highly effective in alleviation of climacteric symptoms (Pelissier et al., 2001). However, the nasal spray had to be taken in combination with continuous or sequential progestrogen, to oppose estrogen’s effect on endometrial hyperplasia and reduce the risk of cancer (Gompel et al., 2000). Although the Aerodiol seems to be well tolerated and accepted by the patients, Servier Laboratories discontinued its manufacturing and marketing in 2006 for commercial reasons (Aerodiol nasal spray, n.d.). Nasal administration of midazolam can be used for sedation before surgical, dental, or diagnostic procedures and for treatment of seizures in pediatric and adult patients. However, the low solubility of the drug hampers the development of suitable nasal formulation. Application of 12% (w/w) of RAMEB allowed the preparation of isotonic nasal solution containing 30 mg/mL of midazolam. Such approach provided the formulation containing higher drug dose compared with that prepared using propylene

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glycol (27.8 mg/mL) or propylene glycol and PEG 400 as cosolvents (25 mg/mL). Further drawback of cosolvent-based formulations was in their hyperosmolarity that caused nasal tissue irritation upon administration. An open-label, sequential trial demonstrated that RAMEB-based formulation administered nasally at 3 mg dose was well tolerated and presented absolute bioavailability of 85 6 8% with tmax and cmax of 13 minutes and 68.9 ng/mL, respectively. For comparison, tmax and cmax for intravenously applied midazolam at 1 mg dose was 2.1 minutes and 87.6 ng/mL, respectively. Addition of chitosan (0.5% w/w) as an absorption enhancer provided decrease in tmax and increase in cmax (7.2 minutes and 80.6 ng/mL, respectively), but the overall bioavailability was somewhat reduced (76% 6 12%) with comparison to that of chitosan free formulation (Haschke et al., 2010). Although the maximal serum concentrations of midazolam after nasal administration of RAMEB formulation are not reached as quickly as after intravenous injection, such formulation could provide a mean of noninvasive drug application in case of seizures and has a potential for further clinical development. Another example of clinically successful nasal formulation development is the fixed combination product containing budesonide solubilized with SBEβCD and azelastine hydrochloride. Clinical practice showed that concomitant antihistamine and corticosteroid nasal administration are more efficient than either medication alone in the treatment of seasonal allergic rhinitis. The randomized, double-blind, placebo-controlled, single-center, three-treatment, six-sequence, and three-period crossover clinical study demonstrated that such combined CD-based formulation provides fast and long acting relief of allergic rhinitis symptom that is superior to that obtained by sequential nasal administration of budesonide suspension (Rhinocort Aqua, Astra Zeneca) and azelastine hydrochloride solution (Astelin, MEDA Pharmaceuticals Inc.). The advantage of combined CD-based formulation is reduced time to onset of action, superior uniformity of dosing with no need to resuspend the drug before use, providing the convenience of dosing with both drugs by one nasal spray product. All these supports the further development of such CD-based formulation toward registration and introduction into everyday clinical practice (Salapatek et al., 2011). Proteins and peptides are nowadays rapidly growing class of active pharmaceutical ingredients, presenting almost 30% of newly approved drugs by the United States Food and Drug Administration in 2015 18 period. However, such compounds cannot be administered orally as they are highly susceptible to degradation in the gastrointestinal tract and are poorly absorbed across the intestinal epithelium because of their high molecular mass and hydrophilicity (Anselmo et al., 2018). Nasal administration appears as the promising strategy for efficient noninvasive delivery of peptide and protein drugs, at least those that are therapeutically active at low doses, such as desmopressin, oxytocin, nafarelin, and calcitonin (Al Bakri et al., 2018). However, until now the potential of nasal delivery of peptide and protein drug is not fully exploited. Although

Cyclodextrin-based drug delivery systems

nasally administered glucagon can counteract insulin-induced severe hypoglycemia in diabetic patients, a potentially life-threatening complication accompanying everyday therapy of diabetes mellitus, the currently available glucagon emergency kits is developed as intramuscular injection. Moreover, this formulation contains glucagon in the form of lyophilized powder that requires a reconstitution with provided diluent immediately prior the administration, due to instability of glucagon in the aqueous solution. Therefore, there is an unmet need for needle-free, ready-to-use formulation of glucagon, as an alternative to injectable glucagon for caregivers, family members, friends, and colleagues who may someday face the difficult task of treating severe hypoglycemia in insulin-using children or adults. In 2010 Eli Lilly company launched development of new nasal glucagon formulation for treatment of severe hypoglycemia. The novel glucagon nasal formulation is developed as a dry powder-containing βCD and dodecylphosphocholine as absorption enhancer that is administered by simple nasal applicator (Pontiroli and Ceriani, 2018). Clinical studies in controlled environment demonstrated that single nasal dose of glucagon (3 mg) can be used in children and adults to treat insulin-induced hypoglycemia, with a glycemic response similar to that of weight-based intramuscular glucagon (0.5 or 1 mg) (Rickels et al., 2015; Sherr et al., 2016). The ease of use and efficacy of intranasal glucagon has also been demonstrated in clinical studies in which intranasal glucagon was dispensed to adults with diabetes type 1 or to caregivers of children with diabetes type 1 to treat incident events of hypoglycemia in the home and school settings. Over 90% of caregivers reported that they were very satisfied with the ease of use of the product (Pontiroli and Ceriani, 2018), so there is a high potential that such needle-free, CD-based nasal glucagon formulation would be commercialized in the near future. 2.3.1.4 Cyclodextrins in oral formulations CDs have an enormous potential in development of oral formulations. In liquid formulations, CDs can be used to solubilize the drug and increase its chemical stability in the aqueous solution, masking in the same time the unpleasant drug taste (Loftsson and Brewster, 2010). Itraconazole oral solution 10 mg/mL was developed using HPβCD. The pharmacokinetic study demonstrated that such solution, when administered in fasted state provided oral bioavailability of itraconazole and its active metabolite that is 30% 33% and 35% 37% higher when compared with that obtained by the itraconazole capsules administered at the same dose. Other pharmacokinetic parameters, like tmax, cmax, and t1/2 were similar for both formulations (Barone et al., 1998). Nowadays HPβCD itraconazole solution is registered as Sporanox 10 mg/mL oral solution (Janssen-Cilag, EU) and is indicated for the treatment of oral and/or esophageal candidosis in HIV-positive or other immunocompromised patients or as prophylaxis of deep fungal infections susceptible to itraconazole, when standard therapy is considered inappropriate (Sporanox 10 mg/ml Oral Solution - Summary of

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Product Characteristics (SmPC) - (eMC), n.d.). Furthermore, HPβCD itraconazole at a dosage of 2.5 mg/kg twice a day was well tolerated and efficacious for the treatment of oropharyngeal candidiasis in HIV-infected pediatric older than 5 years, and even patients with fluconazole-resistant isolates responded well to such treatment (Groll et al., 2002). It is generally anticipated that inclusion complex formation can improve oral absorption of the drugs of class II and IV according to the biopharmaceutical classification system, with an average increase of the oral bioavailability of 150%, depending on the CD derivative used, stability constant of the complex formed and drug to CD molar ratio in the tested product (Loftsson et al., 2016). However, in some cases, such effect could be absent, as the volume for dissolution/dilution in the gastrointestinal tract is relatively low, so the release of the drug from the complex and its absorption can be hampered in case of highly stabile complexes or the use of excess CD in the formulation. Finally, the effects of CD complexation will have negligible or even negative effect in case of class III and class I drugs, respectively (Loftsson et al., 2016). Inclusion complexation of piroxicam, a well-established nonsteroidal antiinflammatory drug used worldwide in the treatment of musculoskeletal diseases, in 1:2.5 drug to βCD ratio lead to the formation of amorphous, hydrophilic, and readily wettable compound which dissolves rapidly. Such complex is nowadays produced by supercritical CO2 technology, avoiding completely the use of potentially toxic solvents (Scarpignato, 2013). Clinical studies have shown that the increased aqueous solubility and dissolution rate of the complex lead to more rapid drug absorption rate with 61% higher ka compared with that of piroxicam tablet containing the drug in the free form (Feldene, Pfizer, SAD), retaining all the analgesic and antiinflammatory properties of the parent compound (Deroubaix et al., 1995). Preclinical studies on rabbit model showed that the increase of piroxicam absorption rate is proportional to the amount of the βCD used in the formulation, increasing up to 1:2.5 drug to βCD molar ratio, whereas further increase in the βCD content (i.e., 1:3 and 1:4 molar ratio) in turn decreases its absorption (Skiba et al., 2013). Because of improved drug solubility and dissolution upon complexation with βCD, the therapeutic drug concentration in systemic circulation are reached in 30 minutes after oral administration of piroxicam/ βCD complex, whereas for conventional oral tablets of piroxicam, this is achieved after 2 hours. However, the overall oral bioavailability for both formulations is the same (Wang et al., 2000). In addition to the more rapid onset of the analgesic effect, another benefit of βCD complexation of piroxicam is in its improved gastrointestinal tolerability. The piroxicam gastrointestinal toxicity is caused by both the nonselective inhibition of cyclooxygenase activity in the gastric mucosa and prolonged contact of poorly soluble acidic drug with the gastric mucosa upon oral administration. The improvement of piroxicam solubility and dissolution rate achieved through inclusion complexation with βCD reduces the contact time and consequent topical irritation of

Cyclodextrin-based drug delivery systems

the gastrointestinal mucosa, leading to improved gastrointestinal drug tolerability, that was demonstrated in several clinical studies (Scarpignato, 2013). Moreover, it seems that complexation with βCD and HPβCD can produce reduced gastric damage, lower incidence of gastric lesion, and the extent of ulceration associated with nimesulide, etodolac, phenybutazone, naproxen, indomethacin, and ketorolac, thereby providing a mean to improve the safety profiles of the drugs (Fioravanti et al., 2002; Scavone et al., 2016). Piroxicam/βCD was a first CD-based formulation introduced on European market in 1988 as Brexin (Chiesi Farmaceutici, Italy) (Crini, 2014) and nowadays is available as oral tablet, sachets, and effervescent formulation in several countries in European Union and Asia. Compared with regular tablet, piroxicam/ βCD effervescent formulation provides even more rapid drug absorption, with cmax being reached in 15 minutes after oral administration (Scarpignato, 2013). Clinical studies of piroxicam/βCD sachets abnormal postural sway in patients with chronic low back pain demonstrated that such formulation can provide greater improvement in pain score and disability index than those obtained with piroxicam tablets, producing lower incidence of adverse effects, thereby expanding the spectrum of analgesic effect of the drug in the treatment of chronic low back pain (Chang et al., 2008). Another study compared the efficacy and tolerability of nimesulide/βCD complex with the parent compound in patients undergoing arthroscopic surgery, clearly demonstrated that βCD tablet formulation has a more rapid onset of action than that containing nimesulide alone, reaching a statistically significant reduction in pain intensity 15 minutes after oral administration and providing an equivalent duration of analgesic effect as plain drug tablet formulation (Vizzardi et al., 1998). The very poor aqueous solubility and wettability of the drug gives rise to difficulties in the pharmaceutical development of oral formulation and leads to a variable bioavailability. However, this problem could be efficiently solved through inclusion complex formation with βCD (Singla et al., 2000). The efficiency of naproxen/βCD in symptomatic treatment of osteoarthritis appears to be equivalent to that of naproxen, used as a standard therapy, and is associated with lower incidence of gastrointestinal adverse reactions (Fioravanti et al., 2002). From limited clinical data available, it appears that nimesulide/βCD oral tablet (Nimedex, Novartis, Europe) may provide a useful and well-tolerated therapeutic alternative in cases where a fast relief of pain is required. Another example of CD-based oral formulation in clinical practice is the inclusion complex of ethinyl estradiol with βCD that is coadministered with drospirenone, which is registered worldwide as Yaz (Bayer, Germany), an oral contraceptive. Here βCD complexation is used enhance chemical stability of ethinyl estradiol and to prolong the product shelf-life. Clinical evaluation demonstrated that βCD, although efficiently improving chemical stability of ethinyl estradiol through inclusion complex formation, did not affect oral bioavailability of ethinyl estradiol and coadministered drospirenone nor altered pharmacokinetic profile of these hormones (Blode et al., 2008).

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Chewable tablets are developed to enable the oral drug administration without the use of water, providing easier swallowing of the formulation compared to standard solid oral dosage forms, that is of particular importance for pediatric and geriatric patients. Here, as in oromucosal formulations, one of the critical quality attributes is the taste and aftertaste of the formulation during and after chewing (Mistry and Batchelor, 2017). CD complexation inhibits the binding of the drugs to the taste buds receptors present on the tongue, interfering with the drug taste sensation. In general, better taste-masking efficiency will be attained with CD derivative that is able to form more stable inclusion complexes with the drug in question, thereby would more efficiently lower the concentration of the free drug in saliva that is able to interact with the taste receptor. In addition, the sweet taste of CDs also contribute to their bitter taste-masking effect (Arima et al., 2012). Examples of such formulations on the market where bitter drug taste was efficiently masked by βCD include Zyrtec (Losan Pharma, Germany), chewable tablet with cetirizine indicated for symptomatic relief of allergy symptoms and Stada-Travel (Stada, Germany), chewable tablets containing fixed combination of diphenhydramine hydrochloride and chlorotheophylline to combat motion sickness (Conceição et al., 2018). Recent literature review revealed that CD complexation has been used in the development of at least 26 different solid oral dosage forms available for the clinical use worldwide, including tablets, capsules, and chewable tablets, with βCD being the most frequently used (19/26) derivative. Other CD derivatives as αCD (2/26), HPβCD (2/26) and SBEβCD (1/26) are seldom used (Conceição et al., 2018; Kurkov and Loftsson, 2013). Majority of formulations contain highly potent drugs (i.e., the ones therapeutically effective at doses lower than 50 mg) and CDs are used to increase drug solubility and dissolution rate, improve its chemical stability, or to mask the unpleasant taste, as discussed earlier. However, the potential of CD use in the development of solid oral dosage forms goes far beyond that, as CDs can be used to modify the drug release behavior providing pH-dependent or sustained release, act as osmotic agents in osmotically controlled drug-delivery systems or provide colon drug targeting (Salústio et al., 2011). Therefore considering multifunctionality of CDs as excipients in solid oral formulations, one could expect a further increase in their number on the marketed and in clinical practice. 2.3.1.5 Cyclodextrins in oromucosal formulations Drug delivery via mucosal surfaces of the oral cavity, especially by sublingual route is nowadays a favored rote for noninvasive administration of emergency drugs, able to provide rapid onset of the drug action, circumventing the drug degradation in the gastrointestinal tract and the first-pass metabolism in liver (Brandl and Bauer-Brandl, 2019). Moreover, orally disintegrating tablets (ODTs), that rapidly disintegrate in the mouth forming drug suspension or solution, appeared recently as a relatively novel

Cyclodextrin-based drug delivery systems

formulation type, providing the benefits of more convenient oral drug administration, especially for children and elderly. The drug released from ODTs and related formulations can act locally in the oral cavity or are absorbed either directly via the oral mucosa or after swallowing with saliva in the gastrointestinal tract (Slavkova and Breitkreutz, 2015). CDs can be efficiently utilized in the development of both sublingual tablets and ODT, as they would provide an efficient mean to ensure the drug dissolution in the limited volume of saliva, what is a prerequisite for efficient oromucosal drug absorption. Furthermore, CDs as soluble excipients would gain to the product disintegration, thereby providing faster dissolution and consequently faster onset of drug action (Conceição et al., 2018). Finally, CD complexation could reduce the unpleasant drug taste (Arima et al., 2012), improving the palatability of the formulation, thereby increasing patient compliance with the new product. As already mentioned, the first CD-based product introduced into clinical practice in 1976 was the sublingual tablet Prostarmon E (Ono Pharmaceuticals, Japan), containing βCD complex of prostaglandin E2, that significantly increased its solid-state chemical stability. Prostarmon E is highly effective and represents a significant clinical advance, especially for the labor induction in oxytocin-insensitive individuals, but also for its tendency to produce less bleeding after delivery (Davis and Brewster, 2004). Another example of CD-based sublingual formulation is Nicorette sublingual tablets (Pharmazia, Sweden), where bitter and astringent taste of nicotine was masked by complexation with βCD (Conceição et al., 2018). Its sublingual administration provides pulsed systemic delivery of nicotine that can mimic more precisely pharmacokinetic plasma profile of nicotine achieved by smoking, therefore, is more efficient in providing the relief of smoking cessation symptoms than transdermal nicotine. However, in some cases, sublingual nicotine administration (as short-acting form) should be combined with transdermal delivery (as a long acting form) to obtain the desired effect (Hukkanen et al., 2005). βCD complexation of testosterone was used in the development of sublingual tablet formulation. The administration of this formulation to hypogonadal men at 2.5 and 5 mg testosterone dose caused a rapid increase of serum testosterone levels with the peak concentration attained after 20 minutes, followed by a gradual decrease to original values in 4 6 hours. Therefore, CD-based sublingual formulation can mimic the pattern of endogenous testosterone release, but multiple daily administrations are needed. However, in the 60-day study period, obtained testosterone plasma concentrations were lower compared to that obtained by testosterone enathate 200 mg dose, applied intramuscularly once every 20 days. Despite the differences in the testosterone serum levels achieved by different androgen replacement therapies, all enrolled patients showed significant improvements in sexual motivation and performance, with no significant difference between the treatment groups. Because of that, sublingual testosterone may be a useful addition to the currently available injectable and transdermal delivery systems for treatment of hypogonadism. It seems especially suitable for

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treatment of boys with delayed puberty and older men with androgen deficiency, where lower serum testosterone levels would be beneficial. However, sublingual CD testosterone formulation would not be efficient for male contraception with androgen alone, where constantly high levels of testosterone are required to suppress the gonadotropins (Salehian et al., 1995). Nowadays, sublingual CD-based testosterone in doses ranging from 0.25 to 0.75 mg is also tested in premenopausal women, with the aim to treat female sexual dysfunction. A pharmacokinetic study demonstrated that, in a patient group with low sex hormone binding globulin (SHBG) levels, sublingual CD testosterone administration lead to a dose-dependent increase of free-testosterone, but such effect was less pronounced in women with high SHBG levels (van Rooij et al., 2012). It was shown that sublingual CD testosterone at 0.5 mg dose produces an increase in sexual motivation and desire in sexually functional woman, about 3 4 hours after its peak plasma levels (van Rooij et al., 2014). This led to the development of two advanced formulations, Lybridos and Lybrido (Emotional Brain NY Inc., SAD), both intended for sublingual administration followed by oral administration, to treat hypoactive sexual desire disorder in premenopausal women. Lybridos is a single fixed dose combination containing the quickly dissolving outer coating made of testosterone (0.5 mg) complexed with CDs, that is applied as a film on the core containing buspirone (10 mg) with time-delayed release properties. A clinical study demonstrated that such formulation produces immediate testosterone pulse absorption followed by adequate absorption of buspirone (80 % relative to the conventional tablet) with a time delay in its release of 3.3 hours (van Rooij et al., 2014). Lybrido, is also single fixed dose combination, containing rapidly dissolving testosterone/CD complex in the outer coating (0.5 mg) on the time-delayed release core loaded with 50 mg of sildenafil. A clinical study confirmed the desired pharmacokinetic profile of both drugs, characterized by the immediate pulse of testosterone, followed by sildenafil release and absorption after 2.75 hours (Bloemers et al., 2016). As already mentioned, there is lag time of about 3 4 hours in the pharmacodynamics effect of sublingual testosterone on genital arousal in women and other cognitive and affective functions. Therefore either the sildenafil (Lybrido) or buspirone (Lybridos) needs to be administered or released approximately 2 3 hours after testosterone administration. While Lybridos is intended for woman who have a dysfunctional activation of sexually inhibitory mechanisms during sexual stimulation, Lybrido is an on-demand therapy for women in whom sexual disorder is caused by low sensitivity to sexual stimuli, increasing sexual motivation, and improving blood flow to the genitals. Both formulations are currently under clinical testing (Bloemers et al., 2016; van Rooij et al., 2014). The advantages of CD in the development of clinically efficient ODT formulation could be clearly demonstrated on examples of ketotifen fumarate fast-dissolving sublingual tablets (Kharshoum and Salem, 2011), lornoxicam ODT (Moutasim et al.,

Cyclodextrin-based drug delivery systems

2017) and vardenafil ODT (Al-Gethmy et al., 2019). The formation of equimolar ketotifen HPβCD inclusion complex promoted fast degradation of the sublingual tablet (21.2 seconds) and fast dissolution of this poorly soluble compounds (more of 80% of the drug dose in 1 minute). Pharmacokinetic study on healthy volunteers showed that this novel sublingual formulation, when compared with Zaditen oral tablet (Novartis, Switzerland), provided shorter Tmax (1.08 vs 1.92 hours), higher cmax (30.80 vs 16.08 ng/mL), and 152% increase in overall bioavailability (Kharshoum and Salem, 2011). In case of lornoxicam, complexation of the drug with βCD was performed to increase drug solubility and mask its bitter taste. When such inclusion complex was formulated as ODT, it exhibited in vivo disintegration time of 13 seconds with more than 95% of the drug dose released in 6 minutes. When applied, ODT formulation provided significantly shorter Tmax (1 vs 2.5 hours) and 105.5% higher overall bioavailability compared with Zeficam oral tablet (Eva Oharma, Egypt). Therefore such ODT has the potential to provide fast relief of pain accompanying rheumatoid arthritis. Furthermore, sensory evaluation of the developed ODT clearly demonstrated its acceptable palatability, where bitter taste of the drug was successfully masked by the use of βCD (Moutasim et al., 2017). Improved solubility and completely masking of the bitter drug taste were accomplished by inclusion complex formation between vardenafil and βCD in 1:2 molar ratio. Such complex was also successfully implemented into ODT formulation, that compared with commercially available Levitra film-coated oral tablet (Bayer, Germany) presented significantly higher cmax (18.18 vs 12.29 ng/mL) that was achieved in a twice shorter time period (tmax of 1.0 and 2.0 hours, respectively), with relative increase drug bioavailability of 122.5% (Al-Gethmy et al., 2019). In some cases, the selection of CD-derivative type can have dominant effect on the ODT product quality attributes and performance. Complexation of famotidine with carboxymethyl-βCD (CMβCD) provided strong stabilizing effect against drug degradation in the acidic environment, whereas SBEβCD had an opposite effect. This favorable stabilizing effect of CMβCD played an important role in enhancing the oral bioavailability of the drug. Further advantages of include the ability to formulate ODT by direct compression technique, preparing tablets with an acceptable physical properties and an adequate masking of the bitter drug taste (Mady et al., 2010). The presented examples clearly show that the use of CD-based technology would provide efficient oromucosal delivery of the drugs that require fast onset of action or are subjected to significant metabolism during the fast passage through liver. The ease of application and acceptable palatability could additionally promote the patient acceptance and compliance of such formulations. 2.3.1.6 Cyclodextrins in dermal formulations CDs are introduced into cosmetic industry in the mid of 1990s and nowadays are widely used in cosmetic products. Recent market analysis revealed that at least

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87 cosmetic products containing CDs are available worldwide (Braga and Pais, 2018). Majority of them, including face and body skin care, cleansing, and decorative cosmetics, are developed using βCD (48.3%), whereas HPβCD, γCD, and RAMEB are present in 39.1%, 11.5%, and 1.1% of the products, respectively. Here CDs are stabilizing the chemically liable compounds against decomposition, increasing solubility of poorly soluble cosmeceuticals, reducing their volatility, and providing prolonged release of active compounds and perfumes. Because of their affinity for interaction with skin lipids, CDs are often introduced into cleaning lotions or in dry hair shampoo (Buschmann and Schollmeyer, 2002; Tarimci, 2011). Furthermore, parent CDs are capable to stabilize the emulsions through the formation of insoluble solid complexes on the phase interface (occurring at high CD concentration) or by the inclusion complexation of lipid compounds resulting in the formation of amphiphilic complexes, mostly occurring at low CD concentration (Mathapa and Paunov, 2013). In this manner, a potentially irritant surfactants could be excluded from the dermal formulations. Finally CDs can interact with some molecules responsible to cause body malodor, thereby acting as deodorants (Lopedota et al., 2015). Natural CDs and their hydrophilic derivatives are not able to permeate skin barrier in significant amounts (0.02% and 0.3% of the applied dose for HPβCD and RAMEB, respectively) and are generally considered to be nonirritant to the skin (Cal and Centkowska, 2008). Moreover, CD complexation can reduce the irritant effect of come cosmeceuticals like retinoic acid. Inclusion complexation with HPβCD improved aqueous solubility and photostability of retinoic acid, providing better skin tolerance and reducing the level of skin inflammation compared to that of the free retinoic acid. Furthermore, CD complexation did not negatively affect antiaging effect of retinoic acid assessed by wrinkle scores, skin elasticity, and wrinkle area measurement. It is hypostatized that the reduction of irritant effect could be attributed to a prolonged release of retinoic acid from the CD complex upon the application to the skin (Miura et al., 2012). This may broaden the efficiency of tretinoin treatment, providing the possibility to apply higher concentration of retinoic acid to the skin to provide more enhanced antiaging effect without significant skin irritation. Higher efficiency and better tolerability of topical retinoic acid/βCD in the treatment of acne vulgaris was demonstrated in an earlier clinical study (Anadolu et al., 2004), providing the possibility to formulate retinoic acid in hydrogel or moisturizing base, that is more convenient for the treatment of oily skin affected with acne vulgaris, thereby further enhancing patient compliance with the treatment. Insulin is known to promote wound healing by affecting proliferation, migration, and extracellular matrix secretion by keratinocytes, endothelial cells, and fibroblasts. As such, it might have a potential to promote the pressure ulcer healing, a complex process that is difficult to achieve. Complexation of insulin with HPβCD improved its solubility, stability, and biological activity. When such complex was incorporated into hydrogel

Cyclodextrin-based drug delivery systems

comprising Carbopol 940 polymer, better healing of the lesion was observed. Although the effect obtained with insulin/HPβCD formulation was not statistically significant in regard to control gel, it was possible to visualize that the clinical conditions of the pressure ulcers were improved with the gel containing insulin/HPβCD complex, showing a gradual reduction in the pressure ulcer size, absence of necrotic tissues, and revitalization of the tissues, suggesting a better response to treatment compared to the control gel and the gel with insulin. Furthermore, such formulation did not show any effect on glucose levels in treated patients and as such it appears to be promising wound healing treatment in hospital care (Valentini et al., 2015). It seems that association of CDs with hydrogels may lead to development of an optimal wound-dressing material. The hydrogel component will maintain the moist environment required for the healing process, whereas CDs can protect and control the release of bioactive molecules. Nowadays, numerous different systems are in preclinical development phase and may be expected to enter into clinical practice (Pinho et al., 2014) Current application of CDs in the development of advanced drug delivery systems for dermal application is mainly directed toward the improvement of the drug loading into nanoparticulate drug delivery systems like liposomes and solid lipid nanoparticles (Gharib et al., 2015), better control of the drug release rate from such nanocarriers (Braga and Pais, 2018), and to increase the (trans)dermal absorption of drugs where RAMEB appeared as very efficient absorption enhancer (Mennini et al., 2016; Mura, 2015). Although some advanced formulations developed using CDs showed promising results, their therapeutically potential still needs to be confirmed in clinical trials.

2.3.2 Cyclodextrin as novel therapeutically active pharmaceutical ingredients The background of CD use as novel therapeutically active agents lies in their ability to interact with different biological molecules, like cholesterol and triglycerides, or with the administered drug within the body, thereby achieving a wide variety of potentially therapeutic effects (di Cagno et al., 2016). 2.3.2.1 Sugammadex The first marketed CD-based drug in clinical use was sugammadex (Bridion, Merck), approved first in Europe in 2008 than in Japan 2010, followed by the United States and Canada in 2015 and 2016, respectively. This highly water-soluble modified γ-CD derivative is the first representative of a new class of drugs called steroidal muscle relaxant encapsulators, introducing complete new possibility of reversal of a neuromuscular block into anesthesia practice (Schaller and Lewald, 2016). Administered intravenously at recommended doses, ranging from 2 to 16 mg/kg, sugammadex provides rapid and predictable reversal of rocuronium, vecuronium, and pipecuronium-induced moderate and deep neuromuscular block (Asztalos et al., 2017; Herring et al., 2017; Schaller and

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Lewald, 2016; Tassonyi et al., 2018). Furthermore, sugammadex is proven to be more efficient and safer than neostigmine, a cholinesterase inhibitor agent that is commonly used to antagonize neuromuscular blocker (Hristovska et al., 2017). Backbone of sugammadex is a γ-CD molecule with lipophilic central cavity extended through addition of eight carboxyl thioether groups to position 6 of glucopyranose unit (Fig. 2.3), tailored to enable a complete encapsulation of the steroidal muscle relaxant. Negatively charged terminal carboxyl groups prevent side-chains from entering the cavity and electrostatically bind the positive-charged ammonium group in the rocuronium molecule, contributing to its strong binding within the CD with Ks of 25 3 106 M21 (Akha et al., 2010). Other steroidal muscle relaxants like vecuronium and pancuronium are also bound, but with a much lower affinity. After intravenous injection, sugammadex binds free intravascular rocuronium, leading to formation of a concentration gradient that shifts the drug from the peripheral compartment (i.e., neuromuscular junction) toward intravascular compartment, where is also encapsulated by the ligand and rapidly eliminated from the body through the glomerular filtration. This rapidly restores neuromuscular transmission and muscle function if sugammadex is applied within 3-minute time frame after steroidal muscle relaxant

Figure 2.3 The chemical structure of sugammedex, a steroidal muscle relaxant encapsulator drug.

Cyclodextrin-based drug delivery systems

administration (Schaller and Lewald, 2016). Compared to neostigmnine, sugammadex provides a faster reversal of the neuromuscular block within 1.9 3.8 minutes, depending on the muscle relaxant used, dose of sugammadex administered and the depth of the neuromuscular block at the time of reversal, leading to accelerated postoperative discharge of patients after general anesthesia. Under the same condition, neostigmine produces reversal of the neuromuscular block within 10.6 67.6 minutes (Carron et al., 2017; Herring et al., 2017). Furthermore, the use of sugammadex is accompanied by lower incidence and severity of postoperative nausea and vomiting than neostigmine (Tas Tuna et al., 2017). One of the most commonly reported adverse reactions on sugammadex application is cough that occurs in approximately 70% of patients reversed with single 2 mg/kg bolus dose of sugammadex. However, staggering the dose by application of 1 mg/kg sugammadex prior to extubation followed by another 1 mg/kg immediately after extubation reduces the incidence of severe cough to 12.5%, whereas time taken for emergence and reversal was not statistically different (Loh et al., 2017). Sugammadex alone at low, medium, or high clinical doses has no effect on anesthetic depth (Fassoulaki et al., 2017) and produces limited transient (,1 hour) increase of activated partial thromboplastin time and prothrombin time that is not associated with increased risk of bleeding versus usual care (Rahe-Meyer et al., 2014). The presence of antibiotics like kanamycin gentamicin, vancomycin, clindamycin and bacitracin prior to the administration of sugammadex did not affect the recovery time from rocuronium-induced neuromuscular block when sugammadex 4.0 mg/ kg was administered at least 15 minutes after the last dose of rocuronium, suggesting that prophylactic antibiotic use is unlikely to have a major impact on the recovery time from rocuronium-induced neuromuscular block with sugammadex reversal (Hudson et al., 2014). In addition, sugammadex administered at a dose of 4 mg/kg is not associated with adverse effects on serum levels of steroid hormones like progesterone and cortisol, whereas it may lead to a temporary increase in serum concentrations of aldosterone and testosterone (Gul et al., 2016). Regarding the sugammadex administration to specific patient groups, dose adjustments of sugammadex should be considered when rapid recovery from deep neuromuscular block is needed in elderly adults. Some studies reported that in patients aged .65 years, time to recovery from rocuronium-induced neuromuscular block after sugammadex administration is prolonged on average by 1 2 minutes compared with younger patients (Carron et al., 2018). Because of that, some authors are considering that elderly patients should require 1 mg/kg higher sugammadex doses than young ones (Shin et al., 2016). Although sugammadex has not received approval for administration in pediatric population, its use has already been reported in several clinical scenarios such as children with neuromuscular diseases including myasthenia gravis, Duchenne muscular dystrophy and myotonic dystrophy. The pediatric data are limited to reversal of rocuronium-induced neuromuscular blockade except for a single case

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report involving vecuronium, and it seems that sugammadex may be clinically advantageous in certain conditions where acetylcholinesterase inhibitors are relatively contraindicated, including myotonic dystrophy and in cardiac transplantation patients (Tobias, 2017). However, future clinical trials are needed to further define the use of sugammadex in the pediatric population. A growing proportion of patients undergoing surgical procedures are obese, presenting anesthesiologists with numerous challenges for patient management. A pooled analysis of sugammadex use in obese patients showed no clinically relevant correlation between body mass index (BMI) and recovery time, indicating that recommended sugammadex dosing regimen based on actual body weight provide rapid recovery from neuromuscular blockade in both obese and nonobese patients (Monk et al., 2017). Sugammadex is one of the most expensive compounds used in anesthesiology, with the price of about 80h plus taxes for 200 mg dose (Schaller and Lewald, 2016). Therefore several studies have reported administration of a lower-than-recommended dose of sugammadex, as a potential cost-saving strategy. It seems that such practice is not beneficial as underdosing of sugammadex leads to an increased risk of recurrence of neuromuscular block after initial successful (but transient) reversal and attendant postoperative complications (De Boer et al., 2018).

2.3.3 Treatment of Niemann Pick disease, type C1 disease with HPβCD Niemann Pick disease, type C (NP-C) is a rare autosomal recessive neurodegenerative disorder with estimated incidence of 1 in 120,000 to 150,000 live births. It is caused by mutations in NPC1 (95% of cases) or NPC2 gene (about 5% of cases) that yield deficient function of the corresponding proteins that normally bind and transport cholesterol from late endosomes and lysosomes to other cellular compartments. This leads to endolysosomal storage of unesterified cholesterol and sphingolipids in brain, liver, spleen, and other peripheral tissues. The clinical presentation of NP-C is nonspecific and highly heterogeneous, depending on the age of onset. Affected individuals typically exhibit ataxia, swallowing problems, seizures, and progressive impairment of motor and intellectual function in early childhood and usually die in adolescence (Megías-Vericat et al., 2017; Ory et al., 2017; Ottinger et al., 2014). Currently, there is no curative treatment for NP-C and miglustat (Zavesca, Janssen-Cilag), a small iminosugar that is approved for use in some countries, does not mobilize intracellularly accumulated cholesterol, thereby having only modest efficacy in delaying the disease progression (Ottinger et al., 2014). Because of pronounced cholesterol binding properties (Szente and Fenyvesi, 2017), CDs have been investigated as a potential therapy of NP-C disease. Among them, HPβCD has been shown to reduce the cholesterol accumulation in NPC1 mutant cells. It seems that in addition to HPβCD, HPγCD could

Cyclodextrin-based drug delivery systems

also be therapeutically active. Both CDs upregulate the expression of the lysosomal membrane protein LAMP-1 and facilitate cholesterol trafficking from late endosomes and lysosomes to other cellular compartments, rescuing the cholesterol accumulation defect in NP-C patient-derived fibroblast cells (NPC1 mutant) (Singhal et al., 2018). Preclinical studies on mice and feline model showed that to confer attenuated neurodegeneration and prolong survival, direct HPβCD administration to central nervous system is necessary, as this CD derivate does not cross blood brain barrier (Singhal et al., 2018). Some early clinical studies on NP-C patients demonstrated that intrathecal administration of escalated HPβCD dose (50 900 mg) during 1 month slow down the disease progression and produced some clinical benefits, like improved neuronal cholesterol homeostasis and decreased neuronal pathology (Ory et al., 2017). Longterm intrathecal administration (2.5 3 years) of escalating HPβCD doses (200 1200 mg) has stabilized the disease over extended period of time and showed improvements in function of neurological domains typically impacted by the disease (Berry-Kravis et al., 2018). From a safety standpoint, doses up to 1200 mg were generally well tolerated, however, all patients developed functional deficits in hearing, confirming ototoxic effect of HPβCD that was already observed on animal models (Crumling et al., 2017). The ototoxicity after HPβCD administration is generally considered as expected and acceptable adverse event, when placed in the context of a disease with fatal outcome (Ory et al., 2017). It should be considered that hearing loss is a manifestation of NP-C disease, but it is getting worse after HPβCD administration. However, the extent of changes in hearing caused by HPβCD administration is highly individual, and these deficits did not appreciably affect daily communication when suitably managed with hearing aids (Ory et al., 2017). In an attempt to overcome the limitations connected with of HPβCD administration, βCD-based linear polymer prodrug (ORX-301) constructed of βCD monomer units linked together by a short degradable ketal linkage has been developed (Kulkarni et al., 2018). Subcutaneously injected ORX-301 extended the mean life span of NP-C mice at a dosage fivefold lower (800 mg/kg, body weight) than those of HPβCD (4000 mg/kg). ORX-301 was safe when administered subcutaneously and no clinical signs of toxicity were observed up to a dose of 2000 mg/kg. Compared to HPβCD, ORX-301 showed prolonged systemic circulation up to 48 hours, higher accumulation in liver and spleen. It also penetrates the blood brain barrier and counteracts neurological impairment, thereby represent a substantial improvement of presently available HPβCD-based therapy. However, its efficiency and safety in human patients should be still needs to be confirmed by clinical trials. 2.3.3.1 Cyclodextrins in control of obesity and hyperlipidemia The incidence of obesity has reached historically unprecedented levels with approximately 75% of adults in the United States that are either overweight or obese, while

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in Western Europe incidence of overweight or obese persons is approximately 61% in men and 48% in women. One of possible approaches to combat this situation could be the development of food ingredients that would limit fat absorption by fat binding or preventing their absorption (Gallaher and Plank, 2015). Some early in vivo animal studies showed that αCD binds dietary fat, preferentially saturated ones and reduces its intestinal absorption which may provide some health benefits. The human studies, although very successful in beneficial changes of blood lipid profile, gave disappointing results concerning body weight reduction related to the expectations based on animal experiments (Fenyvesi et al., 2016). In general, αCD supplementation was safe and reasonably well tolerated in a treated population and did not interfere with the intake of lipophilic vitamins (A, D, E) (Amar et al., 2016; Grunberger et al., 2007). However, the effect of αCD supplementation in oral dose of 6 g/day on plasma lipid levels in a relatively healthy population resulted only in 10% reduction in small LDLparticle number when subjects were on αCD versus placebo, with no other changes in the lipid and lipoprotein profiles, the body weight, serum insulin, and HbA1C (Amar et al., 2016). In contrast to that, a previous clinical study of αCD effect on serum lipids level in hypertriglyceridemic obese patients with BMI .30 kg/m2, which also used the oral dose of 6 g/day, demonstrated a reduction in LDL of 11.9 6 4.2 mg/dL after 3 months treatment, whereas in the placebo group, an increase in LDL of 8.5 6 6.2 mg/dL (P , .01) was observed. Furthermore, placebo group continued to gain weight during the study, whereas those in the αCD group maintained their weight (Grunberger et al., 2007). Another study conducted with overweight individuals having a mean BMI of 26.9 kg/m2 and a mean age of 43.3 years, demonstrated that 1 month of αCD supplementation led to significant weight loss in the compliant participants (20.4 6 0.2 kg, P , .05), in the absence of any change of energy intake. In the same time, a decrease in total cholesterol by 25.3%, LDL cholesterol (26.7%), and adiponectin B (25.6%), in the absence of any other dietary modifications was also observed (Comerford et al., 2011). Interestingly subjects who had the highest total cholesterol and LDL cholesterol tended to show the greatest reductions in those parameters when supplemented with αCD. From the presented data, it seems that αCD may be more effective in serum lipid reduction in a more dyslipidemic and obese population, but this will have to be more definitively established in larger clinical trials. Another interesting result of αCD supplementation in healthy population is the decrease of fasting plasma glucose level (1.6%, P , .05) and insulin resistance index (11%, P , .04) in treated group versus placebo (Amar et al., 2016). Furthermore, some early studies showed that CDs, especially αCD and γCD, can reduce postprandial glucose and insulin levels, without causing carbohydrate malabsorption (Asp et al., 2006; Buckley et al., 2006). This is in line with several reports showing the ability of natural CDs to reduce the digestion of starch in the mammalian organism (Fenyvesi et al., 2016). The underlying mechanism is in the fact that α- and

Cyclodextrin-based drug delivery systems

βCDs are not only resistant to usual starch degrading enzymes, but also inhibit the hydrolysis of amylose by binding to the active sites of enzymes. Furthermore, CDs interact with starch as evidenced by a reduced thermal stability of amylose in the amylose βCD complex and a decrease of extractable βCD from 60% to 51% after their complexation. All this contributes to the ability of CDs to reduce the glycemic index of food. Furthermore, clinical trial showed that consumption of αCD, beside reduction of the serum lipids also lead to significant increase in blood levels of adiponectin, indicating an increase of insulin sensitivity (Comerford et al., 2011). The EFSA Panel on Dietetic Products, Nutrition and Allergies recommended that to obtain the claimed effect, at least 5 g of αCD per 50 g of starch should be consumed (European Food Safety Authority, 2012). Similar effect of γCD on postprandial glycemia was also observed (Asp et al., 2006), however, as γCD is rapidly metabolized and absorbed in the small intestine, the exacts mechanism by which this digestible carbohydrate changes the glycemic index of food is still unclear (Fenyvesi et al., 2016). The presented examples of the CD use in clinical practice clearly demonstrate their multifunctionality as enabling excipients and drug carriers in the development of conventional and advanced dosage forms. The choice of CD derivative to be applied in the formulation development is directed by the intended administration route as well as by the most favorable affinity of the drug in question for the inclusion complexation. The benefits obtained by inclusion complexation include increase of drug solubility and chemical stability, taste masking, controlled drug release and increased bioavailability, providing a valuable mean to develop highly loaded drug formulations devoid of potentially irritant excipients like surface-active compounds and cosolvents. In addition, CD complexation decreased the irritant properties of drugs, extending their efficiency and safety. Due to all that, CD-based technology enables the development of novel formulations containing different poorly soluble drugs and extends the application routes and clinical efficiency of numerous drugs, being a useful strategy in life-cycle management of medicines. Nowadays, CDs are in focus of scientific research as potent novel pharmaceutically active compounds, especially in treatment of diseases for which an adequate therapy is not available. Numerous new chemically modified CD derivatives have been synthetized and tested as potentially novel treatment for artheriosclerosis, dyslipidemia, and degenerative brain diseases (di Cagno et al., 2016). Of specific interest are CD derivatives that can act as tumor-targeting carrier of anticancer drugs or show anticancer activity on their own (Qiu et al., 2017), as well as those that can enhance antibiotic delivery to bacterial biofilm (Thomsen et al., 2017) or even act as broad-spectrum antibiotics with pore-forming proteins as the targets (Karginov, 2013). Also, there is a growing interest for CDs as building blocks in regenerative medicine (Alvarez-Lorenzo et al., 2017). Although preclinical development of such systems clearly demonstrated their potential, there is a persistent problem in translation of many laboratory efficient and clinically promising CD-based

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formulation into clinically useful product that would be cost effective. Constant advances in the development of preclinical drug evaluation models, able to mimic more closely the human pathophysiological conditions would provide deeper insight into its stability, efficiency, performance, and safety upon administration, providing valuable inputs for further product optimization and development, contributing to more successful clinical translation of such systems.

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Ohira, A., Hara, K., Jóhannesson, G., Tanito, M., Ásgrímsdóttir, G.M., Lund, S.H., et al., 2015. Topical dexamethasone γ-cyclodextrin nanoparticle eye drops increase visual acuity and decrease macular thickness in diabetic macular oedema. Acta Ophthalmol. 93, 610 615. Available from: https://doi. org/10.1111/aos.12803. Ory, D.S., Ottinger, E.A., Farhat, N.Y., King, K.A., Jiang, X., Weissfeld, L., et al., 2017. Intrathecal 2hydroxypropyl-β-cyclodextrin decreases neurological disease progression in Niemann-Pick disease, type C1: a non-randomised, open-label, phase 1 2 trial. Lancet 390, 1758 1768. Available from: https://doi.org/10.1016/S0140-6736(17)31465-4. Ottinger, E., Kao, M., Carrillo-Carrasco, N., Yanjanin, N., Shankar, R., Janssen, M., et al., 2014. Collaborative development of 2-hydroxypropyl-β-cyclodextrin for the treatment of Niemann-Pick type C1 disease. Curr. Top. Med. Chem. 14, 330 339. Available from: https://doi.org/10.2174/ 1568026613666131127160118. Pelissier, C., de Kervasdoue, A., Chuong, V.T., Maugis, E.-L., de Mouillac, F., Breil, M.-H., et al., 2001. Clinical evaluation, dose-finding and acceptability of AERODIOLs, the pulsed estrogen therapy for treatment of climacteric symptoms. Maturitas 37, 181 189. Available from: https://doi.org/ 10.1016/S0378-5122(00)00175-4. Pinho, E., Grootveld, M., Soares, G., Henriques, M., 2014. Cyclodextrin-based hydrogels toward improved wound dressings. Crit. Rev. Biotechnol. 34, 328 337. Available from: https://doi.org/ 10.3109/07388551.2013.794413. Pontiroli, A.E., Ceriani, V., 2018. Intranasal glucagon for hypoglycaemia in diabetic patients. An old dream is becoming reality? Diabetes Obes. Metab. 20, 1812 1816. Available from: https://doi.org/ 10.1111/dom.13317. Popielec, A., Loftsson, T., 2017. Effects of cyclodextrins on the chemical stability of drugs. Int. J. Pharm. 531, 532 542. Available from: https://doi.org/10.1016/j.ijpharm.2017.06.009. Preskorn, S.H., 2005. Pharmacokinetics and therapeutics of acute intramuscular ziprasidone. Clin. Pharmacokinet. 44, 1117 1133. Available from: https://doi.org/10.2165/00003088-200544110-00002. Qiu, N., Li, X., Liu, J., 2017. Application of cyclodextrins in cancer treatment. J. Incl. Phenom. Macrocycl. Chem. 89, 229 246. Available from: https://doi.org/10.1007/s10847-017-0752-2. Rahe-Meyer, N., Fennema, H., Schulman, S., Klimscha, W., Przemeck, M., Blobner, M., et al., 2014. Effect of reversal of neuromuscular blockade with sugammadex versus usual care on bleeding risk in a randomized study of surgical patients. Anesthesiology 121, 969 977. Available from: https://doi.org/ 10.1097/ALN.0000000000000424. Rickels, M.R., Ruedy, K.J., Foster, N.C., Piché, C.A., Dulude, H., Sherr, J.L., et al., 2015. Intranasal glucagon for treatment of insulin-induced hypoglycemia in adults with type 1 diabetes: a randomized crossover noninferiority study. Diabetes Care 39. Available from: https://doi.org/10.2337/dc151498dc151498. Rodriguez-Aller, M., Guinchard, S., Guillarme, D., Pupier, M., Jeannerat, D., Rivara-Minten, E., et al., 2015. New prostaglandin analog formulation for glaucoma treatment containing cyclodextrins for improved stability, solubility and ocular tolerance. Eur. J. Pharm. Biopharm. 95, 203 214. Available from: https://doi.org/10.1016/J.EJPB.2015.04.032. Saari, K.M., Nelimarkka, L., Ahola, V., Loftsson, T., Stefánsson, E., 2006. Comparison of topical 0.7% dexamethasone-cyclodextrin with 0.1% dexamethasone sodium phosphate for postcataract inflammation. Graefe’s Arch. Clin. Exp. Ophthalmol. 244, 620 626. Available from: https://doi.org/ 10.1007/s00417-005-0124-2. Salapatek, A.M., Lee, J., Patel, D., D’Angelo, P., Liu, J., Zimmerer, R.O., et al., 2011. Solubilized nasal steroid (CDX-947) when combined in the same solution nasal spray with an antihistamine (CDX313) provides improved, fast-acting symptom relief in patients with allergic rhinitis. Allergy Asthma Proc. 32, 221 229. Available from: https://doi.org/10.2500/aap.2011.32.3444. Salehian, B., Wang, C., Alexander, G., Davidson, T., McDonald, V., Berman, N., et al., 1995. Pharmacokinetics, bioefficacy, and safety of sublingual testosterone cyclodextrin in hypogonadal men: comparison to testosterone enanthate--a clinical research center study. J. Clin. Endocrinol. Metab. 80, 3567 3575. Available from: https://doi.org/10.1210/jcem.80.12.8530600.

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Tarimci, N., 2011. Cyclodextrins in the cosmetic field. Cyclodextrins in Pharmaceutics, Cosmetics, and Biomedicine. John Wiley & Sons, Inc., Hoboken, NJ, pp. 131 144. Available from: https://doi.org/ 10.1002/9780470926819.ch7. Tas Tuna, A., Palabiyik, O., Orhan, M., Sonbahar, T., Sayhan, H., Tomak, Y., 2017. Does sugammadex administration affect postoperative nausea and vomiting after laparoscopic cholecystectomy. Surg. Laparosc. Endosc. Percutan. Tech. 27, 237 240. Available from: https://doi.org/10.1097/ SLE.0000000000000439. Tassonyi, E., Asztalos, L., Szabó-Maák, Z., Nemes, R., Pongrácz, A., Lengyel, S., et al., 2018. Reversal of deep pipecuronium-induced neuromuscular block with moderate versus standard dose of sugammadex: a randomized, double-blind, noninferiority trial. Anesth. Analg. 127, 1344 1350. Available from: https://doi.org/10.1213/ANE.0000000000003719. Thomsen, H., Benkovics, G., Fenyvesi, É., Farewell, A., Malanga, M., Ericson, M.B., 2017. Delivery of cyclodextrin polymers to bacterial biofilms — an exploratory study using rhodamine labelled cyclodextrins and multiphoton microscopy. Int. J. Pharm. 531, 650 657. Available from: https://doi.org/ 10.1016/j.ijpharm.2017.06.011. Tobias, J.D., 2017. Current evidence for the use of sugammadex in children. Paediatr. Anaesth. 27, 118 125. Available from: https://doi.org/10.1111/pan.13050. Tolbert, D., Cloyd, J., Biton, V., Bekersky, I., Walzer, M., Wesche, D., et al., 2015. Bioequivalence of oral and intravenous carbamazepine formulations in adult patients with epilepsy. Epilepsia 56, 915 923. Available from: https://doi.org/10.1111/epi.13012. Valentini, S.R., Nogueira, A.C., Fenelon, V.C., Sato, F., Medina, A.N., Santana, R.G., et al., 2015. Insulin complexation with hydroxypropyl-beta-cyclodextrin: spectroscopic evaluation of molecular inclusion and use of the complex in gel for healing of pressure ulcers. Int. J. Pharm. 490, 229 239. Available from: https://doi.org/10.1016/j.ijpharm.2015.05.037. van Rooij, K., Bloemers, J., de Leede, L., Goldstein, I., Lentjes, E., Koppeschaar, H., et al., 2012. Pharmacokinetics of three doses of sublingual testosterone in healthy premenopausal women. Psychoneuroendocrinology 37, 773 781. Available from: https://doi.org/10.1016/j.psyneuen. 2011.09.008. van Rooij, K., de Leede, L., Frijlink, H.W., Bloemers, J., Poels, S., Koppeschaar, H., et al., 2014. Pharmacokinetics of a prototype formulation of sublingual testosterone and a buspirone tablet, versus an advanced combination tablet of testosterone and buspirone in healthy premenopausal women. Drugs R. D. 14, 125 132. Available from: https://doi.org/10.1007/s40268-014-0047-7. VFEND (voriconazole) for injection, for intravenous use, n.d. ,www.fda.gov/medwatch. (accessed 20.02.19.). Vizzardi, M., Sagarriga Visconti, C., Pedrotti, L., Marzano, N., Berruto, M., Scotti, A., 1998. Nimesulide beta cyclodextrin (nimesulide-betadex) versus nimesulide in the treatment of pain after arthroscopic surgery. Curr. Ther. Res. 59, 162 171. Available from: https://doi.org/10.1016/S0011-393X(98) 85012-1. Voriconazole 200mg Powder for Solution for Infusion - Summary of Product Characteristics (SmPC) (eMC), n.d. ,https://www.medicines.org.uk/emc/product/7316/smpc. (accessed 20.02.19.). Voss, M.H., Hussain, A., Vogelzang, N., Lee, J.L., Keam, B., Rha, S.Y., et al., 2017. A randomized phase II trial of CRLX101 in combination with bevacizumab versus standard of care in patients with advanced renal cell carcinoma. Ann. Oncol. 28, 2754 2760. Available from: https://doi.org/10.1093/annonc/mdx493. Wang, D., Miller, R., Zheng, J., Hu, C., 2000. Comparative population pharmacokinetic-pharmacodynamic analysis for piroxicam-beta-cyclodextrin and piroxicam. J. Clin. Pharmacol. 40, 1257 1266. Weiss, G.J., Chao, J., Neidhart, J.D., Ramanathan, R.K., Bassett, D., Neidhart, J.A., et al., 2013. Firstin-human phase 1/2a trial of CRLX101, a cyclodextrin-containing polymer-camptothecin nanopharmaceutical in patients with advanced solid tumor malignancies. Invest. New Drugs 31, 986 1000. Available from: https://doi.org/10.1007/s10637-012-9921-8. Wen, H., Jung, H., Li, X., 2015. Drug delivery approaches in addressing clinical pharmacology-related issues: opportunities and challenges. AAPS J. 17, 1327 1340. Available from: https://doi.org/ 10.1208/s12248-015-9814-9. Zhang, J., Ma, P.X., 2013. Cyclodextrin-based supramolecular systems for drug delivery: recent progress and future perspective. Adv. Drug Deliv. Rev. 65, 1215 1233. Available from: https://doi.org/ 10.1016/j.addr.2013.05.001.

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CHAPTER THREE

Lipid vesicles for (trans)dermal administration Silvia Franzè, Umberto M. Musazzi and Francesco Cilurzo Department of Pharmaceutical Sciences, University of Milan, Milan, Italy

3.1 (Trans)dermal drug-delivery systems (Trans)dermal drug-delivery systems are defined as prolonged-release dosage forms that, when applied on intact skin, are able to deliver the drug through the skin to reach systemic circulation or to obtain local (cutaneous) or regional effects. The cutaneous application of a drug by (trans)dermal delivery systems is an attractive alternative to both parental and oral administration because it is associated with a higher compliance. Indeed when a systemic effect is desired or, in other words, when a certain therapeutic concentration in the blood to treat diseases remote from the application site is needed (Ghosh and Pfister, 1997), they allow to solve biopharmaceutical issues such as peaks and valley profile, and first-pass effect reducing the drug side effects. In the case of locoregional treatments, they avoid significant blood concentration improving again the patient compliance and adherence to therapy. Nevertheless, although the skin is the widest and most exposed tissue of human body, it is not easily accessible. In fact skin evolved to exert a barrier function against all xenobiotics, and this makes the delivery of drugs in and through the skin one of the most challenging examples of drug targeting. The most responsible of the diffusional resistance of the skin is the stratum corneum (SC), the superficial layer of the tissue having unique chemical and morphological features. This is a very thin layer composed of dead cells, the corneocytes, filled with keratin and imbibed in an extracellular lipid matrix in which the constituents have a highly ordered, interdigitated configuration with the formation of gel-phase (not fluid) membrane domains (Naik et al., 2000). The absorption of molecules through the skin may be considered as a combination of processes of repartition and diffusion (governed by the Fick’s law). When a dosage form is applied on the skin, the first step is the repartition of the drug from the vehicle to the SC. This step is as faster as higher is the affinity of the drug for the SC environment and then requires a certain hydrophobicity. Then the drug has to diffuse through Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00003-5

© 2020 Elsevier Inc. All rights reserved.

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the SC that is the rate-limiting barrier and then to repartition in the different layers of the skin up to the dermis while absorption (distribution in the bloodstream) takes place. Based on the above considerations, the ideal candidate for transdermal administration should have low molecular weight (,500 Da) and an appropriate oil/water partition coefficient (log P 1 3), and when a systemic effect is desired, it has to be a very potent drug (low-daily dose). It is clear that, according to these rules, a wide pool of molecules that can be used both for locoregional and systemic pharmacological treatments should be excluded. Thus for these molecules several chemical or physical enhancement strategies have been exploited to perturb reversibly or irreversibly the skin barrier. The most investigated approach relies without doubt on the use of chemical penetration enhancer, namely, small compounds able to intercalate in the SC temporarily altering its barrier properties or to improve the thermodynamic activity of the drug in the vehicle (Alkilani et al., 2015). More recently the use of skin-penetrating peptides has also raised a great interest (Nasrollahi et al., 2012). Otherwise physical techniques such as iontophoresis, electroporation, sonophoresis, and microneedles can be exploited alone or in combination with chemical enhancer. However, some of these approaches (i.e., microneedles) determine an irreversible disruption of skin structure (Swain et al., 2011). In the past 30 years, the field of (trans)dermal drug delivery has been characterized by an intensive attempt to exploit nanotechnologies to overcome the skin barrier without compromising the integrity of the tissue. Among all the nanocarriers tested for this purpose, liposomes and derivatives have been the most investigated and promising vectors and will be the main object of this chapter. The main characteristics of all classes of lipid vesicles used for (trans)dermal application will be analyzed along with the main results of the in vitro and preclinical studies carried out. Finally the clinical outcomes will be discussed.

3.1.1 Human skin barrier to xenobiotics The skin is the largest organ of human body, covering a total surface area of about 1.8 m2 and accounting for more than 10% of the body weight. Skin organization is highly complex, and the anatomical and physiological properties vary drastically from one layer to another of the tissue. In general skin consists of four strata that, moving from the surface to the interior of the organ, are the SC, the viable epidermis, the dermis, and the subcutaneous tissue (or hypodermis). The last one is connected to the overhead dermis through collagen and elastin fibers, and it is essentially composed of flat cells organized in lobules, which may represent a local depot for hydrophobic compounds penetrated up to this level. This layer acts as an energy storage system and provides isolation and protection against injuries and insults

Lipid vesicles for (trans)dermal administration

(Walters and Roberts, 2002). The dermis is the connective layer of the skin and it is characterized by a little cell content (especially fibroblasts and adipocytes) and a huge amount of collagen bundles, being therefore the main one responsible for the mechanical elasticity of the skin. This elastic network supports nerves, glands, lymphatic, and blood capillaries. In fact at this level is localized an important vascular network that accounts for the overall nutrition of the skin, including that of the epidermis, which, being not vascularized, receives the nutrients through the papillae reached by the capillary loops which drain the blood from the dermis (Schaefer and Redelmeier, 1996). The basal lamina separates the dermis from the epidermis, the outer layer of the skin that acts as a first-line barrier against environmental insults and dehydration. The epidermis comprises several strata of keratinizing epithelial cells at different levels of differentiation along with some Langerhans cells, Merkel cells, and melanocytes (Schaefer and Redelmeier, 1996). In the basal layer of the epidermis, the keratinocytes are attached to the matrix and are still in a proliferative state. In the above stratum spinosum, keratinocytes are strictly linked to each other through desmosomes and are characterized by a large cytoplasm containing numerous filaments and organelles. In the outer part of the stratum spinosum, granules enclosing the intracellular membranes appear, thus signaling the transition to the granular layer. The last stage of keratinocytes differentiation evolves in the SC cells. At the limit between the granulosum and corneum strata lysis enzymes are released, the cells loose the nuclei and the organelles and convert into elongated, flat, and highly keratinized cells, the corneocytes, which represent the cell population of the SC (Walters and Roberts, 1996). This is a 10 20μm-thick layer that, observed in longitudinal section, appears as a wall composed of several columns of partially overlapped corneocytes. Each brick of this wall is formed by clusters of 12 15 layers of corneocytes, sustained by a protein scaffold of keratin filaments. Individual corneocytes are tightened through desmosomes and the intercellular spaces are filled with a fluid continuous lipid phase. The extracellular spaces instead are filled with a highly dense and crystalline lipid matrix, organized in periodic lamellae. At the lateral junctions of corneocytes, some hydrophilic pores exist, which are the result of the imperfect overlapping of cell membranes (Cevc, 2004). In addition to these few hydrophilic pathways, the SC is essentially a dry layer with a water content that does not overcome the 15% w/w (against the 75% of the viable epidermis). Finally along with the channel-like hydrophilic furrows at the level of the imperfect lateral junction between corneocytes clusters, some gaps also exist among the polar head of lipids across the intercellular lipid matrix. Due to its peculiar structure, the so-called “bricks” (corneocytes clusters) and “mortar” (intercellular lipids) structure, the SC defines the main barrier unit of the human skin. The potential pathways of penetration of compounds through the human skin are the intercellular, the transcellular, and the appendagael routes. The last one refers to

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the penetration mostly through the hair follicles, the most abundant appendages of the skin but, as they represent a few percentage of the SC composition (0.1%), this way is not a preferential one (Trommer and Neubert, 2006). The intercellular route presumes the passages through the intercellular, lipid-rich spaces between corneocytes, whereas the transcellular route, even higher resistant to penetration of xenobiotics, requires the crossing of cytoplasm of SC cells and both the lipid and the keratin matrix which fills corneocytes (Trommer and Neubert, 2006).

3.2 Lipid vesicles for breaching the skin barrier The existence of nanosized opening (ranging from 0.4 to 36 nm) within the SC structure raised the possibility of exploiting nanotechnology to breach the skin barrier. Ideally the design of a nanocarrier able to pass through the SC without losing its payload would be the turning point to succeed finally to deliver through the skin any kind of active molecule, regardless of the physicochemical nature, after decades and decades of fighting. In fact the delivery of drugs through the skin would depend on a physical process, namely, the penetration of the carrier, rather than Fickian diffusion. Polymeric nanoparticles fail in that purpose as they have little interaction ability with the skin environment, they do not penetrate the skin, either intact or compromised, but simply act as a drug reservoir on the skin surface (Campbell et al., 2012). In contrast, lipid vesicles are being considered as the carriers of election to gain access to the epidermis and dermis layers via the tortuous pathway of the SC along which the spaces are filled with a lipid glue (Dubey et al., 2007). Liposomes are spherical vesicles composed of lipid bilayers enclosing aqueous compartments. Due to this peculiar composition, liposomes have unique properties as drug-delivery systems because they are able to deliver different kinds of molecules (hydrophilic, hydrophobic, and amphipathic), being the most versatile carrier among all the nanotechnologies. Moreover liposome technology offers a wide range of simple approaches to modify the intrinsic properties of the carriers, in terms of dimension, surface properties, and membrane fluidity (Bozzuto and Molinari, 2015). Therefore besides the widespread use of liposomes as biocompatible carriers for parental administration with several formulations already in clinical use, liposome technology opens a wide range of therapeutic opportunities also in (trans)dermal drug delivery. In fact liposomes applied on the skin can enhance the penetration of compounds through several mechanisms. First liposomes can improve the apparent solubility of the applied drug increasing the flux through the skin. Moreover they can act as occlusive

Lipid vesicles for (trans)dermal administration

vehicles, thus increasing the hydration level of the SC which affects the spatial organization of lipid network of the SC (Casiraghi et al., 2002). This alteration determines a decrease of the barrier properties and therefore increases the flux of the loaded drug through the SC (Gennari et al., 2016). Finally analogously to what takes place at cell membrane level, liposomes may fuse and mix with the SC lipids favoring the partitioning of some molecules in the skin (El Maghraby et al., 2008). Liposomes may act as drug carriers to be used as prolonged-release systems for reaching therapeutic concentrations of drugs into the systemic circulation or may be used for targeted delivery to skin appendages (El Maghraby et al., 2008). Liposomes may also be exploited to form a local depot of drugs to maximize the effect in the target tissue (i.e., dermis for antibiotics and corticosteroids) minimizing the systemic absorption. The prominent effect exerted by liposomes once applied on the skin strongly depends on liposomes composition. Several classes of liposomes have been tested or specifically developed for (trans)dermal drug-delivery purposes, namely conventional liposomes, deformable (or elastic) liposomes, ethosomes, transethosomes, invasomes, and skin penetration enhancer-containing vesicles (PEVs).

3.2.1 Conventional liposomes The application of liposomes for delivering active molecules to the skin tissue was pioneered by Mezei and Gulasekharam (1980). They reported the first evidence that lipid vesicles may improve the localization of drugs in the epidermis and dermis layers. Indeed they found a higher triamcinolone acetonide concentration in epidermis and dermis after the topical application of liposomes compared to the corresponding free drug solution and gel formulation (Mezei and Gulasekharam 1980, 1982). Similar findings were obtained in the case of progesterone, econazole, minoxidil, and lidocaine. In the last case, multilamellar vesicles loaded with lidocaine applied on human volunteers provided a more efficient analgesia, which also lasted longer with respect to the cream formulation. Electron micrographs taken on treated guinea pigs seemed to evidence the presence of intact liposomes in the dermis (Foldvari et al., 1990). By the end of 1980s these encouraging results led to the development and approval of the first liposomal product intended for topical application, namely Pevaryl Lipogel containing econazole, to be applied once a day for the treatment of mycoses in different countries (Elsayed et al., 2007). Although many other studies claimed the potentiality of using liposomes for localized effect, the experimental setup of the first studies was questioned and several contrasting results appeared on the scientific literature about the usefulness of these systems in the field. Indeed in the same years, Ganesan and coworkers reported that the skin penetration ability of progesterone and hydrocortisone entrapped in liposomes was similar to that of free drugs and ruled out the possibility that both intact lipid vesicles and liposome phospholipids would diffuse across hairless mouse skin (Ganesan et al.,

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1984). This result was confirmed in another study performed on excised human skin. Using confocal laser scanning microscopy (CLSM) to follow the pattern of penetration of 1,2-dimyristoyl-sn-glycero-3-phosphocholine liposomes, it was derived that lipid vesicles dry on the surface of the skin forming a lipid film on the SC (Zellmer et al., 1995). Neither intact vesicles nor carried compounds were able to penetrate the skin barrier. It should be mentioned that the lipid film formed by the fusion of vesicles on the skin surface represents a further barrier to be overcome by hydrophilic drugs. Nevertheless, the applied liposomes modified the transition phase of SC lipids, suggesting an interaction between lipid components and SC ones. This finding would lead to postulate a skin penetration enhancement-like behavior for conventional liposomes. However, the pretreatment of skin with empty liposomes did not have any effect on the skin penetration of both hydrophilic and hydrophobic molecules (du Plessis et al., 1994). On the contrary, Kirjavainen and coworkers reported that the pretreatment of the skin with liposomes favored the penetration of the applied fluorescent labels. The result, however, was strongly dependent on the liposome composition (Kirjavainen et al., 1996). Differences in liposomes’ features may be the explanation for the contrasting results reported in literature about conventional liposomes, along with the experimental design of the studies. Lipid composition, surface charge, particle size, lamellarity, and method of preparation used seem to have a role, even if also in this case the available data are ambiguous. For sure lipid composition is the most important parameter. To facilitate the interaction with the SC environment, liposomes should be in the fluid state in which the mobility of the lipid chains is maximal. Lipids with transition temperature (Tm) values below the skin temperature (32 C) should be preferred. In fact as an example, the topical application on forearm of human volunteers of soy phosphatidylcholine liposomes assured a higher skin penetration of fluorescein with respect to hydrogenated phosphatidylcholine-based liposomes (Perez-Cullell et al., 2000). It should be mentioned that authors detected only the superficial penetration into the SC after stripping. For the same reason, increasing the cholesterol amount in conventional liposome formulations, the drug permeation through the skin decreases accordingly because of the reduction of the lipid chains’ mobility and the higher thickness of the bilayer ( Jain et al., 2017). Kirjavainen and coworkers demonstrated instead that the presence of the fusogenic lipid DOPE increased the depth of penetration of fluorescein in the SC (15 μm depth after 24 hours). However, the fluorescent probe did not reach the granulosum layer of the epidermis also after long exposure (Kirjavainen et al., 1996). The replacement of phospholipids with the SC lipids, namely, ceramides, cholesterol, free fatty acids, and cholesteryl sulfate (SC lipid liposomes, SCLL) seem to favor the interaction with the cutaneous tissue and to be useful for localized therapy. SCLL resulted in a higher retention of 5-aminolevulinic acid and corticosteroid drugs in the epidermis (Pierre et al., 2001; Fresta and Puglisi, 1997).

Lipid vesicles for (trans)dermal administration

Analogously the entrapment of interferon (INF) in SCLL led to a better reduction of lesions caused by herpes simplex virus type 1 compared to INF phospholipid liposomes (Egbaria et al., 1990). The impact of surface charge on the skin penetration of liposomes is more controversial. It is recognized that the lipid lamellae of the SC are preferentially formed by negatively charged lipids (due to the phosphate and carboxylate groups). For this reason, some groups have claimed that cationic liposomes would stick on the superficial layers of the skin due to the electrostatic interaction with SC lipids, rather than penetrate. Nevertheless, many experimental evidences highlighted higher performances of cationic liposomes with respect to negatively charged ones. Montenegro et al. (1996) found a greater retinoic acid skin permeation using positively charged liposomes. Similar findings were obtained by Sinico and coworkers for tretinoin (Sinico et al., 2005). The stabilization of dipalmitoylphosphatidylcholine liposomes with cationic polymers such as chitosan and Eudragit E PO resulted in an increased permeation of minoxidil and aciclovir through the skin (Hasanovic et al., 2010). Cationic 1,2-dioleoyl-3-trimethylammonium-propane-based liposomes significantly increased the permeation of doxorubicin into the deeper dermal layers of the skin (Boakye et al., 2015), whereas negatively charged liposomes remained confined to the SC. In fact what was hypothesized is that in presence of high amounts of strongly positive vesicles, the skin penetration occurs at lower rate (due to the interaction with the anionic film of the SC), but once the SC charge is neutralized, most vesicles can penetrate deeper into the skin. Particle size does not seem to be a critical property for conventional liposomes as it is for other classes of lipid vesicles. This may be easily explained by the fact that conventional liposomes do not cross the SC as intact vesicles, and therefore, their action is independent of the ability to pass through nanosized pores. As further proof of that, no significant differences in terms of drug flux and skin retention have been found using multilamellar and unilamellar liposomes, whatever lipid composition was considered (Sinico et al., 2005; Yu and Liao, 1996). Also the lipid-penetrating ability in the SC was found to be not correlated to the lamellarity and mean diameter of the vesicles (Lymberopoulos et al., 2017). Beyond the parameters that can affect the performances of topically applied conventional liposomes, and excluding the claims of the pioneering studies, it is universally recognized that conventional liposomes cannot penetrate the skin as intact vesicles; therefore, they are useless as carriers for (trans)dermal delivery purposes. The skin penetration enhancer mechanism is still questioned instead. For sure, they are solubilizing and occlusive agents and then might improve the penetration pattern of some drugs and form a depot on the skin surface. The extent of drug penetration will depend again on the physicochemical properties of the carrier compounds.

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3.2.2 Transfersomes Cevc and Blume (1992) proposed a mixed lipid aggregate with improved elastomechanical properties thanks to the elastic curvature of the bilayer which lends to a stress adaptability. These liposomes were introduced as Transfersomes. Such vesicles are composed of phospholipids, or more generally amphipats-forming bilayer, and one destabilizing agent, that can be either a compound allocated in the bilayer or a hydrophilic moiety that assembles close to the bilayer yielding to more fluid vesicles (Elsayed et al., 2018). The basic deformable liposome formulation contains a single chain surfactant in the bilayer that has affinity for curved configurations. Then when the vesicles undergo an anisotropic stress (penetration into narrow pores), the surfactant accumulates spontaneously in the most curved region, namely, the areas in which the tension is maximal, allowing the vesicles to elongate in the pore at minimum energetic cost (Fig. 3.1). Due to this peculiar behavior, it was supposed that deformable liposomes may succeed to pass through the nanosized channels of the SC, although smaller than liposomes original diameter (Cevc et al., 1998; Wachter et al., 2008). The advantage is that the change of morphology is not accompanied by a modification of the vesicle volume; this means that, in theory, deformable liposomes would maintain the drug payload while crossing the tortuous SC pathway. It is thought that “xerophobia” is the basis of deformable liposomes penetration. In other words, when applied on the skin surface, liposomes try to escape from dry environment to remain swollen, thus penetrating in the skin according to the natural osmotic gradient of the tissue (from SC with 15% water content toward the epidermis

Figure 3.1 Scheme of a vesicles with the destabilizing agent distributed in the areas of maximum curvature.

Lipid vesicles for (trans)dermal administration

with up to 70% water content) (Morilla and Romero, 2016). Deformable liposomes should then work only under nonocclusive conditions as occlusion would lead to the abolishment of the driving force for penetration. This mechanism was longer questioned as the osmotic gradient is not constant within the tissue and would limit the repartition of the vesicles from the epidermis to the deeper skin layers (Morilla and Romero, 2016). Moreover some regions of epidermis are less hydrated than the central area of the SC. As a matter of fact, Subongkot et al. found intact vesicles in the granulosum layer of the epidermis also after occlusive application of deformable liposomes on porcine skin. Using a combination of TEM and CLSM, they proposed that deformable liposomes act both as drug carriers and skin penetration enhancers through corneodesmosome degradation (Subongkot et al., 2014). Nevertheless, different experiments performed by comparing the performances of deformable liposomes under occlusive and nonocclusive conditions demonstrated that occlusive dressing causes a decrease of the efficiency of the systems (Cevc and Blume, 1992; Cevc and Gebauer, 2003; Honeywell-Nguyen et al., 2003). As further proof, in vivo, the occlusive application of a deformable liposome formulation containing ketoprofen (Diractin) led to a almost 10 times reduction of the dose of drug delivered through living mammalian skin (Cevc et al., 2008). Honeywell-Nguyen and coworkers using freeze fracture electron microscopy observed the presence of intact vesicles preferentially in the hydrophilic channel-like regions of the SC both in vitro and in vivo, supporting the hypothesis that elastic vesicles passage occurs through the channel-like regions and according to the osmotic gradient (Honeywell-Nguyen et al., 2002, 2003). Rigid vesicles instead fused and disaggregated on the skin surface. The same research group observed that the extent and the depth of intact vesicles or vesicle material penetration increases with the volume of formulation applied. Instead by prolonging the application time, vesicles fusion was observed (Honeywell-Nguyen et al., 2003). Analogously after application of deformable liposomes on human skin in vitro, we found the presence of intact vesicles up to the stratum spinosum of the epidermis (Franzè et al., 2017). Again, the vesicles moved in clusters through preferentially pathways keeping their morphology and diameter during passage. Moreover we found that the depth of penetration was function of the elastic modulus of the vesicles (Franzè et al., 2017). This is in line with what reported by Cevc, who, using CLSM, observed that the depth of fluorescence intensity depended on the deformability of the vesicles. In the case of deformable liposomes, a certain fluorescence intensity was detected between cells and propagated up to the lower part of SC, at the limit with epidermis. Using rigid liposomes or micelles, the fluorescence was restricted to the superficial layers of the skin (Cevc, 2003). After 4 hours of application of radioactivelabeled transfersomes on intact murine skin, a significant amount of label was found in the bloodstream (Cevc et al., 2003). The intact vesicle penetration hypothesis seems to

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be also supported by the capability of elastic vesicles to improve the skin penetration of both hydrophilic and hydrophobic compounds. In fact in the first case, the limiting step for permeation would be the repartition in the dry and fat environment of the SC; therefore, for hydrophilic molecules with favorable chemical physical properties, occlusion or SC perturbation may help. In contrast, hydrophobic compounds would tend to be retained in the SC due to the low affinity for the epidermis layer. In the last case, the encapsulation in a carrier able to penetrate the skin is even more essential. As an example, the encapsulation of quercetin (solubility 7 μg/mL) in deformable liposomes improved the drug deposition at epidermis/dermis level and resulted in preventing skin inflammation induced by ultraviolet ray B in vivo (Liu et al., 2013). Despite the several scientific evidences, the ability of transfersomes to squeeze through the pores of SC as intact vesicles is still questioned, and studies supporting other possible mechanisms of action can be found in the literature. Honeywell-Nguyen et al. conducted a study with elastic vesicles composed of a bilayer-forming surfactant (L-595) and a micelle-forming surfactant (PEG-8-L) containing pergolide aimed to go deep into the mechanism of action of the vesicles. They proposed that elastic vesicles can undergo a fast repartition in the SC, thanks to the great elastomechanical properties but they cannot reach the epidermis as intact vesicles. Therefore elastic vesicles bring the associated drug in the SC with them but then the diffusion of the drug toward the viable layers of the skin will occur in free drug form. The release rate of the drug from the vesicles along with the solubility of drug in the SC would determine the flux of permeation (Honeywell-Nguyen and Bouwstra, 2003). The coadministration of elastic vesicles and saturated free drug solution did not exert the same effect, thus confirming that for improving the drug penetration, the drug has to be encapsulated in the vesicles. What is never questioned is that the stress adaptability of deformable liposomes is the determinant of the success of these drug-delivery systems for (trans)dermal administration. The deformability of these systems has been studied and proven by several approaches. The basic and most used method to assess the deformability of the vesicles relies on the extrusion process. When rigid vesicles (conventional liposomes) are forced to pass through pores having a diameter half of their own size, vesicles undergo fragmentation as the applied pressure creates a tension that overcomes the rupture tension of the lipid membrane (Patty and Frisken, 2003). In contrast, deformable liposomes are able to pass through pores at least five times smaller than their diameter without experimenting particle size variation. Then the deformability of the vesicles in the simplest approach can be estimated by comparing the particle size distribution of the liposome dispersion before and after extrusion. Otherwise when the extrusion process is carried out at a constant pressure, it is possible to measure the flux of the unfragmented vesicles which increases with the square of the pores’ diameter (Cevc, 1995). Recently, Franzè

Lipid vesicles for (trans)dermal administration

et al. proposed another approach to measure the flexibility of the vesicles. In particular as the penetration is a process in which the entity that diffuses across the membrane changes the resistance opposed, we developed a modified extrusion assay in which the test is performed at constant rate, and the extruder holder is connected to a dynamometer (Franzè et al., 2017). Using this approach, it is possible to directly measure during the extrusion the trend of the resistance opposed by the membrane at the passage of the vesicles. Moreover it is possible to derive from the slope of the plot force (N) versus displacement (mm) registered by the instrument, the constant of the deformability (K, N/mm). Both these values (resistance and K) resulted very sensitive to highlight variations of elasticity of the vesicles due to formulation variables. In addition, a linear correlation was evidenced between the resistance values registered by this approach and the Young’s moduli calculated by atomic force microscopy (Franzè et al., 2017). The last one is another well-exploited approach to assess vesicles deformability. It is a very powerful technology that allows to obtain at the same time the topographic map of the sample and the maps of forces. Other techniques used are micropipette aspiration, ellipsometry, and electron spin resonance (ESR, Elsayed et al., 2018). Besides the mechanism of action, deformable liposomes resulted were able to deliver through the skin efficiently a wide range of molecules, regardless of the molecular weight. Lidocaine- and tetracaine-loaded transfersomes caused a pain-relief sensation comparable to that obtained after subcutaneous injection of the anesthetic compounds (Planas et al., 1992). The hypoglycemic effect of topically applied Transfersulin was slower but similar to the subcutaneous injection (Cevc et al., 1998). The incorporation of ethinylestradiol in deformable liposomal vesicles caused a strong inhibition of ovulation in albino rats in vivo accompanied by an increase of the thickness of the endometrius and the inhibition of the release of luteinizing hormone (effective contraception). Nondeformable liposomes and plain drug solution were tested in parallel and resulted ineffective (Garg et al., 2006). Transfersomes were also proposed for (trans)dermal vaccination. In two separate studies, it was demonstrated that deformable liposomes were able to deliver 60% of carried-hepatitis B surface antigen (HBsAg) and 50% of Plasmodium falciparum surface antigen (PfMSP-119) across human skin in vitro (Mishra et al., 2006; Tyagi et al., 2015). The immune response determined with elastic liposomes resulted comparable to that induced by PfMSP-119 injected intramuscularly (Tyagi et al., 2015). Cationic deformable liposomes were also used for transfection. After (trans)dermal application in vivo on plasmid DNA carrying vesicles on mice, gene expression was detected in several organs such as liver and lungs (Kim et al., 2004). Hattori and coworkers followed the pattern of penetration of 6carboxyfluorescein-labeled siRNA and rhodamine-labeled deformable lipoplexes through mice skin by CLSM. They observed fluorescence up to a 70 μm depth and, since the siRNA fluorescence signal overlapped that of liposomes, they suggested that siRNA lipoplexes penetrated as intact vesicles (Hattori et al., 2013).

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If molecular weight does not pose a limit to the development of efficient carriers, the encapsulation of hydrophobic drugs into transfersomes may raise some concerns. In particular hydrophobic drugs having high molecular weight or being able to establish hydrophobic interactions with the bilayer, may compromise the deformability of the systems (Perez et al., 2016). In contrast, in the case of drugs that do not intercalate strongly in the bilayer a concern of drug leakage emerges. In fact it is important to underline that the bending rigidity of the lipid bilayer increases following the same trend of the Tm. It means that to have poorly rigid membrane, transfersomes are formulated using only phospholipids with very fluid chains (low Tm). Then the release of the drug through the loose lipid network is favored still at storage temperature. To overcome this drawback, sorts of dual carrier systems have been proposed. Maestrelli et al. (2010) proposed to use cyclodextrins inclusion for the entrapment of anesthetics in deformable liposomes. However, authors focused more on the therapeutic efficacy of the novel system rather than stability over time. Recently we developed a novel system (DiMiL) in which hydrophobic drugs are entrapped in the aqueous core of deformable liposomes through micelles solubilization. DiMiL demonstrated to slow down the drug-release rate in vitro in comparison to conventional deformable liposomes, to be stable against drug leakage at least for 2 months after preparation and, more important and unexpectedly, the novel system resulted even more efficacious of deformable liposomes in delivering the model drugs through human skin (Franzè et al., 2019). The mechanism aspects of these systems are to be deepened but they opened a new frontier in (trans)dermal administration of poorly permeable compounds.

3.2.3 Ethanol-based lipid vesicles 3.2.3.1 Ethosomes Ethosomes are highly fluid lipid vesicles which contain a huge amount of ethanol (20% 45%). They were introduced by Touitou and coworkers, when it was clear that in this range of concentration ethanol does not solubilize and break lipid vesicles (Touitou et al., 2000). The use of alcohol is a further expedient to obtain “soft” carriers. In fact it is well known that ethanol is able to interact through hydrogen bonds with the polar heads of phospholipids, thus decreasing the packing of the bilayer as result of the minimized interactions and increased distance between the hydrophobic alkyl chains of phospholipids. The result is an ethanol concentration-dependent reduction of both the bilayer thickness and the Young’s modulus of the lipid vesicles (Stetter and Hugel, 2013). It was also hypothesized that an intercalation of ethanol in the bilayer was accompanied by a lowering of the surface tension of the vesicles (Mbah et al., 2019). With respect to transfersomes, ethosomes have smaller size, which is governed by the ratio of alcohol present rather than by phospholipid concentration, more negative ζ-potential and, then, higher colloidal stability (Morilla and Romero, 2016).

Lipid vesicles for (trans)dermal administration

The drug leakage concern is instead even more stressed for these carriers. Ethosomes also work under occlusive conditions making them suitable to be delivered in transdermal patches. Ethosomes show higher capability of delivering drugs in the systemic circulation with respect to transfersomes due to the dual loosening effect of ethanol on lipid vesicles bilayer and SC barrier (Fig. 3.2). Ethanol is a skin-penetration enhancer that disturbs the SC organization creating some gaps that become selective pathways for very soft vesicles, which can reach the deeper skin layers where ethosomes fuse with cell membranes and release their payload (Touitou and Godin, 2007). Moreover the evaporation of ethanol leads to a push enhancing mechanism through the increase of the thermodynamic activity (Mahmood et al., 2018). Mbah et al. (2019) using differential scanning calorimetry (DSC), observed in skin samples treated with ethosomes signs of reversible disorganization of the SC, namely, intercellular lipids fluidization and keratin denaturation. Using CLSM and fluorescence-labeled ethosomes loaded with rhodamine, Yang and coworkers observed the penetration of the payload up to the deep dermis. Nevertheless, the green fluorescence signal due to the vesicles was

Figure 3.2 Supposed mechanism of action of ethosomes. Reprinted with permission from Touitou, E., Dayana, N., Bergelson, L., Godin, B., Eliaz, M., 2000. Ethosomes—novel vesicular carriers for enhanced delivery: characterization and skin penetration properties. J. Controlled Release 65, 403 418.

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confined to the viable epidermis, suggesting that ethosomal vesicles actually break and fuse with cell at that level, and then the drug continues to diffuse in the free form. A certain penetration of both payload and vesicles was also found through the hair follicles (Yang et al., 2017). Using Nile red-loaded ethosomes, another research group suggested that vesicles even do not penetrate beyond SC but can be used to target dermatophytes in the hair follicles (Marto et al., 2016). In contrast, Elsayed et al. (2006) claimed that drug has to be encapsulated in the ethosomes to permeate because the flux of the drug after coadministration with ethosomal vesicles was significantly lower, as further proof that the performances of ethosomes do not rely merely on the enhancement effect of the alcohol. Beyond the mechanism of action, the value of ethosomes as carriers for (trans)dermal delivery has been proven in several in vitro and in vivo studies and confirmed by clinical evaluations. Campani et al. demonstrated that ethosomes facilitated the permeation of the hydrophobic vitamin K1 through porcine skin with almost 80% of the loaded drug permeating through the barrier after 24 hours. Comparing the performances of ethosomes and transfersomes, authors concluded that the latter is the most suitable carrier to limit the action of the drug in the deep skin by minimizing the systemic effect (Campani et al., 2016). Similar findings were obtained when preparing vitamin E and caffeine ethosomal and transfersomal formulations (Ascenso et al., 2015). Testosterone ethosomal patches provided 30-fold higher drug permeation in vitro through rabbit skin compared to the commercial patch Testoderm and higher testosterone blood level in vivo after 5 days of patch application (Touitou et al., 2000). Raloxifene-loaded ethosomes showed higher flux in vitro compared to both liposomal and hydroethanolic gels. What sounded strange was that the drug fluxes decreased while increasing the ethanol content in the formulation. This is in contrast with that recently reported by Mbah and coworkers, namely, that the permeation of griseofulvin increased increasing ethanol content in ethosomes formulation, even if any significant difference in the fluxes of the formulations (in the range 0% 45% ethanol content) was observed (Mbah et al., 2019). Anyway in the in vivo studies performed on Sprague-Dawley strain rats, optimized ethosomal composition improved the main pharmacokinetic parameters (Cmax and AUC) of raloxifene compared to the oral drug suspension (Mahmood et al., 2018). In another study on rats, the antiinflammatory activity of curcumin was found to be higher when the drug was administered in the form of ethosomes than as oral formulation (Inayat et al., 2018). Similarly Yan et al. (2016) found a double flux and almost fivefold higher skin deposition of sinomenine with ethosomes with respect to the ethanol solution and a significant antiinflammatory activity of the drug in vivo. Ethosomal carbomer gel was able to deliver ovalbumin into the skin of mice and stimulate specific IgG secretion, as demonstrated in in vivo immunoassays (Zhang et al., 2018). In vivo studies carried out on human volunteers demonstrated instead a less plasmatic concentration of glimepiride delivered

Lipid vesicles for (trans)dermal administration

from ethosomal transdermal film with respect to commercially available drug tablets. Nevertheless, ethosomal formulation was able to maintain an effective drug concentration in the plasma for 12 hours, being a possible alternative to the oral dosage form to reduce the side effects associated to the systemic treatment (Ahmed et al., 2016). Ethosomes were also proposed as means to treat deep skin infection minimizing the bacterial resistance due to their efficiency to localize the antibiotics in the dermis resulting in effective eradication of Staphylococcus aureus-induced intradermal infections in vivo (Godin et al., 2005). For a quick overview of the extent of the in vivo studies carried out on ethosomes readers can refer to Abdulbaqi et al. (2016) and Ashtikar et al. (2016). The number of clinical evaluation is instead more limited, and it will be discussed in Section 3.3.3. 3.2.3.2 Transethosomes Transethosomes is derived from the combination of transfersomes (deformable liposomes) and ethosomes, containing both ethanol in high percentage and another destabilizing agent such as a single chain surfactant. The synergic effect of surfactant and ethanol on the surface tension and bilayer curvature promotes the formation of irregular nonspherical vesicles (Song, 2012). Moreover as result of this peculiar composition, transethosomes are the most fluid in state among the lipid-based nanocarriers (higher elasticity), and then they show the highest capability of improving the skin permeation of the delivered drugs (Ascenso et al., 2015). Transethosomes loaded with imiquimod significantly improved the deposition of the drug in the epidermis/dermis layer in vitro and in vivo compared to both ethosomes and commercially available topical formulation. CLSM studies demonstrated a deeper penetration of lipid components of transethosomes with respect to ethosomes (Ma et al., 2015). We recently proposed hyaluronan-decorated deformable liposomes to localize nifedipine in the epidermis for the treatment of Raynaud’s syndrome. The labeling of hyaluronan on the surface of liposomes significantly compromised the flexibility of deformable liposomes. Only the combination of the single chain surfactant Tween 80 with a high percentage of ethanol (40%) allowed to maintain the softness of the vesicles, resulting in a significant improvement of nifedipine retention in human epidermis (Franzè et al., 2018). Habib et al. (2018) found a two order of magnitude higher permeation of ondansetron from transethosomes with respect to the drug suspension. Transethosomes were also successfully exploited for photodynamic therapy against resistant melanoma. The application of the vesicles on animal models led to a complete regression of small tumors without recurrence over a period of 8 months (Rady et al., 2018). As seen for the other lipid vesicles, also in this case, the mechanism of action is matter of debate. On one hand, the increased flexibility of the carrier should increase the probability of the intact vesicles penetration phenomenon. However, as usual contrasting data are reported in literature. Lei et al. (2015) using Fourier transform infrared spectroscopy stated that

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vesicles fuse with the skin as lipid fluidization and extraction were evidenced in porcine skin samples treated with transethosomes. According to the authors, the perturbation of SC barrier is the one that is mainly responsible for the enhancement properties of these vesicles. Similar findings were obtained by Ma and coworkers. Nevertheless, when caffeine was delivered by transethosomes in huge fraction as nonencapsulated hydrophilic compound, any significant difference in drug permeation was observed between vesicles and control solutions thus suggesting that the skin penetration enhancer mechanism is not sufficient to explain the performances of transethosomes. On the other hand, the small fraction of caffeine entrapped in the aqueous core of the vesicles was efficiently delivered to epidermis/dermis layer, confirming the great potentiality of transethosomes to act as drug carriers (Ascenso et al., 2015). Rady et al. (2018) claimed the presence of intact vesicles in the stratum granulosum of mice skin as revealed by TEM, although in the images provided it was not so easy to distinguish vesicles from melanosomes. Again Moolakkadath and coworkers suggested that transethosomes vesicles favored the penetration of fisetin through a fluidizing action on rat’s skin lipids. However, CLSM studies evidenced a great skin distribution of rhodamine B-loaded transethosomes vesicles up to the deeper layers. The same pattern was not observed with the rhodamine B-hydro alcoholic solution, ruling out the hypothesis that the disorganization of lipid barrier is the unique mechanism of action of transethosomes (Moolakkadath et al., 2018). 3.2.3.3 Other lipid vesicles Starting from 2008, a novel series of vesicles containing skin penetration enhancers has been proposed for improving the skin penetration of poorly permeable drugs. Invasomes are lipid vesicles enriched of terpenes which is added to unsaturated phospholipids and ethanol (3% 10% w/v) (Dragicevic-Curic et al., 2008). The basic principle behind the development of invasomes is that if terpenes, as it is known, acts perturbing the SC lipid organization, they can also have fluidizing and destabilizing effects on lipid vesicle bilayers, as already seen for ethanol. Cryo-TEM images of invasomes prepared using a mixture of terpenes (cineole, citral, and limonene) evidenced the presence of some deformed vesicle (Fig. 3.3). Nevertheless, the amount of temoporfin delivered to the human dermis was very poor after topical administration of the developed invasomes formulation (DragicevicCuric et al., 2009). Later the same research group, using ESR measurements, demonstrated that there is not a significant difference in the rotational freedom of lipid chains close to the polar heads of invasomes and conventional liposomes made of low-Tm phospholipids, but the addition of terpenes increases the overall fluidity of the bilayer, as supported also by the decrease of the Tm measured by DSC (Dragicevic-Curic et al., 2011). Using CLSM, Kamran et al. found that invasomes loaded with a hydrophobic dye were able

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Figure 3.3 Cryo-electron micrographs of invasomes with 1% terpenes mixture: (A) unilamellar (short arrows) and bilamellar (arrows of medium length) invasomes; (B) long arrows represent deformed vesicles. Reprinted with permission from Dragicevic-Curic, N., Gräfe, S., Albrecht, V., Fahr, A., 2008. Topical application of temoporfin-loaded invasomes for photodynamic therapy of subcutaneously implanted tumours in mice: a pilot study. J. Photochem. Photobiol. B Biol. 91, 41 50.

to increase the depth of the dye penetration up to 168 μm. In vivo olmesartan-loaded invasomes led to a decrease of Cmax compared to the oral tablet but the AUC results improved (Kamran et al., 2016). When the performances of invasomes were compared to those of ethosomes, the last carrier resulted anyway the most efficient in delivering the carried drug through the skin (Chen et al., 2010). After terpenes a wide range of skin penetration enhancers was tested to obtain modified liposomal vesicles, named as PEVs. Both hydrophobic and hydrophilic penetration enhancers are used, even if the former is preferred as they can intercalate in the bilayer and induce the most changes in vesicles characteristics, whereas the latter can interact only with lipid polar heads (Manconi et al., 2012). In fact PEVs prepared with transcutol did not show improved deformability with respect to conventional liposomes according to extrusion assay, although, as previously seen with invasomes, they favored the transition toward the fluid state of vesicles (Manconi et al., 2011a,b). In contrast, the increase of the hydrophobicity of the penetration enhancer (e.g., labrasol) was accompanied by a more prominent alteration of the lipid packing with an increase of both bilayer softness (as demonstrated by SAXS) and index of deformability (Caddeo et al., 2015). The enhanced fluidity of the bilayer along with the disorganizing effect provided by free penetration enhancer on the skin barrier was claimed to be responsible for the improved penetration of the carried compound. Moreover SEM images collected on porcine skin treated with PEVs showed a higher hydration of the SC which causes an enlargement of the intercellular SC regions (Manconi et al., 2011a,b). The application of PEVs favored the deposition in the epidermis of diclofenac both in the acid and salt form, whereas the pretreatment of porcine skin with empty PEVs had an opposite effect, suggesting that PEVs can act as real drug carriers (Manconi et al., 2011a,b). When compared to the solution and the commercially available topical formulation

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(Voltaren), PEVs were capable of enhancing dermal delivery of diclofenac. Moreover on mice the in vivo application of PEVs led to a better antiinflammatory activity with a significant and visible reduction of the lesions induced by 12-Otetradecanoylphorbol13-acetate, more than both ethosomes and Voltaren (Caddeo et al., 2013). PEVs were able to increase the localization of different drugs in the skin, among those also minoxidil and tretinoin (Mura et al., 2009, 2013; Manconi et al., 2011a,b). Nevertheless, using CLSM it was recently evidenced that PEVs have less capability of penetrating themselves into the skin compared to transfersomes. PEVs seem to remain in great extent on the SC creating a local depot that sustains the drug release (Manca et al., 2019).

3.3 Liposomal formulation in clinics Browsing the literature resources, more than 3000 articles dealing with the in vitro and preclinical data about lipid vesicles for topical and transdermal application can be found, but among those not more than 100 papers report results on clinical trials performed on such systems. Moreover few topically applied products have been really reaching the market in the industrialized countries as cosmetics, medical devices, or medicinal products. Indeed excluding the liposomal products marketed as cosmetics ( Jain et al., 2017), less than 10 products containing lipid vesicles have been authorized as topically applied medicinal products or medical devices in the of Organisation for Economic Co-operation and Development countries [e.g., Canada, European Union (EU), and Israel]. Such products are thought for the locoregional treatment of human diseases or symptoms, such as infections (e.g., antimycotics and antivirals), pain [e.g., local anesthetics and nonsteroidal antiinflammatory drugs (NSAIDs)], and Peyronie’s disease (e.g., enzymes). The main outcomes of the clinical data available, divided by class of lipid vesicles (conventional liposomes, transfersomes, and ethosomes), are briefly described in the following sections.

3.3.1 Conventional liposomes All the researches carried out on the topical application of conventional liposomes were traduced in the commercialization of three products, for local treatments, namely, Pevaryl Lipogel, LMX4, and ELA-MAX. Pevaryl Lipogel was the first topical liposomal product approved for the market in different European (i.e., Switzerland, Italy, Belgium, and Norway) and Central America countries (i.e., Costa Rica, Dominican Republic, El Salvador, Guatemala, Honduras, Mexico, Nicaragua, and Panama), for the treatment of cutaneous mycoses. It is a

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conventional liposomal formulation containing econazole which assured a faster onset of action with respect to 1% econazole-loaded cream allowing to modify the regimen of administration from twice-a-day to once-a-day (Naeff, 1996). It should be also underlined that the few and quite old clinical data available in the literature are related to the eradication of tinea corporis or tinea pedis (Szepietowski et al., 1994; Wilkowska et al., 1995). LMX4 lidocaine 4% w/w cream or ELA-MAX 4% w/w cream, which is now on the market with the trade name Maxilene cream, are topical dosage forms containing lidocaine used for local anesthesia. Such formulations were authorized on the basis of clinical evidences that demonstrated a superior efficacy with respect to conventional semisolid preparation (e.g., creams and ointments). In 1990, Foldvari and coworkers found that lidocaine-loaded liposomes permitted not only a massive drug penetration in the dermis but also higher and longer anesthesia in comparison to conventional lidocaine-loaded cream (Foldvari et al., 1990). Indeed high painless scores at pinprick test were recorded when human volunteers were treated with liposomal lidocaine for 1 hour in occlusive condition with respect to nonliposomal drug-loaded cream. The efficacy of liposomes as local anesthetic carriers was demonstrated by further studies both in children (Taddio et al., 2005) and in adults (Bucalo et al., 1998). In particular Bucalo et al. demonstrated that a 5% w/w lidocaine liposomal cream guaranteed higher anesthesia than equivalent liposome-free cream and ointment when they were applied for 30 min in both occlusive and nonocclusive conditions. The application time could be further reduced to 5 min, maintaining the same anesthetic effect during a venepuncture, when the application of liposomes followed the pretreatment of the patients with physical methods (e.g., ultrasounds) (Zempsky et al., 2008). These evidences suggested that the efficacy of liposomal systems was strongly influenced by the time elapsed between the topical application and the painful intervention. Other works demonstrated that it was possible to reduce the lidocaine dose from 5% to 4% without altering the efficacy profile of the topical treatment (Bucalo et al., 1998; Tang et al., 2004), improving the advantages of using liposomes instead of the conventional creams (at 5%). Indeed the performances of 4% and 5% liposomal cream formulations were also compared in several studies to a 5% w/w eutectic combination of lidocaine and prilocaine marketed under the trade name of EMLA cream which is considered the gold reference for the local anesthesia. The 5% drug-loaded liposomal cream showed a comparable pain control to EMLA cream at pinprick test (Bucalo et al., 1998). In contrast the clinical results of 4% liposomal cream are not univocal. In a work focused on treatment of heat-induced pain, the performances of lidocaineloaded liposomal resulted slight inferior to the eutectic combination after 60-minutes application in occlusive conditions (Tang et al., 2004). In another paper, its anesthetic

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effect was comparable to EMLA in reducing pain during the electrodesiccation of benign and malignant skin lesions (Carter et al., 2006). Despite these contradictory results that may be due to the different trial design, it is noteworthy that the liposomal 4% lidocaine cream permits to obtain a comparable efficacy to the eutectic composition loaded with the 5% of drug with a lower drug-applied dose, which could be related to a slight reduction exposition to local anesthetics. Recently the liposomal lidocaine medical products were also tested for reducing pain during vaccination. The basic idea of researchers is to minimize the cost of vaccines and needles in children and, therefore, to increase the adherence of population to vaccination program. Taddio and colleagues studied the ability of liposomal 4% lidocaine cream to prevent injection-induced pain in adults (Taddio et al., 2010) and infants, respectively (Taddio et al., 2017). In adults such a nanomedicine strategy resulted superior to the self-distraction of the patient but showed comparable efficacy to tactile stimulation of injection site, made by a nurse immediately prior the vaccine injection, or to the administration of anesthetic-loaded vapocoolant spray (Taddio et al., 2010). On the other side, the topical application of liposomal lidocaine resulted superior to other conventional pain-minimizing and soothing methods in infants (Taddio et al., 2017). Conventional liposomes were also tested as carriers for enzymes. In particular a recombinant Cu/Zn human superoxide dismutase liposomal gel (trade mark: Lipoxysan) was developed for the management of Peyronie’s disease (Riedl et al., 2005). The etiopathology of Peytonie’s disease involves the penile corpora carvernosa inducing pain, plaque formation, penile deviation, and erectile dysfunction. After 8 weeks of twice daily topical application, the liposomal preparation reduced the pain in more than 80% of treated subjects. These clinical outcomes were attributed to the reduction of reactive-oxygen species-induced inflammation by the enzyme amounts penetrated in the penile tunical tissues. However, minimal changes in plaque characteristics and in other clinical outcomes were observed after 8 weeks, allowing authors to speculate that longer treatment time and dose assessment should be tested. In addition to the different efficacy outcomes, the available studies reported no local and systemic side effects, suggesting that the topical application of liposomal systems may be relatively safe (Riedl et al., 2005).

3.3.2 Transfersomes Following the studies of Cevc and colleagues, the possible clinical application of transfersomes as carriers of antiinflammatory drugs was in depth studied for the symptomatic treatment of locoregional diseases, such as musculoskeletal disorders (e.g., osteoarthritis, muscle soreness following exercises, and eccentric muscle contraction). In addition to the disease etiopathology, most of the musculoskeletal disorders cause

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pain and movement discomfort or disability, affecting the quality of life of patients. The current clinical treatment aimed to relieve pain, to reduce muscle/joint stiffness, and to restore their physiological function. Oral NSAIDs are the gold standard treatment for most of musculoskeletal disorders. However, their chronic use is associated with several systemic adverse effects, such as hepatotoxicity (paracetamol), cardiovascular events (selective and nonselective NSAIDs), gastrointestinal complications. To overcome these issues, Rother and colleagues performed several clinical trials (more than 1600 patients involved) to assess the efficacy of ketoprofen-loaded transfersomal gel versus conventional oral NSAIDs (e.g., naproxen and celecoxib) in patients affected by knee osteoarthritis (Conaghan et al., 2013; Kneer et al., 2009, 2013; Rother et al., 2009; Rother and Conaghan, 2013). The product efficacy was assessed using the Western Ontario and McMaster University (WOMAC) Osteoarthritis index for scaling pain, physical function, and patient global assessment of responses. The overall results showed that the efficacy of ketoprofen-loaded transfersomal preparations was superior to oral placebo and comparable to the oral celecoxib. Contrasting results were obtained comparing the efficacy of different applied doses of drug-loaded transfersomal preparations (Conaghan et al., 2012; Kneer et al., 2013). Similar results were obtained with the same preparations in the treatment of eccentric muscle contractions and muscle soreness induced by exercise (Rother et al., 2009; Seidel et al., 2016). The ketoprofen-loaded transfersomal gel was well tolerated by patient: only transient skin irritation and erythema were observed after application in all the clinical trials performed. Based on these results, the clinical data of ketoprofen-loaded transfersomal gel was submitted to EMA to obtain the market authorization (MA) in the EU under the trade name Diractin. However, the application was withdrawn by the applicant as the clinical efficacy of Diractin was not sufficiently documented (EMEA, 2008). Interestingly the clinical trials performed to investigate the efficacy and safety of ketoprofen-loaded transfersomes highlighted that also drug-free transfersomes, called Sequessome vesicles, resulted effective as well in reducing pain and stiffness of osteoarthritic knee, other than improving joint function (Conaghan et al., 2014; Rother and Conaghan, 2013). In particular the metaanalysis carried out by Conaghan and coworkers demonstrated that the clinical effect of Sequessome vesicles was noninferior to oral celecoxib and superior to the placebo. Moreover these drug-free nanovesicles were well tolerated by patients. The mechanism of action seemed related to biolubrification of damaged joints due to the permeation of drug-free transfersomes and their localization at the cartilage surface. Based on these findings, the drug-free transfersomal gel preparation was registered in EU and marketed as a medical device (trade mark: FLEXISEQ) for the topical treatment of osteoarthrosis (NICE, 2016). On the contrary, a clinical trial aimed to evaluate the efficacy of Sequessome vesicles in treating cutaneous inflammatory diseases, such as Rosacea subtype 1, showed a modest efficacy of such nanomedicine product in comparison to placebo (Luger et al., 2015).

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3.3.3 Ethosomes The first medicinal product (i.e., Supra-vir cream) containing ethosomal technology was marketed based on the clinical data supporting its efficacy in promoting the permeation of acyclovir in labial herpetic infection. In particular a two-arm, doubleblind, and randomized clinical trial was performed including 40 volunteers (Horwitz et al., 1999). It was demonstrated that the novel 5% acyclovir ethosomal preparation, applied four times per day, was more effective than both placebo vehicle and a conventional 5% drug-loaded cream. The time to crusting of lesion and to loss of crust were selected as clinical endpoints of the study. In particular time to crust formation was reduced to 1.6 6 1.4 days, which was a significant and clinical-relevant shorter period in comparison to placebo vehicle (4.8 6 2.1 days) and conventional cream (4.3 6 1.9 days). Moreover the time needed to crust loss followed a similar trend: ethosomal preparation (3.5 6 3.1 days) , placebo (6.1 6 3.1 days)  conventional cream (6.4 6 3.2 days).

3.4 Final remarks Despite the wide pool of preclinical data available in literature claiming the worth of fluid and deformable lipid vesicles for (trans)dermal drug delivery, clinical data seem to suggest that the goal of using lipid vesicles to assure plasma concentrations of drug suitable for the treatment of systemic pathologies is far to be reached. This discrepancy can be ascribed to the experimental design of the preclinical studies and, mainly, to the selection of animal models having a skin barrier significantly different from human one, in terms of the distribution, lipid composition, and thickness of hair follicles. As a matter of fact, the data obtained using animal models can be transferred to humans, but with difficulty. liposomal products open instead a new frontier for the development of medicinal products to be used for the treatment of locoregional diseases and relative symptoms and for cutaneous pathologies, thanks to the possibility to target drugs to a particular skin layer or cellular population.

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CHAPTER FOUR

Stimuli-responsive nanocarriers for drug delivery Maria Chountoulesi1, Nikolaos Naziris1, Natassa Pippa1,2, Stergios Pispas2 and Costas Demetzos1 1 Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece 2 Theoretical and Physical Chemistry Institute, National Hellenic Research Foundation, Athens, Greece

4.1 Introduction Stimuli-responsive nanosystems are an emerging platform for numerous applications in the field of nanotechnology and therapy of diseases. This technology is the realization of one of the most important scientific efforts, which aims to mimic the responsive behavior of cells and living organisms to all the surrounding physical and chemical stimuli. In this approach, advanced applications are built for biomedical purposes, including drug delivery and release of bioactive molecules to pathological tissues, imaging, and diagnosis, as well as the promising field of theranostics. Although conventional nanotherapeutics have shown promising results in preclinical and clinical practice, the precisely controllable release of their content exactly to the targeted pathological tissue remains a challenge. For example, one of the major limitations is the undesired content release in healthy tissues, causing toxicity from side effects. Stimuliresponsive nanosystems are designed to exploit in a “smart” way the altered conditions (e.g., temperature, pH, and enzyme concentration) that take place in pathological tissues compared to the physiological ones, facilitating triggered content release in the targeted tissue, overcoming the intermediate barriers and resulting in enhanced bioavailability, prolonged blood circulation time, and overall increased therapeutic efficacy. Stimuli-responsive nanocarriers are developed by utilizing functional nanobiomaterials, which may be natural or synthetic and usually belong to the class of soft materials (Naziris et al., 2016; Hruby et al., 2015; John et al., 2015; Viard and Puri, 2015; Liu et al., 2017a). Molecules such as polymers, lipids, and peptides that are biocompatible, responsive to particular conditions, and exhibit self-assembled properties have been characterized as “smart” molecules. In pharmaceutics, such ingredients belong to the class of functional excipients and are studied in formulations that will

Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00004-7

© 2020 Elsevier Inc. All rights reserved.

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enhance the therapeutic effect of therapeutic molecules, while reducing their toxicity (Mura et al., 2013; Alvarez-Lorenzo and Concheiro, 2014). The mechanism of stimuli responsiveness is of great importance in nature and relates to the fundamental concepts of natural systems, including communication and survival of living cells and organisms. It is indeed of great interest to decode, analyze, understand, and utilize this mechanism, which is complex and multifactorial, for building “smart” and functional nanodevices that will be able to realize demanding but promising biomedical applications. A characteristic example is the utilization of stimuli-responsive nanocarriers for targeted drug delivery. These platforms are developed by combining biomaterials that are biocompatible and biofunctional and self-assemble to supramolecular complexes that can transport therapeutic bioactive molecules to specific pathological tissues, based on the abnormal physiological conditions that they exhibit. The systems are bioinspired because they mimic the behavior of living systems when meeting altered conditions in their surrounding environment. The currently utilized conditions include not only endogenous ones, originating from inside the cells or tissues, such as temperature fluctuations, pH variations, redox potential alterations, and differences in the ionic strength, but also exogenous ones, generated from artificial sources, such as heat, light, ultrasounds, and magnetic or electric fields. In this rationale, various classes of nanocarriers can be rendered stimuli-responsive, including liposomes, liquid crystals, polymersomes, micelles, niosomes, dendrimers, and nanocarrier drug conjugates (Naziris et al., 2016; Liu et al., 2017a; Karimi et al., 2016; Lee and Thompson, 2017; Lee and Nguyen, 2013; Li and Xie, 2017; Qiao et al., 2019). In this chapter, we discuss stimuli-responsive liposomes and nonlamellar lyotropic liquid crystalline nanosystems, with emphasis on pH and temperature. Moreover we focus on nanocarriers that have been developed by combining two different categories of biomaterials, namely, lipids and stimuli-responsive polymers. In addition, results from certain methods of analysis and structural characterization are presented.

4.2 Types of stimuli Stimuli exist in the human body but may also be generated from external artificial sources. Endogenous and exogenous types of stimuli have both been explored for the development of advanced and bioinspired nanosystems with functional properties that arise from responsiveness to microenvironmental conditions. Stimuli have been also categorized as intracellular and extracellular, physical, chemical/biochemical, and biological/physiological. Among the different types of stimuli, shear stress and

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Figure 4.1 Physical and chemical stimuli that originate from the cell microenvironment or from external artificial sources, including pH, temperature, light, enzymes, redox potential, ultrasounds, and so on, are exploited for various therapeutic applications by triggering stimuli-responsive nanosystems. Adapted from Karimi, M., Ghasemi, A., Sahandi Zangabad, P., Rahighi, R., Moosavi Basri, S.M., Mirshekari, H., et al., 2016. Smart micro/nanoparticles in stimulus-responsive drug/gene delivery systems. Chem. Soc. Rev. 45 (5), 1457 1501.

temperature fluctuations may occur from both external sources and internal conditions (Karimi et al., 2016; Ge and Liu, 2013; Fleige et al., 2012). A number of very common types of stimuli are shown in Fig. 4.1.

4.2.1 pH-responsive nanosystems The concept of pH responsiveness is one of the most well investigated in the field of stimuli-responsive biomaterials and drug delivery nanosystems. The smartness of the pH-responsive systems is based on the fact that they are able to exploit wellcharacterized pH differences inside the human body, such as pH differences existing between normal blood and pathological tissues (e.g., due to infection, inflammation, and cancer), between certain intracellular compartments (i.e., cytosol, endosomes, and lysosomes), as well as along the gastrointestinal track. The acidic pH is characteristic of specific pathological states like infections, inflammations, ischemia, and tumors (Liu et al., 2014; Felber et al., 2014; Kanamala et al., 2016). Stimuli-responsive nanosystems can also be characterized as bioinspired, because of the viruses, which exploit the intracellular pH gradient through mobilization of their fusogenic peptides, to infiltrate mammalian cells (Hughson, 1995; Lin et al., 2010). Developing pH-sensitive nanosystems against tumors is considered to be an attractive strategy for anticancer research. According to the phenomenon called as “Warburg effect,” tumor cells exhibit rapid proliferation that leads to disorganized

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vasculature and subsequently to hypoxia and insufficiency of their nutritional needs (Bhattacharya et al., 2016). As a result, propensity for glycolytic metabolism, glycolysis, and lactic acid fermentation take place in the cytosol, instead of oxidative phosphorylation in the mitochondria, increasing proton production. The combination of increased proton production, poor lymphatic drainage, and elevated interstitial pressure of cancer tissues causes acidosis of these cells, with extracellular pH dropping to as low as 5.7, for certain types of tumors (Stubbs et al., 2000; Tian and Bae, 2012). Regarding the pH differences taking place between certain intracellular compartments, the pH-responsive systems can also exploit the “endosomal effect” (Xu et al., 2015). When drug carriers are taken up by cells actively, through receptor-mediated endocytosis, they are trapped within the early and late endosome (pH 5 5.0 6.5), which later matures to lysosome (pH 5 4.5 5.0) with the hydrolytic enzymes. The pH-responsive systems can respond to this extreme acidic pH, by releasing their content to the cytosol, avoiding the lysosomal degradation and enhancing the bioavailability (Felber et al., 2014; Kanamala et al., 2016). There are three different categories of stimuli-responsive nanosystems, namely, nanocarriers with protonable/deprotonable groups, nanocarriers with acid labile bonds in the polymers that compose them, like the hydrazine bond, and nanocarriers with pH-responsive “PEG-detachment.” We focus on the first category, which is the most investigated. The biomaterials with protonable/deprotonable groups can be pH-sensitive polymers, either anionic [poly(acrylic acid), PAA] or cationic [poly(2dimethylamino)ethyl methacrylate, PDMAEMA], lipids (dioleoylphosphatidylethanolamine, DOPE), “cage” lipid derivatives (C-DOPE), fusogenic peptides (GALA), and polysaccharides (chitosan derivatives) (Kanamala et al., 2016; Liu and Huang, 2013). Many different types of pH-sensitive nanosystems have been reported since 1990 (Chu et al., 1990; Collins et al., 1990a,b; Liu and Huang, 1990) until now, such as liposomes (Kono et al., 1994a), niosomes (Roux et al., 2002), polymersomes (Du et al., 2012), lipoplexes (Mignet et al., 2008), polymeric micelles (Leroux et al., 2001), polymeric nanoparticles (Du et al., 2014), virosomes with cell-penetrating peptides (Bron et al., 1993), oral insulin delivery promising pH-sensitive hydrogels (Demirdirek and Uhrich, 2015), as well as dual-sensitive hydrogels (pH and thermo) utilized for encapsulation of cardiosphere-derived cells and stem cell cardiac therapy (Li et al., 2016).

4.2.2 Thermoresponsive nanosystems The temperature as a stimuli is based on the fact that the temperature is elevated (40 C 42 C) in tumors, due to the high rates of aerobic glycolysis and fast proliferation (Qiao et al., 2019). The “smartness” of the thermoresponsive nanosystems is based on the ability of their biomaterials to undergo temperature-dependent phase transition and to induce conformational changes in the system structure and morphology that

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promotes the controlled release of their content. Apart from targeted drug delivery, the thermoresponsive nanosystems can also be applied for diagnosis, monitoring, and treatment of various diseases. Poly(N-isopropylacrylamide) (PNIPAM) is one of the most famous thermoresponsive polymers, whose solubility in water depends on temperature. The mechanism of its thermoresponsiveness will be discussed analytically in the following section. It has been utilized in a wide range of technological platforms for biomedical purposes, such as thermoresponsive controlled delivery of bioactive molecules, gene delivery, tissue engineering, self-healing, bioimaging, three-dimensional bioprinting, and other applications. It has been used in a variety of thermoresponsive drug delivery nanosystems, such as liposomes, micelles, polymersomes, films, particles, as well as hydrogel development (Lanzalaco and Armelin, 2017; Ward and Georgiou, 2011). Liposomes are an ideal vehicle, to develop thermoresponsive platforms, because the intrinsic thermodynamic properties of the phospholipid, like the main transition temperature Tm, serve for modulatory controlled drug delivery. Thermoresponsive liposomes have already been introduced in clinical phase III for the delivery of doxorubicin to primary liver cancer, under the trademark Thermodox, containing a lysolipid that provides thermoresponsiveness, categorized as a lysolipid thermally sensitive liposome technology (Sánchez-Moreno et al., 2018; Hongshua et al., 2019; Jhaveri et al., 2014; Kneidl et al., 2014). Alternative approaches for the development of thermoresponsive liposomes include the use of thermosensitive zipper peptides and bubble-generating liposomes. Bubble-generating liposomes, combined well with local hyperthermia and magnetic resonance imaging (MRI), have also improved ultrasound imaging (Al-Ahmady et al., 2012; Chen et al., 2013). However, thermoresponsive liposomes can also be developed by the incorporation of functional polymers to provide a temperature-induced membrane disruption and consequently a controlled release of their content. The chimeric thermoresponsive liposomes containing thermoresponsive polymers will be discussed analytically in the following section. Other thermoresponsive platforms, which have been reported, are pluronic/poly (ethylenimine) nanocapsules exhibiting a thermally reversible swelling/deswelling volume expansion/contraction, for siRNA delivery after endosomal breakage, as “cryotherapy” or “cold shock” (Lee et al., 2008) and “smart nanobombs” that they can “blow up” the tumor cells, releasing the encapsulating DOX inside the tumor tissue (Lee et al., 2009).

4.3 Development of chimeric stimuli-responsive liposomes with incorporated stimuli-responsive polymers Liposomes are considered to be one of the most well-investigated drug delivery nanosystems, presenting major advantages, such as biocompatibility, biodegradability,

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and lack of immunogenicity, as well as their ability of passive targeting to tumors by exploiting the enhanced permeability and retention effect. Unfortunately the conventional liposomal platforms exhibit serious disadvantages, such as their rapid recognition and uptake by reticulo-endothelial system that reduce their plasma half-life, their endosomal trapping with subsequent degradation by lysosomal enzymes after endocytosis, as well as the content leakage that sometimes occurs before they reach the target tissue. In an effort to overcome the abovementioned limitations, stimuli-responsive forms of liposomes have been developed (Eloy et al., 2014). Stimuli-responsive liposomes and especially pH-responsive and thermoresponsive ones have been lately developed by attaching functional polymers on the surface of conventional liposomes, therefore producing chimeric/mixed advanced drug delivery nanosystems (aDDnSs) (Demetzos and Pippa, 2014). The response of the liposome to the stimuli yields a structural change. More analytically this approach is promising to effectively control the membranes disruption or alteration of the surface properties of liposomes, leading to utilities like the targeted and controlled release of therapeutic moieties for various diseases, for example, drug molecules. Because of the fact that this can be achieved in a spatiotemporal way, the amount of drug that reaches the diseased tissues is increased, while healthy tissues are bypassed, resulting in increased effectiveness and reduced toxicity of the drug molecule (Lee and Thompson, 2017; Lee and Nguyen, 2013). In the following sections, we emphasize on the development and evaluation of stimuli-responsive liposomes with incorporated functional stimuli-responsive polymers.

4.3.1 pH-responsive liposomes There are a number of ways to develop pH-responsive liposomes, for example, by using pH-sensitive lipids and fusogenic peptides/proteins or by anchoring pHresponsive polymers on the lipid bilayer, usually through hydrophobic alkyl chains (Felber et al., 2014). We focus on the incorporation of different kinds of pHresponsive polymers, which transforms the conventional liposomes to pH-responsive chimeric ones. pH-Sensitive liposomes are designed to remain stable at physiological pH (7.4), but destabilize due to polymer chain conformational changes under pathological acidic pH conditions. More analytically, in acidic pH, the polymers acquire fusogenic properties due to their ionizable chemical groups and change their conformation. The pH-responsive polymers are classified to anionic (exhibiting, e.g., carboxylic acid groups) and cationic (exhibiting, e.g., tertiary amine groups). Anionic polymers with carboxylic pendant groups, such as PAAs, are commonly used for targeting the acidic pH of tumor tissues, where they get protonated (nonionized) and hydrophobic, causing a shrinkage of the polymer chains and leading eventually to destabilization of the membrane and content release. At physiological pH 7.4, their groups are deprotonated (ionized) and hydrophilic. The negative charge of the

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carboxylic groups at physiological pH makes liposomes to repel each other, preventing fusion and thus destabilization. The tertiary amine group of cationic polymers, such as of the PDMAEMA, remains neutral under basic conditions, but it gets positively ionized in acidic conditions, by accepting protons. Thus the positive charge in acidic environment, that pH-responsive liposomes acquire due to the incorporated polycations, may induce cellular uptake via the negatively charged membranes of pathological tissues with lower values of environmental pH and also facilitate endosomal escape due to the proton sponge effect (Liu et al., 2014; Felber et al., 2014; Kanamala et al., 2016; Lin et al., 2010; Kim et al., 2016).

4.3.2 Thermoresponsive liposomes The incorporation of thermoresponsive polymers into liposomes has been proven to be a useful strategy toward the development of chimeric thermoresponsive liposomal systems. Thermoresponsive polymers exhibit lower critical solution temperature (LCST) in aqueous solutions, indicating water solubility below this temperature. Increasing temperature through LCST causes coil-to-globule conformational phase transition, due to the breakage of the hydrogen bonds existing between the polymer segments and water molecules, yielding dehydration and thus a decrease of the hydrophilicity of the polymer. By incorporating this kind of polymers into the liposomes, the abovementioned temperature response of the polymer causes an increase of the hydrophobicity also at the liposomal surface, as well as destabilization of liposomes and release of their contents, but only at the pathological tissue with the increased temperature. In contrast, below the LCST, where the polymer is in a highly hydrated state, the polymer chains, stretching out of the liposomal membrane, stabilize the liposomes, protect them partly from blood proteins, improve circulation time in the blood stream, and prevent aggregation phenomena (Lanzalaco and Armelin, 2017; Ward and Georgiou, 2011; Eeckman et al., 2004; Liu et al., 2009, 2017b; Kono, 2001; Kono et al., 1994b; Chountoulesi et al., 2017). PNIPAM is considered to be one of the most popular thermoresponsive polymers and exhibits a LCST at around 32 C (close to human temperature) making PNIPAM ideal for biomedical applications, such as the development of chimeric liposomal drug delivery systems. It also exhibits pHresponsive properties. At temperatures near its LCST, PNIPAM undergoes a spontaneous reversible and endothermic, coil-to-globule transition driven by entropy gain. At temperatures below LCST, the acrylamide groups of the block interact via hydrogen bonding with water molecules of the medium, enhancing the solubilization of the polymer chains. Above the LCST, the hydrogen-bonded network is instantaneously disrupted resulting in the release of the water molecules out of the polymer chain, while the increase of interactions among the polymer chains results in the formation of insoluble globular aggregates (Lee and Nguyen, 2013; Heskins and Guillet, 1968;

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Futscher et al., 2017; Pelton, 2010; Southall et al., 2002). Moreover the LCST of PNIPAM and related biomaterials can be modified, to be close to physiological temperature values (37 C), by copolymerization with other hydrophilic monomers, via control of the molecular weight and architecture and finally, through control of the overall hydrophilicity of the final macromolecule (Liu et al., 2009; Eeckman et al., 2004).

4.4 Thermotropic behavior of stimuli-responsive liposomes Liposomes have been widely analyzed by applying thermoanalytical techniques, especially differential scanning calorimetry (DSC). DSC is a method by which a test and a reference sample are heated/cooled with the same heating/cooling rate, at the same time. The test sample is expected to produce thermal phenomena in a designated temperature range. This happens when differential amount of heat is absorbed or emitted from the two samples, resulting in endothermic or exothermic phenomena, respectively. The mechanism behind this is that the heat capacity of the sample exhibits variations as a function of temperature, compared with the reference, which are monitored real time and are proportional to the energy changes of the system, in this case heat flow. An example of endothermic phenomenon is the melting process and in the case of lipid bilayers and liposomes, the main phase transition from the gel phase (Lβ’) to the liquid crystalline phase (Lα) (Demetzos, 2016; Huang and Mason, 1986). By utilizing DSC on lipid bilayers and liposomes, it is possible to define the various mesophases or mesomorphs that occur in a range of temperature, due to polymorphism, because these are lyotropic liquid crystals, for which both temperature and concentration of various additives affect their phase behavior. The identification and study of these mesophases of liposomal dispersion systems lipid bilayers allows the control over their thermodynamic properties, to rationally design the liposomal system with the most satisfactory physical and thermal stability (Demetzos, 2008). What is also interesting is to identify the metastable phases between the gel and the liquid crystalline phases that arise in lipid bilayers and liposome membranes, induced by incorporated molecules and various additives in the surrounding environment. These have been linked with the biophysical properties and behavior of liposomes (Pippa et al., 2015b).

4.4.1 Thermal analysis on pH-responsive liposomes DSC has been applied on liposomal platforms with incorporated pH-responsive amphiphilic diblock copolymers. This approach has allowed for the identification of

Stimuli-responsive nanocarriers for drug delivery

various phases that are created by this integration, but also how membranes respond to different pH conditions on a thermodynamic level. In a particular study, bilayers of the phospholipid L-a-phosphatidylcholine, hydrogenated (soy) (HSPC) with incorporated poly(2-(dimethylamino)ethyl methacrylate)-b-poly(lauryl methacrylate) (PDMAEMA-b-PLMA) amphiphilic diblock copolymers were developed and studied through DSC (Naziris et al., 2018). Two different in composition copolymers were utilized, PDMAEMA-b-PLMA 1 of weight composition 59% 41% and PDMAEMA-b-PLMA 2 of weight composition 46% 54%. The thermotropic profiles of the chimeric nanosystems were obtained in two different pH environments, in phosphate-buffered saline (PBS) (pH 5 7.4) and in citrate buffer (pH 5 4.5) media, to evaluate their stimuli responsiveness. The results indicated that the insertion of the copolymers led to a homogeneous distribution inside the membranes, which in the case of PBS did not affect their fluidity, but increased their cooperativity and reduced their effective transition enthalpy. However, when exposed to lower pH conditions, both copolymer-induced perturbation and possible disruption in the membranes, diminishing the effective transition enthalpy and affecting all thermodynamic parameters in a concentration- and composition-dependent manner. Naziris et al. (2017) utilized another pH-responsive polymer, consisting of PAA, from the category of anionic pH-responsive polymers. More specifically the pHresponsive amphiphilic homopolymer, the C12H25-poly(acrylic acid) (C12H25-PAA) was incorporated into HSPC bilayers and then the resulting liposomes were loaded with indomethacin (IND). The bilayers were fully hydrated with two different dispersion media, high-performance liquid chromatography (HPLC)-grade water (with pH 5 4.5) and PBS (with pH 5 7.4). The incorporation of C12H25-PAA caused alterations of the thermotropic behavior of HSPC lipid bilayers, affecting extensively the specific enthalpy ΔHm of the main transition in both dispersion media in a concentration-dependent manner but not the main transition temperature Tm, indicating that the interactions of the PAA polymeric segment with HSPC lipids affect the mobility of the polar head groups of lipids. The pretransition was generally eliminated, the cooperativity was decreased especially in HPLC-grade water, but no other remarkable differences were caused by the different dispersion medium. The temperature at which the thermal event starts (Tonset,m) changed significantly only for the systems of the highest polymer amount. Considering the amphiphilic nature of the homopolymer and the hysteresis being observed at the cooling curves, the authors suggested a formation of a new metastable phase, probably of an interdigitated phase, especially at the highest molar ratios of polymer. In these ratios, the increased polymer concentration can cause membrane disruption and phase separation. The calorimetric profile of the bilayers was correlated with the release kinetic profile of the IND drug from the respective pH-responsive liposomes. The abovementioned metastable phase probably acted as a drug-release “barrier” for one of the chimeric nanosystems, exhibiting

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slower in vitro release of IND, comparing with liposomes using lower molar ratios of C12H25-PAA. These differences in the release rate as well as at the total percentage of drug release were attributed to the pH responsiveness of C12H25-PAA. In two other interesting literature examples (Kyrili et al., 2017; Pippa et al., 2016), pH-responsive block copolymers were incorporated into liposomes. The copolymers consisted of the stimuli-responsive hydrophilic part of PAA and the nonresponsive poly(n-butylacrylate) (PnBA) hydrophobic part. Through the hydrophobic PnBA block, the copolymer can be anchored inside the liposomal membrane. Two different polymers with different content of PAA (70% and 85%) were used. This perturbation from the hydrophobic block can alter both the physicochemical and the calorimetric profiles of the liposomal system. Poly(n-butylacrylate)-b-poly (acrylic acid) (PnBA-bPAA) block copolymers were incorporated in both 1,2-dipalmitoyl-sn-glycero-3phosphocholine (DPPC) and HSPC bilayers that were subsequently evaluated by the DSC technique. According to the results, the specific enthalpy ΔH m of the main transition decreased significantly at all molar ratios with a concentration-dependent manner, while the characteristic temperatures, the Tonset,m, and the main transition temperature, Tm, did not change significantly indicating that the interactions of the PAA polymeric segment with DPPC or HSPC lipids affect the mobility of the polar head groups of the lipids. At both lipidic bilayers, DPPC and HSPC, the incorporation of the copolymer caused disappearance of the pretransition peak, suggesting that the interaction between PAA blocks and DPPC or HSPC lipids occurs on the membrane surface. The formation of hydrogen bonds between nonionized polymer carboxyl groups and the lipids polar head groups may be responsible for the above phenomenology. Taking into account the obtained ΔT1/2 values, the cooperativity of all the systems decreased significantly and especially at the highest molar ratio of the polymeric component, where fluidization phenomena of the bilayers may occur, due to the extreme broadening of the main curve. As authors suggest, this phenomenon at the highest polymeric molar ratio can also be correlated with the increased concentration of hydrophobic nanoclusters from the hydrophobic block of PnBA, creating highly concentrated PnBA regions, perturbating and eventually disrupting the lipidic bilayer. The presence of shoulders indicates that the PnBA-b-PAA is not uniformly distributed in the lipid bilayers, suggesting lateral phase separation into polymer-rich and polymer-poor domains. Concerning the observed differences due to the different PAA block content utilized, the increase of the PAA block (85%) in the copolymer caused the creation of a new phase and probably an interdigitated phase, reflected by the hysteresis being observed at the cooling curves. Block copolymers can also be synthesized from two different stimuli-responsive parts. For example, PNIPAM and PAA have been also combined in a dual-responsive diblock copolymer poly(N-isopropylacrylamide)-block-poly(acrylic acid) (PNIPAM-bPAA), dual responsive to pH and temperature changes. DPPC lipidic bilayers

Stimuli-responsive nanocarriers for drug delivery

incorporating the above copolymer were prepared and thermodynamically evaluated, regarding polymer composition, polymeric guest concentration, as well as the nature of the dispersion medium conditions (i.e., pH and ionic strength) (Kolman et al., 2016). Similar calorimetric shifts were also observed in HPLC-grade water (pH 5 5.5) and in PBS (with pH 5 7.4) dispersion medium, but significant differences were observed in acidic medium (hydrochloric acid solution of pH 4.5), due to the PAA block pH responsiveness. Another significant observation was the decrease of the cooperativity in higher temperatures of the LCST of the PNIPAM, probably due to the presence of the PNIPAM block. At higher temperatures, where the PNIPAM block becomes less hydrophilic (coil-to-globule transition), it penetrates the bilayers, forming hydrophobic clusters and finally causing membrane disruption (temperature-dependent perturbation). Thus both the stimuli-responsive blocks affect the calorimetric profile of the whole system. The authors concluded that the study of dual-stimuli-responsive chimeric bilayers can be used as a projection of the behavior of the respective dual-stimuli-responsive chimeric liposomes, being able to respond to both pathological tissue pH and temperature alterations.

4.4.2 Thermal analysis of thermoresponsive liposomes DSC has also been applied on liposomal platforms with incorporated thermoresponsive amphiphilic diblock copolymers. In this case, chimeric membranes are scanned multiple times by heating, to assess the effect of the polymers on them. Bilayers of DPPC and poly(N-isopropylacrylamide)-b-poly(lauryl acrylate) (PNIPAM-b-PLA) amphiphilic diblock copolymers were developed and studied (Naziris et al., 2019). Once more two different in composition copolymers were utilized, PNIPAM-b-PLA 1 of weight composition 66% 34% and PNIPAM-b-PLA 2 of weight composition 50% 50%. The conclusion made was that these chimeric nanosystems exhibit thermoresponsive behavior and must be heated above the newly developed phase transition, to reach a certain equilibrium state. However, the initial metastable and nonequilibrium phase is what renders them functional and might facilitate the controlled release of drug molecules. Pippa et al. (2015a) utilized the thermoresponsive amphiphilic homopolymer C12H25-poly(N-isopropylacrylamide)-COOH (C12H25-PNIPAM-COOH), whose hydrophilic segment is the PNIPAM and the lipophilic is the hydrocarbon end group C12H25. Two polymers (PNIPAM 1 and 2) of different molecular weights (different degree of polymerization of the PNIPAM part) were synthesized by the reversible addition-fragmentation chain transfer polymerization technique. Subsequently thermoresponsive chimeric liposomes were developed by incorporating the two different forms of the above homopolymer to DPPC liposomes. In the newly developed liposomal nanocarriers the hydrophobic drug IND was also incorporated. The C12H25 end group of the polymer was utilized for anchoring into the liposomal membrane,

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PNIPAM was stretching out of the liposomal surface, whereas the end-functional group of COOH made the system sensitive also to pH alterations. According to the DSC results, the incorporation of PNIPAM 1 (20 kDa) in DPPC liposomes affected only the specific enthalpy ΔHm of the main transition. Different shifts of ΔHm were observed as a function of the molar ratio of the polymeric guest into the liposomes. Moreover, the incorporation of highest polymer amounts caused the appearance of new peaks (centered at approximately 30 C), indicating the creation of new metastable phases, as a result of the heterogeneous distribution of the polymer that creates the formation of PNIPAM-1-rich and PNIPAM-1-poor domains. The main transition temperature Tm was slightly reduced, while the system cooperativity was reduced at a concentration-dependent manner (severe decrease at the highest polymer ratios, again due to the abovementioned nonuniform polymeric distribution). From the observed hysteresis on cooling results, authors also suggested the formation of an interdigitated phase at high concentration of the polymeric guest. In contrast to the PNIPAM 1 systems, no new metastable phases were observed for the PNIPAM 2 systems. Thus the polymer characteristics, such as hydrophilic/hydrophobic parts constitution/sequence, composition and molecular weight, can play a key role in the thermal behavior of the whole nanosystem. The calorimetric profile was subsequently correlated with the kinetic release of the incorporated drug IND. PNIPAM 1 exhibited larger incorporation efficiency, immediate, temperature dependant, “burst” release of IND at 37 C and no release at 32 C, due to the LCST of the PNIPAM. The authors suggested that the new metastable phases centered at 30 C may behave as promoters of the release of IND from the highest ratios of DPPC:PNIPAM 1 at 37 C, due to the inhomogeneous drug distribution inside the chimeric liposomal membrane. In contrast, DPPC:PNIPAM 2 systems exhibited absence of metastable phases, had lower incorporation efficiency, and did not release IND at all.

4.5 Physicochemical properties of stimuli-responsive liposomes Based on DSC analysis and results, chimeric liposomes are developed and studied through various techniques. One important aspect of liposomes in general is their physicochemical properties, that is, their particle size.

4.5.1 Physicochemical characterization of pH-responsive liposomes The physicochemical properties of liposomes with pH-responsive amphiphilic diblock copolymers have been studied using dynamic light scattering (DLS).

Stimuli-responsive nanocarriers for drug delivery

This allows for the determination of the effect of the pH-responsive behavior of the polymers on the size (hydrodynamic diameter, Dh) of the nanosystem. Chimeric liposomes of the phospholipid HSPC and the pH-responsive copolymer PnBA-b-PAA were developed through the thin-film hydration method and analyzed through DLS (Kyrili et al., 2017). Concerning the copolymer, two different types were utilized, PnBA-b-PAA 1 of weight composition 15% 85% and PnBAb-PAA 2 of weight composition 30% 70%. Exposure of these chimeric nanosystems to acidic environment led to alterations of their particle size, depending on the type of copolymer but also on the copolymer concentration. In particular, for all systems with PnBA-b-PAA 1, the hydrodynamic diameter was decreased at acidic pH. On the other hand, PnBA-b-PAA 2 led to a tremendous increase of size at phospholipid:polymer 9:0.1 molar ratio, significant increase at 9:0.5 and decrease at the rest of the molar ratios. Authors attributed the above size decreases to the pH-dependent properties of the PAA block (pKaB4.2). At neutral pH (PBS), the PAA block is partially ionized with expanded chains, while at citrate buffer conditions (pH 5 4.0), the acidic groups ( COO 2) of PAA block are protonated to COOH and as a result PAA loses its expanded form and the polymer chains shrink, yielding a greater binding of PAA block to the membrane.

4.5.2 Physicochemical characterization of thermoresponsive liposomes The physicochemical properties of liposomes with incorporated thermoresponsive amphiphilic diblock copolymers have been studied by DLS, to evaluate the thermoresponsive behavior of such chimeric nanosystems and its effect on liposomal size. Thermoresponsive liposomes of DPPC and PNIPAM-b-PLA of weight composition 66% 34% were developed through thin-film hydration method and analyzed through DLS experiments (Naziris et al., 2019). The particle size of the thermoresponsive chimeric liposomes was measured at room temperature and after heating at 45 C, and the results are presented in Fig. 4.2. The thermoresponsiveness of DPPC:PNIPAM-b-PLA was confirmed through DLS, as the physicochemical properties of the chimeric nanoparticles were altered substantially after heating. In specifics the response was profound, the suspensions went from homogeneous to biphasic, and large aggregates were formed that were visible with the naked eye. The degree of the phenomenon was dependent on the concentration of the incorporated polymer. After the samples reached room temperature again, the aggregates were no more visible, but the DLS measurement revealed the size of nanoparticles, which was increased in a polymer concentration dependent manner.

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Figure 4.2 Particle size of chimeric nanosystems DPPC:PNIPAM-b-PLA at room temperature and after heating at 45 C.

4.6 Development of stimuli-responsive lyotropic liquid crystalline nanosystems Nonlamellar, lyotropic liquid crystalline nanoparticles, such as cubosomes and hexosomes, are usually prepared from the lipids glyceryl monooleate (GMO, 2,3dihydroxypropyl oleate) and phytantriol (PHYT, 3,7,11,15-tetramethyl-1,2,3-hexadecanetriol). These lipids can self-assemble into different mesophases as a function of concentration and temperature. The hexagonal (HII) and bicontinuous (QII) cubic phases, in the presence of amphiphilic block copolymers utilized as stabilizers, are further exploited toward the development of stable colloidal dispersions of nanoparticles, hexosomes, and cubosomes, respectively. The stabilization by the Pluronics triblock copolymers, usually by F-127 (Poloxamer P407), is the most well investigated (Azmi et al., 2015; Mulet et al., 2013; Chountoulesi et al., 2018). Nonlamellar liquid crystalline nanoparticles proved to be ideal therapeutic nanosystems, exhibiting tunable structural characteristics, high grade of internal organization, allowing the loading with different kinds of agents, such as amphiphilic, hydrophobic, and hydrophilic agents. Moreover they are able to carry greater volumes of cargo than liposomes and allow the sustained release of their content (Azmi et al., 2015; Lancelot et al., 2014). The development of stimuli-responsive liquid crystalline nanosystems is considered to be an attractive strategy, to further improve their ability of controlled content release and facilitate on demand drug delivery. Similarly to the liposomes, stimuli-responsive

Stimuli-responsive nanocarriers for drug delivery

release from the nonlamellar liquid crystalline nanosystems can be achieved by intrinsic (physiological pH or/and temperature differences between normal and diseased tissues) and external stimuli, such as temperature, light, and magnetic fields (Fong et al., 2016). In the present report, literature examples about stimuli-responsive liquid crystalline nanosystems are presented, that respond to either pH or temperature and employ either stimuli-responsive block copolymers or other types of functional biomaterials. Moreover we also emphasize at the techniques that were utilized, to comprehensively evaluate the properties of the presented stimuli-responsive liquid crystalline nanosystems.

4.6.1 Stimuli-responsive lyotropic liquid crystalline nanosystems using polycation of PDMAEMA Chountoulesi et al. (2019) developed for the first time liquid crystalline nanoparticles with incorporated polycations of PDMAEMA, being suitable for novel drug and gene delivery nanosystems. In this study, the stimuli-responsive amphiphilic block copolymer PDMAEMA-b-PLMA is proposed as a novel stabilizer for liquid crystalline nanoparticles with advanced stimuli-responsive properties, exhibiting double role, both stabilizing and responding to pH and temperature alterations. During the formulation stage, two different lipids were investigated, both GMO and PHYT, as well as different concentrations of the PDMAEMA-b-PLMA stabilizers. All these formulation parameters were found to affect both the physicochemical behavior and the morphology of the nanosystems. Regarding the physicochemical behavior, being investigated by a gamut of light scattering techniques (dynamic, electrophoretic, and static), the GMO nanosystems were proven to be colloidally stable over time, retaining their physicochemical characteristics (size and size distribution). On the other hand, the PHYT systems were less stable over time, indicating different lipid stabilizer interactions between GMO and PHYT cases. High positive values of ζ-potential were observed, probably due to charged amino groups of PDMAEMA in their structure. Regarding their pH responsiveness, all nanosystems presented pH-induced alterations of their charge in the three studied environments/media with different pH values (pH 5 4.2, 6.0, and 7.4) (Fig. 4.3). More specifically their positive charge was increased in more acidic environment and was decreased at neutral pH. This observed pH-responsive charge conversion is considered to be a very useful property, which can facilitate endosomal escape, uptake by highly negative charged membranes and complexation with nucleic acids for both gene and drug delivery applications. Their size and size distribution were not significantly changed upon pH alteration. Apart from pH, also the response to temperature was monitored at three different temperatures (25 C, 37 C, and 55 C), taking into account the temperature responsiveness of the PDMAEMA block. Irreversible physicochemical and morphological alterations were observed, probably due to the temperature-induced shrinkage of

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Figure 4.3 Hydrodynamic radius (Rh, nm) of (A) GMO and (C) PHYT nanosystems, as well as ζ-potential (mV) of (B) GMO and (D) PHYT nanosystems depending on the pH of the dilution medium (K: lipid:PDMAEMA-b-PLMA 9:1, ’: lipid:PDMAEMA-b-PLMA 9:3, ▲: lipid:PDMAEMA-bPLMA:P407 8:1:1). Adapted from Chountoulesi, M., Pippa, N., Chrysostomou, V., Pispas, S., Chrysina, E. D., Forys, A., et al., 2019. Stimuli-responsive lyotropic liquid crystalline nanosystems with incorporated poly(2-dimethylamino ethyl methacrylate)-b-poly(lauryl methacrylate) amphiphilic block copolymer. Polymers (Basel) 11 (9), 1400.

PDMAEMA by its conversion from hydrophilic to more hydrophobic at increased temperatures. The effects from other environmental parameters, such as ionic strength and the presence of serum proteins, were also investigated, revealing aggregation phenomena upon ionic strength increase and some disintegration phenomena due to the presence of proteins. Alterations of their microenvironmental parameters (micropolarity and microfluidity) were observed by fluorescence spectroscopy. Concerning their morphology, cryogenic-transmission electron microscopy (cryo-TEM) studies revealed increased morphological diversity, presenting nanoparticles with different grades of internal organization, strictly dependant on the concentration of the polymeric stabilizer.

4.6.2 pH-responsive cubosomes by using pH-sensitive polymer As an effort to overcome the lack of sensitivity of monoolein cubosomes to pH conditions, Kluzek et al. (2017) used a pH-sensitive polymer designed to strongly interact

Stimuli-responsive nanocarriers for drug delivery

with the lipid structure at low pH, promoting a triggered drug release from the cubic phase into acidic environments. More specifically, they used the pseudopeptidic polymer poly(L-lysine-iso-phthalamide) grafted with L-phenylalanine with 50% grafting degree, designed to mimic the membrane-penetrating peptides. This polymer exhibits carboxylic groups that can promote conformational changes. At neutral pH, there are extended charged polymeric chains, which are transformed to globular state at acidic conditions, promoting binding with the lipidic membrane and eventually causing its disruption. By using cryo-TEM and small-angle X-ray scattering (SAXS) techniques, the authors proved that under the environmental pH alteration from neutral to acidic, a significant amount of structural disruption, with a partial disruption of the cubic phase (totally or partially disordered particles, coexisting with cubic and lamellar ordered particles), is taking place. This property would be very useful toward the development of cubosomal drug delivery nanosystems that can facilitate triggered drug release, for example, in the intracellular acidic conditions.

4.6.3 pH-responsive liquid crystalline nanosystems using pH-responsive molecules Apart from block copolymers, other pH-sensitive molecules, which are able to exhibit protonation/deprotonation, can be incorporated into the liquid crystalline phases promoting phase transitions upon pH alterations. Negrini and Mezzenga (2011) developed a food-grade lyotropic liquid crystal system, capable of responding to pH variations, switching in a reversible way both its structure and physical properties, resulting in pHcontrolled release from a monolinolein cubic phase. The chosen pH-sensitive triggering molecule was linoleic acid. The linoleic acid, as a weak fatty acid, is protonated at acidic pH. This protonation of the incorporated molecule induced structural changes to the liquid crystalline system from the reverse bicontinuous cubic phase at neutral pH (simulating intestine conditions) to reverse columnar hexagonal phase in acidic environment (pH 5 2, simulating stomach conditions), as revealed by SAXS studies. The proposed system was able to retain the drug in acidic conditions typical of the human stomach and release drug in neutral pH conditions, being suitable for oral administration of drugs for targeted delivery in intestine or colon tracts. Negrini et al. (2015) proposed also a liquid crystalline pH-responsive system that exhibited a phase transition from hexagonal phase at higher pH to cubic phase at lower pH, induced by the protonation of the incorporated weak amphiphilic base pyridinylmethyl linoleate at acid pH. The pH-induced hexagonal-to-cubic transition at pH , 5.5 was detected by SAXS measurements. This nanosystem was exploited to promote controlled drug release from the mesophases in the low pH microenvironment of tumors and more specifically of the anticancer drug doxorubicin. According to the results, the system was shown to increase human colon cancer cell death at an acidic pH mimicking tumor conditions. More specifically the in vitro release studies of

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doxorubicin on HT29 human colon cancer cells showed 10-fold faster release and 3-fold increased efficiency in killing cancer cells at pH 5 5.5 versus pH 5 7.4.

4.6.4 Thermoresponsive lipid-based liquid crystalline nanosystems Liquid crystalline nanosystems can be affected by the temperature increase, which can induce transitions of their phases. GMO is proved to perform typical liquid crystal thermal expansivity, while the lattice parameter of all GMO mesophases decreases with increasing temperature (Briggs et al., 1996; Qiu and Caffrey, 2000). In another typical example, emulsified liquid crystalline particles from monolinolein lipid (MLO) were proved to own cubic internal structure at 25 C, that is transformed to inverse hexagonal after heating at 58 C and to inverse micellar phase upon further heating to 87 C. The described phase transitions were reversible upon cooling process with opposite sequence of steps. Apart from this temperature-induced morphological transition, MLO-based particles expel water upon heating (deswelling/shrinkage) and take up water again upon cooling (swelling) also in a reversible way, termed as “breathing mode” (de Campo et al., 2004). Taking into account the temperature-induced transitions of the lipids and applying functional modifying molecules, we can control the temperature responsiveness of the liquid crystalline nanosystems. For example, Barriga et al. (2015) prepared ternary mixtures, combining monoolein, cholesterol, and the negatively charged phospholipids 1,2-dioleoyl-sn-glycero-3-phospho(10-rac-glycerol) and 1,2-dioleoyl-sn-glycero-3phospho-L-serine. The above combinations resulted in highly swollen primitive (Im3m) symmetry bicontinuous cubic phases, with lattice parameters of up to 480  A, highly sensitive to both temperature and pressure. Fong et al. (2009) developed thermoresponsive lipid-based liquid crystalline system, being the first in situ triggered on demand drug delivery system based on a lipid mesophases. Different ratios of PHYT-vitamin E acetate and GMO-oleic acid were investigated. The temperature-induced changes of the nanostructure between reverse hexagonal and bicontinuous cubic were confirmed by crossed polarized optical microscopy and SAXS. By modifying the temperature, the precise control of the release rate of a hydrophilic model drug, glucose, was achieved in vitro. During in vivo experiments in rats (subcutaneously administration of the systems), a heat or cool pack at the injection site was applied to evaluate their potential thermoresponsive ability and stimulated changes in drug release from the matrix were observed.

4.7 Conclusion and future directions There is a great progress in the research field of the stimuli-responsive drug delivery nanosystems. Taking into account their unique smartness of responding to the

Stimuli-responsive nanocarriers for drug delivery

altered conditions of the pathological tissues, stimuli-responsive drug delivery nanosystems have gained much interest and many different categories have already been developed. In this chapter, literature case studies of stimuli-responsive liposomes and nonlamellar lyotropic liquid crystalline nanosystems were presented, prepared from both lipids and stimuli-responsive polymers. The combination of a great variety of different in nature biomaterials, such as lipids and stimuli-responsive polymers, can provide versatility in design and as a result many different formulations of chimeric/mixed aDDnSs with upgraded functionality that can be applied in the area of stimuliresponsive drug delivery nanosystems. However, further development of stimuliresponsive therapeutic nanosystems remains still a challenge, to achieve efficient clinical results and quick translation into clinical applications, apart from the already achieved progress in the preclinical stages. Toward this scope, there is an emerging need for continuous development of smart innovative biomaterials (i.e., polymers) that will upgrade the already existing systems of this promising category of nanotherapeutics. Furthermore, robust techniques are necessary to evaluate the characteristics of the potential newly developed stimuli-responsive nanosystems and deeply understand the interactions taking place between their different biomaterial components.

References Al-Ahmady, Z.S., Al-Jamal, W.T., Bossche, J.V., Bui, T.T., Drake, A.F., Mason, A.G., et al., 2012. Lipid-peptide vesicle nanoscale hybrids for triggered drug release by mild hyperthermia in vitro and in vivo. ACS Nano. 6 (10), 9335 9346. Alvarez-Lorenzo, C., Concheiro, A., 2014. Smart drug delivery systems: from fundamentals to the clinic. Chem. Commun. (Camb.) 50 (58), 7743 7765. Azmi, I.D.M., Moghimi, S.M., Yaghmur, A., 2015. Cubosomes and hexosomes as versatile platforms for drug delivery. Ther. Deliv. 6 (12), 1347 1364. Barriga, H.M., Tyler, A.I., McCarthy, N.L., Parsons, E.S., Ces, O., Law, R.V., et al., 2015. Temperature and pressure tuneable swollen bicontinuous cubic phases approaching nature’s length scales. Soft Matter. 11, 600 607. Bhattacharya, B., Mohd Omar, M.F., Soong, R., 2016. The Warburg effect and drug resistance. Br. J. Pharmacol. 173 (6), 970 979. Briggs, J., Chung, H., Caffrey, M., 1996. The temperature-composition phase diagram and mesophase structure characterization of the monoolein/water system. J. Phys. II 6, 723 751. Bron, R., Ortiz, A., Dijkstra, J., Stegmann, T., Wilschut, J., 1993. Preparation, properties, and applications of reconstituted influenza virus envelopes (virosomes). Meth. Enzymol. 220, 313 331. de Campo, L., Yaghmur, A., Sagalowicz, L., Leser, M.E., Watzke, H., Glatter, O., 2004. Reversible phase transitions in emulsified nanostructured lipid systems. Langmuir 20, 5254 5261. Chen, K.J., Liang, H.F., Chen, H.L., Wang, Y., Cheng, P.Y., Liu, H.L., et al., 2013. A thermoresponsive bubble-generating liposomal system for triggering localized extracellular drug delivery. ACS Nano. 7 (1), 438 446. Chountoulesi, M., Kyrili, A., Pippa, N., Meristoudi, A., Pispas, S., Demetzos, C., 2017. The modulation of physicochemical characterization of innovative liposomal platforms: the role of the grafted thermoresponsive polymers. Pharm. Dev. Technol. 22 (3), 330 335. Chountoulesi, M., Pippa, N., Pispas, S., Chrysina, E.D., Forys, A., Trzebicka, B., et al., 2018. Cubic lyotropic liquid crystals as drug delivery carriers: physicochemical and morphological studies. Int. J. Pharm. 550 (1-2), 57 70.

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Jhaveri, A., Deshpande, P., Torchilin, V., 2014. Stimuli-sensitive nanopreparations for combination cancer therapy. J. Control. Release 190, 352 370. John, J.V., Johnson, R.P., Heo, M.S., Moon, B.K., Byeon, S.J., Kim, I., 2015. Polymer-blockpolypeptides and polymer-conjugated hybrid materials as stimuli-responsive nanocarriers for biomedical applications. J. Biomed. Nanotechnol. 11 (1), 1 39. Kanamala, M., Wilson, W.R., Yang, M., Palmer, B.D., Wu, Z., 2016. Mechanisms and biomaterials in pH-responsive tumour targeted drug delivery: a review. Biomaterials 85, 152 167. Karimi, M., Ghasemi, A., Sahandi Zangabad, P., Rahighi, R., Moosavi Basri, S.M., Mirshekari, H., et al., 2016. Smart micro/nanoparticles in stimulus-responsive drug/gene delivery systems. Chem. Soc. Rev. 45 (5), 1457 1501. Kim, K., Chen, W.C.W., Heo, Y., Wang, Y., 2016. Polycations and their biomedical applications. Prog. Polym. Sci. 60, 18 50. Kluzek, M., Tyler, A.I.I., Wang, S., Chen, R., Marques, C., Thalmann, F., et al., 2017. Influence of a pH-sensitive polymer on the structure of monoolein cubosomes. Soft Matter 13, 7571 7577. Kneidl, B., Peller, M., Winter, G., Lindner, L.H., Hossann, M., 2014. Thermosensitive liposomal drug delivery systems: state of the art review. Int. J. Nanomed. 9, 4387 4398. Kolman, I., Pippa, N., Meristoudi, A., Pispas, S., Demetzos, C., 2016. A dual-stimuli-responsive polymer into phospholipid membranes. J. Therm. Anal. Calorim. 123, 2257 2271. Kono, K., 2001. Thermosensitive polymer-modified liposomes. Adv. Drug Deliv. Rev. 53 (3), 307 319. Kono, K., Zenitani, K., Takagishi, T., 1994a. Novel pH-sensitive liposomes:liposomes bearing a poly (ethylene glycol) derivative with carboxyl groups. Biochim. Biophys. Acta 1193, 1 9. Kono, K., Hayashi, H., Takagishi, T., 1994b. Temperature-sensitive liposomes: liposomes bearing poly (N-isopropylacrylamide). J. Control. Release 30 (1), 69 75. Kyrili, A., Chountoulesi, M., Pippa, N., Meristoudi, A., Pispas, S., Demetzos, C., 2017. Design and development of pH-sensitive liposomes by evaluating the thermotropic behavior of their chimeric bilayers. J. Therm. Anal. Calorim. 127 (2), 1381 1392. Lancelot, A., Sierra, T., Serrano, J.L., 2014. Nanostructured liquid-crystalline particles for drug delivery. Expert Opin. Drug. Deliv. 11 (4), 547 564. Lanzalaco, S., Armelin, E., 2017. Poly(n-isopropylacrylamide) and copolymers: a review on recent progresses in biomedical applications. Gels 3, 36 67. Lee, S.H., Choi, S.H., Kim, S.H., Park, T.G., 2008. Thermally sensitive cationic polymer nanocapsules for specific cytosolic delivery and efficient gene silencing of siRNA: swelling induced physical disruption of endosome by cold shock. J. Control. Release 125 (1), 25 32. Lee, S.M., Nguyen, S.T., 2013. Smart nanoscale drug delivery platforms from stimuli-responsive polymers and liposomes. Macromolecules 46, 9169 9180. Lee, Y., Thompson, D.H., 2017. Stimuli-responsive liposomes for drug delivery. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 9. Available from: https://doi.org/10.1002/wnan.1450. Lee, Y., Park, S.Y., Kim, C., Park, T.G., 2009. Thermally triggered intracellular explosion of volume transition nanogels for necrotic cell death. J. Control. Release 135 (1), 89 95. Leroux, J.C., Roux, E., Garrec, D.L., Hong, K., Drummond, D.C., 2001. N-Isopropylacrylamide copolymers for the preparation of pH-sensitive liposomes and polymeric micelles. J. Control. Release 72, 71 84. Li, R., Xie, Y., 2017. Nanodrug delivery systems for targeting the endogenous tumor microenvironment and simultaneously overcoming multidrug resistance properties. J. Control. Release 251, 49 67. Li, Z., Fan, Z., Xu, Y., Lo, W., Wang, X., Niu, H., et al., 2016. pH and thermal sensitive hydrogels as stem cell carriers for cardiac therapy. ACS Appl. Mater. Interfaces 8 (17), 10752 10760. Lin, Y.L., Jiang, G., Birrell, L.K., El-Sayed, M.E.H., 2010. Degradable, pH-sensitive, membranedestabilizing, comb-like polymers for intracellular delivery of nucleic acids. Biomaterials 31, 7150 7166. Liu, D., Huang, L., 1990. pH-sensitive, plasma-stable liposomes with relatively prolonged residence in circulation. Biochim. Biophys. Acta 1022, 348 354. Liu, R., Fraylich, M., Saunders, B.R., 2009. Thermoresponsive copolymers: from fundamental studies to applications. Colloid. Polym. Sci. 287, 627 643.

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Liu, J., Huang, Y., Kumar, A., Tan, A., Jin, S., Mozhi, A., et al., 2014. pH-Sensitive nano-systems for drug delivery in cancer therapy. Biotechnol. Adv. 32 (4), 693 710. Liu, M., Du, H., Zhang, W., Zhai, G., 2017a. Internal stimuli-responsive nanocarriers for drug delivery: design strategies and applications. Mater. Sci. Eng. C Mater. Biol. Appl. 71, 1267 1280. Liu, P., Song, L., Li, N., Lin, J., Huang, D., 2017b. Time dependence of phase separation enthalpy recovery behavior in aqueous poly(N-isopropylacrylamide) solution. J. Therm. Anal. Calorim. 130, 843 850. Liu, X., Huang, G., 2013. Formation strategies, mechanism of intracellular delivery and potential clinical applications of pH-sensitive liposomes. Asian J. Pharm. Sci. 8 (6), 319 328. Mignet, N., Richard, C., Seguin, J., Largeau, C., Bessodes, M., Scherman, D., 2008. Anionic pHsensitive pegylated lipoplexes to deliver DNA to tumors. Int. J. Pharm. 361, 194 201. Mulet, X., Boyd, B.J., Drummond, C.J., 2013. Advances in drug delivery and medical imaging using colloidal lyotropic liquid crystalline dispersions. J. Colloid. Interface Sci. 393, 1 20. Mura, S., Nicolas, J., Couvreur, P., 2013. Stimuli-responsive nanocarriers for drug delivery. Nat. Mater. 12 (11), 991 1003. Naziris, N., Pippa, N., Pispas, S., Demetzos, C., 2016. Stimuli-responsive drug delivery nanosystems: from bench to clinic. Curr. Nanomed. 6, 1 20. Naziris, N., Pippa, N., Meristoudi, A., Pispas, S., Demetzos, C., 2017. Design and development of pHresponsive HSPC:C12H25-PAA chimeric liposomes. J. Liposome Res. 27 (2), 108 117. Naziris, N., Pippa, N., Stellas, D., Chrysostomou, V., Pispas, S., Demetzos, C., et al., 2018. Development and evaluation of stimuli-responsive chimeric nanostructures. AAPS Pharm. Sci. Tech. 19, 2971 2989. Naziris, N., Skandalis, A., Forys, A., Trzebicka, B., Pispas, S., Demetzos, C., 2019. A thermal analysis and physicochemical study on thermoresponsive chimeric liposomal nanosystems. J. Thermal. Anal. Calorim. accepted manuscript. Negrini, R., Mezzenga, R., 2011. pH-responsive lyotropic liquid crystals for controlled drug delivery. Langmuir 27 (9), 5296 5303. Negrini, R., Fong, W.K., Boyd, B.J., Mezzenga, R., 2015. pH-responsive lyotropic liquid crystals and their potential therapeutic role in cancer treatment. Chem. Commun. 51, 6671 6674. Pelton, R., 2010. Poly(N-isopropylacrylamide) (PNIPAM) is never hydrophobic. J. Colloid. Interface Sci. 348, 673 674. Pippa, N., Meristoudi, A., Pispas, S., Demetzos, C., 2015a. Temperature-dependent drug release from DPPC:C12H25-PNIPAM-COOH liposomes: control of the drug loading/release by modulation of the nanocarriers’ components. Int. J. Pharm. 485 (1-2), 374 382. Pippa, N., Pispas, S., Demetzos, C., 2015b. The metastable phases as modulators of biophysical behavior of liposomal membranes. J. Therm. Anal. Calorim. 120, 937 945. Pippa, N., Chountoulesi, M., Kyrili, A., Meristoudi, A., Pispas, S., Demetzos, C., 2016. Calorimetric study on pH-responsive block copolymer grafted lipid bilayers: rational design and development of liposomes. J. Liposome Res. 26 (3), 211 220. Qiao, Y., Wan, J., Zhou, L., Ma, W., Yang, Y., Luo, W., et al., 2019. Stimuli-responsive nanotherapeutics for precision drug delivery and cancer therapy. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 11 (1), 1527. Qiu, H., Caffrey, M., 2000. The phase diagram of the monoolein/water system: metastability and equilibrium aspects. Biomaterials 21, 223 234. Roux, E., Francis, M., Winnik, F.M., Leroux, J.C., 2002. Polymer based pH-sensitive carriers as a means to improve the cytoplasmic delivery of drugs. Int. J. Pharm. 242, 25 36. Sánchez-Moreno, P., de Vicente, J., Nardecchia, S., Marchal, J.A., Boulaiz, H., 2018. Thermo-sensitive nanomaterials: recent advance in synthesis and biomedical applications, Nanomaterials (Basel), 8. pp. 935 966. Southall, N.T., Dill, K.A., Haymet, A.D.J., 2002. A view of the hydrophobic effect. J. Phys. Chem. B. 106, 521 533. Stubbs, M., McSheehy, P.M.J., Griffiths, J.R., Bashford, C.L., 2000. Causes and consequences of tumour acidity and implications for treatment. Mol. Med. Today 6 (1), 15 19.

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Tian, L., Bae, Y.H., 2012. Cancer nanomedicines targeting tumor extracellular pH. Colloids Surf. B Biointerfaces 99, 116 126. Viard, M., Puri, A., 2015. Stimuli-sensitive liposomes: lipids as gateways for cargo release. In: Igliˇc, A., Kulkarni, C.V., Rappolt, M. (Eds.), Advances in Planar Lipid Bilayers and Liposomes. Academic Press, Cambridge, MA, pp. 1 41. Ward, M.A., Georgiou, T.K., 2011. Thermoresponsive polymers for biomedical applications. Polymers 3, 1215 1242. Xu, H., Paxton, J.W., Wu, Z., 2015. Enhanced pH-responsiveness, cellular trafficking, cytotoxicity and long-circulation of PEGylated liposomes with post-insertion technique using gemcitabine as a model drug. Pharm. Res. 32 (7), 2428 2438.

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CHAPTER FIVE

Biodegradable nanomaterials Katerina Anagnostou1, Minas Stylianakis1, Sotiris Michaleas2 and Athanasios Skouras1,2 1

Department of Electrical & Computer Engineering, Hellenic Mediterranean University Heraklion, Crete, Greece Department of Life Sciences, School of Sciences, European University Cyprus, Nicosia, Cyprus

2

5.1 Introduction Starting from 1970 the interdisciplinary approach between chemical, biological, and pharmaceutical sciences has led to the introduction of new applications in the medical field thanks to greater versatility in controlling the physical state, shape, size, and surface of polymers. This new knowledge has led to considerable innovation in the design and development of drug delivery systems (DDS): in those years began the study of reabsorbable controlled release systems of drugs, formed from biodegradable polymers (BPs). A thorough examination of the structure and properties of this type of materials led to the development of methods for their manufacture and the consequent production of materials—synthetic and not—, specific by type of application (Pillai and Panchagula, 2001). A polymer is a substance composed of repeating units, called monomers, held together by covalent bonds and its macrostructure may present itself as a linear, branched, or reticulated chain. It generally has a high-molecular weight, hence the name of macromolecule. Depending on the composition and arrangement of the monomers, they can be characterized as homopolymers or copolymers: the former consist in repeating the same unit for the entire length of the chain, the latter may have a sequence of two or more types of monomer in a more or less disordered manner. A polymeric material owes its physicochemical and mechanical properties to the composition, structure and molecular weight of the polymer chains of which it is composed. The polymers can be of natural origin (such as polysaccharides, nucleic acids, or proteins) or synthetic [such as polylactic acid (PLA) and polyurethane]. BPs exact discovery time cannot be easily traced, although one of their first medical application can be traced back to at least CE 100 (Nutton, 2013) as catgut suture made from sheep intestines. In regards with synthetic BPs, the first reports surface in the 1980s

Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00005-9

© 2020 Elsevier Inc. All rights reserved.

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Figure 5.1 List of biodegradable polymers showing their medicinal properties.

(Vroman, 2009). Applications of BPs have been utilized in many applications in clinical medicine, as depicted in Fig. 5.1, including tissue engineering and drug delivery. As the name implies, the main characteristic of BPs is their degradation by biological means. Degradation is defined as an irreversible process, which leads to a change of the structure of the material, in the form of loss of mechanical properties, damage, fragmentation, or depolymerization. Degradation is affected from the environment and can present a constant or variable speed over time. The human body uses mainly two processes to decompose large molecules dimensions: hydrolysis and enzymatic degradation. The synthetic polymers used in medicine are mainly decomposed by hydrolysis. The speed of this reaction and, consequently, the residence time of the polymer in the organism, are regulated by several factors: 1. Interaction of the polymer with water: the diffusion coefficient is important and, above all, water absorption in the polymer. The greater the hydrophilicity of the material, the greater the decomposition speed. 2. Crystallinity of the polymer: the amorphous regions of the material are less compact and allow for better penetration of water. The more crystalline a material is, the more it will resist degradation.

Biodegradable nanomaterials

3. Temperature has two effects that accelerates the reaction kinetics and makes more movable chains (if the glass transition temperature is exceeded). 4. Polymer structure: the presence of heteroatoms, hydrophilic groups and, in particular way, ethereal, ester, and urethane bonds facilitates degradation. Other factors such as water concentration, pH, and salt concentration can also influence the hydrolysis speed. The hydrolysis products of biocompatible polymers are often monomers that are disposed of thanks to normal metabolic body processes. For example, poly(lactic-co-glycolic acid) (PLGA) is decomposed into lactic acid and glycolic acid: the former is disposed of, similarly to that produced during intense physical activity in the muscles, through the kidneys or transformed into pyruvate, which enters the cycle of Krebs; the glycolic acid instead is used by the body for the synthesis of glycine and other amino acids. Practically a polymer is defined as “degradable” if it is broken down during the application or immediately thereafter; on the contrary, a material “not degradable” requires a much longer time for degradation than the duration of its application (Gopferich, 1996). BPs offer numerous advantages: • Chemically they are often inert in biological environments and exhibit less toxicity than nonbiodegradable materials. • They degrade in a controlled manner within the body thus allowing it to be easily absorbed or eliminated, avoiding the use of surgery for their removal and thus decreasing the discomfort for the patient. • They can be designed to allow adequate control over the kinetics of drug release. • They are naturally recycled by biological processes, as they are most often derived by plant processing of atmospheric carbon dioxide. BPs, as already stated, can be divided into two large categories: natural and synthetic polymers. The main polymers utilized so far in drug delivery applications are described below.

5.2 Natural polymers Biodegradable natural polymers or biopolymers derive from renewable resources, formed in nature during biological processes of different organisms. Chemical modification of these polymers is often required for improvement of their physicochemical and mechanical properties. Biopolymers can be further divided in polysaccharides, proteins, and bacterial polymers as depicted in Table 5.1.

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Table 5.1 Main biopolymers utilized as drug delivery systems (DDS) and their origin. Biopolymer Main polymers utilized Chemical structure Origin class

Polysaccharides

Proteins

Microbial origin

Starch

Seed, roots, tubers of potatoes, rice, etc.

Cellulose and derivatives

Cell wall of green plants

Chitosan

Crab shells

Hyaluronic acid

Extracellular matrix of tissues

Alginate

Brown algae, bacteria

Collagen

Connective tissue

Gelatin

Hydrolysis of collagen

Albumin

Serum protein

Polyhydroxyalcanoates

Bacterial

Poly-γ-glutamic acid

5.2.1 Polysaccharides Polysaccharides are polymeric carbohydrate molecules composed of monosaccharides bound together by glycosidic bonds. Their main function in living organisms is either structural or storage related. Polysaccharides include a large variety of biopolymers with different physicochemical properties, resulting mainly by the many different monosaccharides that compose them. Nevertheless, all of these biopolymers are biodegradable,

Biodegradable nanomaterials

biocompatible, widely available, and can be readily modified due to the presence of functional groups along their polymeric chain (Liu et al., 2008). Polysaccharides include cellulose, starch, alginic acid, hyaluronic acid, chitin, and chitosan.

5.2.1.1 Chitosan Chitosan is a linear polysaccharide consisting of repeating units of N-acetyl-D-glucosamine and D-glucosamine joined by (1-4)-β-glycosidic bonds. Generally, it is produced by alkaline deacetylation of chitin, widely distributed in nature as the main component of the exoskeleton of insects and crustaceans and of the cell wall of some fungi (Younes and Rinaudo, 2015). Unlike chitin, it is soluble in water under acidic conditions (pH less than 6) per effect of protonation of amine groups (pKa equal to 6.3), which gives the polymer a positive net charge. In addition to being biocompatible, nontoxic, and biodegradable, chitosan possesses a series of interesting biological properties that make it a good candidate for carrier of drugs. Thanks to its mucoadhesive properties, it prolongs the residence time of drugs at the site of absorption leading to a significant increase in their bioavailability (Ahmed and Aljaeid, 2016). Moreover, it exerts a strong antimicrobial action against both Gram-positive and Gram-negative bacteria. It also possesses immunoadjuvant, hemostatic, and cicatrizing properties, which justifies its use in tissue engineering and medicine applications regenerative.

5.2.1.2 Hyaluronic acid The very structure of hyaluronic acid, a polysaccharide composed of disaccharide units of D-glucuronic acid and N-acetyl-glucosamine connected by glycosidic bonds of type β (1,3) and β (1,4), in addition to its high biocompatibility, biodegradability, and nonimmunogenicity, gives it great potential as a DDS (Huang and Huang, 2018). This natural polymer is very soluble in water and thus the preparations obtained from it are highly unstable in physiological liquids and therefore not suitable for controlled drug release. This problem can be solved by derivatization with hydrophobic groups, which leads to the obtainment of temporarily insoluble polymers. An adequate choice of the substituent allows to obtain products without toxicity or immunogenicity and with good biocompatibility. The hyaluronic acid molecule carries a set of free hydroxyl and carboxylic groups, which can be used to directly bind the drug to the polysaccharide. In this way, during the physical process of the erosion and solubilization of the polymer matrix, a significant lowering of the drug release rate is obtained in the surrounding environment, due in large part to the time necessary for the chemical breakdown of these bonds (Khunmanee et al., 2017)

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5.2.1.3 Alginate Alginates are a family of linear polysaccharides consisting of two epimers: α-L guluronic acid (G) and β-D mannuronic acid (M). Its structure is formed from homopolymer sequences (MM or GG) between which sequences are found heteropolymeric (MG) (Lee and Mooney, 2012). The monomeric composition of alginate, the molecular weight, and the extension of the guluronic sequences emannuronics influence its properties (George and Abraham, 2006). The alginate is an important component present in brown algae as well as an exopolysaccharide of different bacteria. Since alginate is prepared in aqueous media offers an attractive alternative for the formulation of compounds that are unstable in organic solvents such as proteins. However, since cation-induced gel formation is reversible, a disadvantage of unmodified, alginate-based delivery systems is rapid drug release when exposed to monovalent ions in physiological media. 5.2.1.4 Starch Starch is a low-cost, abundantly available, environmentally friendly polysaccharide mainly extracted from potatoes, rice, and corn. It is a copolymer of amylopectin and amylose composed of D-glucopyranoside monomers linked with different glycosidic bonds. Starch, however, presents poor mechanical properties and a certain water sensibility. These problems can be addressed either by modifying the hydroxyl functionality as is the case with acetylated starch or by blending starch with other BPs. Starch acetate has a higher content of the linear amylose and is thus more hydrophobic and has better film-forming capability compared to native starch (Heinze and Koschella, 2005). Blending of starch holds the advantage that the material’s properties can be adjusted by modifying the composition of the blend while also being a low-cost process. 5.2.1.5 Cellulose and cellulose derivatives Cellulose is an omnipresent polysaccharide consisting of a linear chain of β(1-4) linked D-glucose units that forms the backbone of many excipients used in marketed drug products. It is the main component of the cell wall of green plants. The characteristic intramolecular hydrogen bonding results in cellulose being insoluble in most solvents, including water. This was addressed by the development of chemically modified celluloses (e.g., hydroxypropyl) that can be soluble both in water and organic solvents, thus opening the way for their use as DDS (Salimi et al., 2016; Moon et al., 2011).

5.2.2 Proteins Proteins are high-molecular weight polymers where amino acids act as monomers, linked together by characteristic amidic bonds. Being the main structural components in the human body, proteins have been extensively researched for various applications,

Biodegradable nanomaterials

including DDS. The most successful proteins so far in drug delivery include collagen, gelatin, and albumin. 5.2.2.1 Collagen Collagen is the main component of connective tissues and is synthesized by fibroblasts of connective tissue and bone osteoblasts. There are many types of collagen, united by three polypeptide chains that are assembled in a triple-helix supramolecular structure. Collagen in mammals can be rapidly degraded in its amino acid components by enzymes such as collagenase and metalloproteinases. On the other hand, it is a biomaterial difficult to sterilize without altering its original structure; therefore, the use of collagen for drug delivery could be problematic, also due to adverse immune reactions (Parenteau-Bareil et al., 2010). A valid alternative is represented by gelatin (Santoro et al., 2014). 5.2.2.2 Gelatin Gelatin is produced by acid-, alkali-, or enzymatic hydrolysis of collagen. To obtain gelatin, the triple-helix structure of the collagen is disintegrated, transforming the progenitor structure into a single-helix polymer (Kuijpers et al., 1999). Gelatin has the same properties of biocompatibility and biodegradability of collagen but is less prone to the induction of adverse immune responses. Gelatin is crosslinked with agents such as glutaraldehyde in order to achieve lower drug release rates (Santoro et al., 2014). 5.2.2.3 Albumin Albumin, a globular protein, is the most expressed serum protein with its main functions being, maintaining oncotic pressure and acting as a molecule carrier in the plasma (Hawkins et al., 2008). It has low immunogenicity and can be chemically modified easily. The natural transport function, the presence on its surface of aminic, carboxyl, and thiol moieties able to form covalent bonds with different drugs or other proteins, and cellular interactions provides rational for the exploitation of albumin for drug delivery (Larsen et al., 2016). Another interesting strategy is the fusion of albumin with drug/proteins by connecting the protein gene to that of albumin and subsequently expressing it in a suitable host. The main challenge for albumin remains the complete understanding of cellular interactions, which could lead to intracellular drug delivery applications.

5.2.3 Biopolymers of bacterial origin Bacterial polyesters and polyamides are another class of biopolymers presenting interesting properties, including biodegradability, biocompatibility, and almost no toxicity. They are produced by microorganisms while their composition can be modified by

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modifying the nutrients and/or the culture conditions. Polyhydroxyalkanoates and poly (γ-glutamic acid) are the main polymers studied as DDS. 5.2.3.1 Polyhydroxyalcanoates Polyhydroxyalcanoates (PHAs) are produced by bacteria as an energy reserve and source of intracellular carbon, under nutrient limitation conditions and in excess of carbon source. Several bacterial species can produce PHA, while the nature of the synthesized polymer depends on the starting monomers. The bacteria of the genus Ralstonia produce short-chain PHA (scl) using intermediates from fatty acids catabolism, while from the bacteria of the genus Pseudomonas originate medium-chain PHA (mcl) starting from intermediates of β-oxidation. The poly(3-hydroxybutyrate) (PHB) is the best known and studied both from the biosynthetic point of view and from that of the properties and uses (Kai and Loh, 2014). It was first identified in 1926 by Lemoigne, a microbiologist at the Pasteur Institute in Paris, as a constituent of the microorganism Bacillus megaterium in the form of lipid-like sudanophilic inclusions soluble in chloroform. Pure and perfectly isotactic and always presents the R configuration and is therefore optically active. The PHB extracted from bacteria has crystallinity, ranging from 55% to 88%. PHB has a highglass transition temperature and low-impact resistance, a problem addressed with the insertion of 3-hydroxyvalerate units in the PHB, which results in a poly (3-hydroxybutyrate-co-3-hydroxyvalerate) P copolymer with less fragility (Shrivastav et al., 2013). 5.2.3.2 Poly(γ-glutamic acid) This is a nonimmunogenic, biocompatible, anionic, BP made up of repeating units of L-glutamic acid, D-glutamic acid, or both, produced by microbial fermentation. The glutamic acids in poly(γ-glutamic acid) (γ-PGA) are polymerized via γ-amide linkages, so they are not susceptible to proteases. Its applications range from food industry to medical research including drug delivery and tissue engineering (Khalil et al., 2017). Optimization of the production and purification procedures in order to reduce costs is required for γ-PGA to become a more viable option in drug delivery.

5.3 Synthetic polymers Although for the first controlled release resorbable systems polymers of natural and semisynthetic origin were used, synthetic polymers have proven over time much more suited to this type of applications since being of nonnatural production, makes it possible to exercise greater control over their degradation profile and functionality. The main classes of synthetic BPs utilized thus far are summarized in Table 5.2.

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Table 5.2 Main synthetic biodegradable polymers utilized as drug delivery systems. Polymer class Main polymers General chemical Benefits utilized structure (where applicable)

Polyesters

Polyanhydrides

Poly lactic acid

Low cost Excellent mechanical properties

Polyglycolic acid

Monomer of natural origin Many commercial vendors

Poly(lactic-coglycolic acid)

Tunable degradation rate Highly processable

Poly-ε-caprolactone

High flexibility and processability

Poly(sebacic acid)

Monomer flexibility Tunable degradation rate

Poly-(orthoesters)

Controllable and pHsensitive degradation

Poly(alkyl cyanoacrylates)

Tunable degradation rate

Synthetic Pseudoaminoacids poly(amino acids)

Poly(ethylene glycol)-conjugated aminoacids

Variable structures depending on aminoacid

Enzyme-specific degradation Aminoacids as degradation products Increased circulation time

Polyphosphazenes

Synthetic flexibility Phosphorus polyvalency

Polyphosphates

Unique degradation kinetics Phosphorus polyvalency

5.3.1 Aliphatic polyesters Polyesters are the materials most studied in this field because of the presence of the ester bond in their skeleton, which can be hydrolyzed in the biological environment. The most used are polyglycolic acid, polylactic acid, and polylactic-co-glycolic acid surface, thus protecting the drug within the matrix (Singh and Tiwari, 2010). Aliphatic polyesters represent the class of synthetic BPs mainly researched. The synthesis of aliphatic polyesters with polycondensation of diols and dicarboxylic acids has

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been known since 1930. However, the low-melting point, high-hydrolytic instability, and low-molecular weights of the polymers initially obtained limit their use in mechanical applications. On the other hand, the susceptibility in hydrolysis of these polymers serves a variety of medical applications like absorbable sutures. These limitations placed research focus in new synthetic routes in order to form polymers of highmolecular weights. As a result, new routes like the ring-opening polymerization of cyclic diesters were developed. The monomers mainly used for biomedical applications are lactide, glycolide, and caprolactone. 5.3.1.1 Polyglycolic acid Polyglycolic acid was the first synthetic BP used as a resorbable and implantable DDS. Its use has given way to the design of devices that do not require an explant procedure, as opposed to the systems based on nondegradable materials previously used. Its first application dates back to 1962, as a bioabsorbable suture produced by the American Cyanamide Company (Singh and Tiwari, 2010). It is an aliphatic polyester obtainable at low cost and in a highly crystalline form (up at 45%
55% of crystallinity); it has a high-elastic modulus and a low solubility in solvents organic. It is characterized by a glass transition temperature is of 35 C
45 C and a very high-melting point (more than 200 C). It loses its strength in 1
2 months degrading through hydrolysis and undergoes a loss of mass within 6
12 months. Moreover, thanks to its high crystallinity, it has excellent mechanical properties: used in a copolymer, reinforces the structure of the material more than any other degradable polymer used for medical applications (Lakshmi, 2007). 5.3.1.2 Polylactic acid In the last 15 years, PLA has established itself among the most used BPs for the production of resorbable implants and devices. It belongs to the family of aliphatic polyesters and derives from lactic acid (monomer), of natural origin, produced by the bacterial fermentation of carbohydrates ( Jiang and Zhang, 2017). Unlike glycolic acid, lactic acid is a chiral molecule and occurs as enantiomer of type L or D. The polymerization of these monomers leads to the formation of semicrystalline polymers. The polymer deriving from the L-type enantiomer is the poly-L-Lactide (PLLA), having a degree of crystallinity of 37%, depending on the molecular weight and the production process. Presents a glass transition temperature of 55 C
65 C and a melting temperature of about 175 C. An important feature to consider is the very long degradation time, when compared with other polymers used as absorbable DDS: PLLA takes more than a year to degrade in the body through the hydrolysis process, and from 2 to 5 years to be totally expelled from the organism (Middleton and Tipton, 2000). This property derives from the high hydrophobicity of the polymer and is also influenced by the degree of the material’s porosity.

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The poly-DL-Lactide (PDLLA) (or more simply PLA) derives instead from the polymerization of a mixture of monomers of type L and D; it is totally amorphous, with a glass transition temperature of 45 C
60 C. Due to its noncrystalline nature, it has an elastic modulus lower than the PLLA, and for the same reason degrades faster: in 1
2 months it undergoes a loss of mechanical resistance and loses mass in 12
16 months (Maurus and Kaeding, 2004). Depending on the specific applications and needs, one can choose to use PLLA or PDLLA; generally, the polymer deriving from poly-D-Lactide (PDLA) does not find a great use as DDS. PLA can be easily obtained in the form of fibers, films, and sheets; it is degraded followed by hydrolysis of ester bonds and does not require enzymes that catalyze this reaction. Precisely for this reason, following manufacture, thermal stability is required to prevent spontaneous degradation and maintain the properties and molecular weight; in fact, for temperatures above 200 C, it degrades by hydrolysis, by cleavage of the main chain due to oxidation, and by intra- or intermolecular transesterification. The degradation of PLA depends from time, temperature, and the presence of lowmolecular weight impurities and catalysts (Cai et al., 1996). 5.3.1.3 Polylactic-co-glycolic acid It is a statistic copolymer obtained by copolymerization of cyclic dimers glycolic and lactic acid with the purpose of modulating the properties of the two homopolymers. Often identified by the ratio between monomers used, the PLGA, in its various forms, tends to be more amorphous than it is crystalline, with a Tg between 40 C and 60 C. Soluble in many common solvents, unlike the constituent homopolymers, the PLGA degrades to a speed directly proportional to the content of glycolic acid. Used for the first time in 1974 as a suture material, today finds various applications also thanks to the modularity of its properties: for example, in the form of nanospheres, nanofibers, microspheres, and microcapsules are used for controlled release devices of drugs, drugs chemotherapeutics, antibiotics, proteins, analgesics, antiinflammatories, and molecules of RNA. The problems related to the use of this polymer, however, are the difficulty of modulation of the degradation rate and the high acidity of the products resulting from the latter (Makadia and Siegel, 2011). 5.3.1.4 Poly-ε-caprolactone Poly-ε-caprolactone (PCL) is a semicrystallized aliphatic polyester obtained with a polymerization process by opening the caprolactone ring. It is used in some applications such as DDS for its biocompatibility and degradation properties in physiological environment. The PCL in fact, while undergoing a process of hydrolysis much slower compared to that of PLA, it is degraded in the body thanks to the action of enzymes. It has high flexibility and processability and can be obtained in the form of fibers or film. PCL is almost always used to form copolymers with PLA: high strength

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and low-glass transition point of the polycaprolactone give the final material greater ductility and robustness ( Jiang and Zhang, 2017). Traditional polyester-based DDS absorb water during degradation which leads to a bulk erosion of the system leading to stability problems, such as the hydrolysis of sensitive drugs. In addition, zero-order kinetic drug release was difficult to achieve. Therefore, investigators have begun research on highly hydrophobic polymers with low-hydrolytic stability, where degradation of the polymer may be limited to the surface, thus protecting the bulk material. The surface would degrade gradually thus giving even more control on the release kinetics, leading to the development of DDS composed of different BPs such as poly-(orthoesters) (POE), polyanhydrides, poly (alkylcyanoacrylates), and synthetic poly-aminoacids.

5.3.2 Poly-(orthoesters) Poly-(orthoesters) are hydrophobic polymers developed in the early 1970s and have evolved through four families, POE I
IV. Resulting hydrophobicity of the matrix limits water penetration, thus confining erosion to the surface and leading in a more controlled release of the drug, in contrast with bulk erosion of polylactides. POE IV is essentially a derivative of the most hydrophobic POE II containing a segment based on lactic or glycolic acid. As POE II is extremely hydrophobic, presence of water needed for the hydrolysis of the ortho esters is limited, leading to low-degradation rates. Addition of these segments provides a catalyst (ɑ-hydroxy acids produced by hydrolysis of the segments) for the hydrolysis of the ortho esters. Concentration of the ɑ-hydroxyl acids is directly related to the degradation rate, allowing for better control of the erosion process (Heller et al., 2002).

5.3.3 Polyanhydrides Polyanhydrides were first used in 1996 for controlled drug delivery devices after the Food and Drug Administration, characterized their degradability and biocompatibility properties through various in vivo and in vitro studies (Katti et al., 2002). The polyanhydrides derive from monomers containing the CO
O
CO functional group and were first reported in 1909 by Bucher and Slade. They are among the polymers that undergo the fastest hydrolytic degradation process, due to their aliphatic bonds in the basic structure extremely reactive in the presence of water. A property that has made them suitable for the production of DDS is to undergo surface erosion, thanks to their high degree of hydrophobicity combined with the tendency to a fast process of hydrolysis ( Jain et al., 2005). Synthesis is mainly achieved by condensation of diacids, open-chain anhydride polymerization, interfacial condensation, and dehydrochlorination of diacid chlorides.

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5.3.4 Poly(alkyl cyanoacrylates) Poly(alkyl cyanoacrylates) (PAC) present the unique property of undergoing ester side-chain hydrolysis by esterases. The biodegradation process is followed by solubilization and, precisely thanks to this characteristic, these polymers have been used for many years as a glue for tissues in surgery (Vauthier et al., 2003). The main advantage of PAC is that their degradation rate can be tuned by modifying the length of the alkyl chain. PAC rate of degradation is also correlated with toxicity (Kante et al., 1982). More specifically, smaller alkyl chains with higher rate of degradation exhibit some toxicity, whereas larger alkyl chains with lower rate do not. Little control on many factors regarding drug loading and release has led to different results and subsequently to no commercial DDS.

5.3.5 Synthetic poly(amino acids) Synthetic poly(amino acids) present many potentials advantages like biomaterials. Given the large number of existing amino acids, many different types of polymers and copolymers can be produced that easily bind drugs and small peptides. Many are highly insoluble and not processable, their degradation is difficult to control in vivo because they are enzymatically degraded, and the levels of enzymatic activity vary from person to person. Only a few of these can be used for drug delivery applications (e.g., polyglutamic acid derivatives). To tackle these problems pseudo-polyamino acids were proposed in 1984 (Kohn et al., 1986). The pseudo-polyamino acids, thanks to the presence of nonamide bonds between the amino acids, have better chemical
physical characteristics compared to polyamino acids. Other attempts include the preparation of copolymers containing blocks of aminoacids and synthetic polymers like poly(ethylene glycol) (PEG).

5.3.6 Inorganic biodegradable polymers So far, we have reported synthetic BPs of organic origin. However, there are also inorganic polymers that in their backbone instead of carbon have other atoms such as nitrogen or phosphorus. These polymers are considered remarkable biomaterials because of their exceptional features. 5.3.6.1 Polyphosphazenes Polyphosphazenes are hybrid polymers combining an inorganic backbone with organic side chains. More specifically the main chain consists of phosphorus and nitrogen atoms with alternating single and double bonds while organic substituents are linked to the phosphorus atoms as side groups. The phosphorous
nitrogen backbone exhibits extraordinary synthetic flexibility, while variation of side chains lead to polymers with different physicochemical characteristics and degradation rates

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(Teasdale and Brüggemann, 2013). In fact, depending on the side groups, polyphosphazenes can be biodegradable or nonbiodegradable. In contrast to organic BPs, presence of pentavalence phosphorus allows the development of prodrugs and or conjugation of targeting moieties. 5.3.6.2 Polyphosphates Polyphosphates is another class of phosphorus-containing polymers widely studied because of their structural similarities to biological macromolecules like DNA (Chaubal et al., 2003). Their backbone consists of repeating phosphoester groups, which possess unique degradation kinetics, and they can be modified with the introduction of functional groups to the pendant alkoxy or aryloxy groups (Singh et al., 2006). Phosphorus multivalency means their physicochemical properties can be easily tweaked. Under physiological conditions they are degraded due to hydrolysis or enzymatic cleavage of phosphate linkages.

5.4 Polymeric nanoparticles 5.4.1 Introduction Polymeric NPs are a novel addition to the field of pharmaceutical nanotechnology. They are a nanotechnology-based system fabricated for pharmaceutical purposes, otherwise known as nanopharmaceuticals. Nanopharmaceuticals may aid pharmaceutical nanotechnology in tissue engineering, disease prevention and treatment, and improvement of DDS (Bhatia, 2016). However, issues such as safety concerns, toxicity hazards, low biocompatibility, and physiological challenges result in limitations for the activity of nanopharmaceuticals. There is certainly a need for safe, biocompatible nanosystems with tunable properties and controllable activity. Polymeric nanoparticles (NPs) are currently being engineered and investigated as carriers for safe, efficient, and stable drug delivery to desired administration sites. Polymeric NPs show many novel properties which make them worthy competitors for other DDS. They are biodegradable, biocompatible, and nontoxic which makes them a safe option as carriers for DDS. An equally important feature is their tunable physical, chemical, and biological properties which provides controllable drug release motives, the ability to codeliver multiple drugs and specificity resulting in more efficient drug delivery and targeting of the administration site. Lastly, there are various fabrication methods for polymeric NPs with the possibility for production in large quantities.

Biodegradable nanomaterials

5.4.2 Properties—advantages of polymeric nanoparticles As stated earlier, biodegradable polymeric NPs present many properties that make them exquisite candidates for carriers in drug delivery. Their main advantages include the following: 1. Size: Polymeric NPs typically range from 10 to 1000 nm in size. The size of NPs gives them an advantage as a drug carrier because they are better suited for intravenous delivery compared to previously developed microsized DDS (Hans and Lowman, 2002). The smallest capillaries in the body have a diameter of 5
6 μm. Therefore, the size of particles circulating through the bloodstream must be significantly smaller than 5 μm, without forming aggregates, to avoid the formation of an embolism caused by the particles themselves. 2. Shape: Different NP structures can be engineered using various methods depending on the properties of the selected polymer (or polymers) and the drug that is to be loaded and delivered. In this context, two major categories of polymeric NP drug carriers can be named based on their form: nanospheres (100
200 nm) and nanocapsules (100
300 nm) (Kumari et al., 2010; Letchford and Burt, 2007). The two forms differ in the way that the polymer is organized structurally. In the case of a nanosphere, the polymer forms a colloidal particle, whereas in the case of a nanocapsule the polymer forms an outer shell wherein exists an aqueous or oily environment (Calzoni et al., 2019). In both cases, depending on the structural organization of the polymeric NP, the drug can either be adsorbed to the surface of the NP or become entrapped within the NP itself (encapsulation). An addition to the concept of a nanosphere is that of a polymersome (5 nm
5 μm). These particles are inspired by the structure of liposomes and consist of a bilayer of amphiphilic block copolymers which surrounds an aqueous environment. The amphiphilic polymers mimic the lipid bilayer of liposomes which inspired the fabrication of this particular polymeric NP and the aqueous center provides an environment in which hydrophilic drugs can be loaded (Letchford and Burt, 2007). Amphiphilic polymers also form polymeric micelles through spontaneous selfassembly due to hydrophobic interactions. Self-assembly formulations involve intermolecular and intramolecular interaction between the polymer molecules and the drug itself. In the case of micelles, amphiphilic polymers accumulate in such a way creating a two-phase system. The polymers orient themselves in such a way that the hydrophobic component phases inward, away from the hydrophilic environment, whereas the hydrophilic end of the polymer molecules is attracted to the environment and remains on the outer part of the micelle structure (Letchford and Burt, 2007; Karlsson et al., 2018). 3. Biocompatibility and biodegradability: Natural polymers, especially, namely, polysaccharides and protein-based polymers, show excellent biocompatibility as they

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4.

5.

6.

7.

8.

can be broken down into polysaccharides and peptides respectively by enzymatic degradation. These biomolecules can then be easily metabolized by the body without any harmful side effects (Nair and Laurencin, 2007). The biodegradability of polymeric NPs not only contributes to their biocompatibility, put it is also a major factor in their drug release ability. Safe, less toxic, nonimmunogenic: The use of polymeric NP’s as carriers in anticancer drug delivery has shown decreased toxicity, as the polymeric component of the carrier provides protection for the drug and limits its interaction with healthy cells (Brewer et al., 2011). Engineered specificity: Depending on the method of preparation polymeric NP are a suited choice for intracellular and site-specific drug delivery (Bhatia, 2016). These systems have the ability to deliver higher concentrations of the drug to specific administration sites, since they can be specially modified in terms of size and surface characteristics in order to reach a specified target cell. Tunable physical, chemical, and biological properties: The use of polymeric NPs as drug carriers provides the liberty of controlled and long-term release rates, prolonged circulation times, and prolonged bioactivity. All these features are crucial in the efficient activity of the drug carrier. The ability to modify drug release motives from the polymeric NPs have made these systems promising contenders for cancer therapy, vaccine delivery, delivery of antibiotics, and contraceptives. Adding PEG to the polymer of the NP has been shown to increase the circulation time of the nanocarrier by hindering uptake by the reticuloendothelial system (RES). PEG achieves this by inhibiting the binding of proteins on the carrier’s surface therefore intercepting its recognition by the RES (Gref et al., 2000). Stability: Polymeric NPs can stabilize volatile pharmaceuticals, protecting them from the environment until the reach the administration site. There exist a variety of polymers from which suitable ones may be chosen for loading of either hydrophobic or hydrophilic drugs. Multiple drug codelivery: Ability to codeliver multiple drugs with synergistic effects (Brewer et al., 2011). The codelivery of multiple drugs in the same polymeric NP carrier overcomes issues such as multidrug resistance of tumor cells and achieves synergistic effects between different drugs (Bhatia, 2016). Wang et al. developed a polymeric NP carrier consisting of hyaluronic acid-decorated PLGA, pluronic F127 and chitosan for the codelivery of doxorubicin hydrochloride and irinotecan, which are a hydrophilic and hydrophobic anticancer drug, respectively. This dual-DDS shows high efficacy and increased effect against cancer stem cells in vitro and in vivo, which the researchers attributed to the synergistic activities of the two anticancer drugs as well as the capability of the polymeric NPs in efficiently delivering higher amounts of the drugs to the target cells (Wang et al., 2015).

Biodegradable nanomaterials

Figure 5.2 Surface adsorption and encapsulation of drug in nanospheres and nanocapsules. Courtesy from Kumari, A., Kumar Yadav, S., Yadav, S.C., 2009. Biodegradable polymeric nanoparticlesbased drug delivery systems. Coll. Surfaces B Biointerf., doi:10.1016/j.colsurfb.2009.09.001.

9. Multiple fabrication methods: The most suitable fabrication method for synthesizing polymer NPs depends on the properties of the polymer and the nature of the drug that is to be loaded for delivery. The two general categories of fabrication processes are top down and bottom up (Figs. 5.2 and 5.3).

5.4.3 Polymeric nanoparticles preparation Nanoparticulate systems can be prepared by utilizing different methods depending on the characteristics of the polymer and active ingredient, the site of action, and the therapeutic regime. The most commonly used techniques for the formulation of NPs for drug delivery involve the use of preformed polymers. 5.4.3.1 Emulsion—solvent evaporation This method, widespread due to its simplicity and versatility, was originally adopted for the encapsulation of lipophilic drugs. The polymeric material and the active ingredient are dissolved in a volatile organic solvent, immiscible with water (dichloromethane, chloroform, or acetonitrile), and the solution obtained is emulsified with an aqueous phase containing suitable stabilizing agents. The solvent is then evaporated, generally at elevated temperatures and reduced pressure, with consequent precipitation

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Figure 5.3 Schematic summary of different methods for nanoparticles synthesis.

of the polymer and formation of solid NPs. In the case of hydrophilic drugs, instead, a multiple water/oil/water emulsion (w/o/w) is required. The aqueous solution of the drug is emulsified with an organic phase consisting of solvent and polymer, so as to obtain a primary emulsion w/o (water/oil). The mixing of this first emulsion with an excess of water, under continuous stirring, generates the second emulsion w/o/w, from which the solvent can be removed by evaporation or extraction (Chiellini et al., 2008). The main drawback of this method lies in the use of organic solvents, which, in addition to being significantly toxic, can compromise the stability of the drug incorporated within the NPs (Park et al., 2005). 5.4.3.2 Coacervation Coacervation is a physical phenomenon of phase separation, typical of polymer dispersions, which occurs due to changes in pH and temperature or due to the addition of salts or solvents. This process is widely exploited for the preparation of micro- and nanocapsules, constituting one of the most common approaches also at the industrial level. After dispersing the active principle in a solution of the chosen polymer, coacervation of the system is induced leading to deposition of the polymeric material on the surface of the drug particles to form a continuous layer. The solidification of the

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coacervate, obtained by cooling or desolvation, leads to the formation of a rigid and resistant membrane containing the drug (Vilar et al., 2012).

5.4.3.3 Nanoprecipitation, coprecipitation, and dialysis In the previous techniques, the formation of NPs occurs as a consequence of the removal of the solvent from the polymer solution in the presence of a nonsolvent. In the nanoprecipitation method, the polymer solution containing the drug and any stabilizing agents is added dropwise to a nonsolvent, miscible with the solvent used to dissolve the polymer. NPs are formed instantly thanks to the rapid desolvation and aggregation of the polymer chains. The encapsulation efficiency is poor in the case of hydrophilic drugs, which tend to spread rapidly in water as soon as the organic phase in which they are dissolved (acetone and ethanol) comes into contact with an aqueous solvent (Legrand et al., 2007). A variant, certainly more advantageous for the encapsulation of water-soluble molecules, is represented by coprecipitation. The polymeric material is dissolved in a water miscible solvent and injected into an aqueous solution of the active ingredient. This procedure does not require the use of aggressive solvents and therefore allows the incorporation of proteins and peptides into the NPs without altering their functionality. In addition, with the aforementioned methods, dialysis also allows obtaining particles with a very narrow size distribution. A solution of the polymer, the active ingredient, and a surfactant in an organic solvent is prepared and placed inside a dialysis tube with adequate cutoff and dialyzed against a nonsolvent miscible with the previous one. The solvent inside the semipermeable membrane is displaced, which determines the loss of solubility and the progressive precipitation of the polymer in the form of NPs. The formation of large aggregates and their interaction with the dialysis membrane considerably limit the applicability of this process (Piras et al., 2010). 5.4.3.4 Ionic gelation NPs can form following the interaction of polyelectrolytes of opposite charge, generally operated in aqueous media. Polymers most used in this type of formulation are of natural and semisynthetic origin, such as alginate, chitosan, hyaluronic acid, pectin, and carboxymethylcellulose. The mixing of a polymeric solution containing the drug with an aqueous solution of polyvalent ions of opposite charge determines the three-dimensional crosslinking of the polymeric chains with consequent formation of stable complexes. Thanks to its mild preparation conditions, ionic gelation is particularly suitable for labile active ingredients such as proteins and nucleic acids (Chiellini et al., 2008).

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5.4.3.5 Spray drying Spray drying is a relatively new method, which offers several advantages, such as speed, good reproducibility, and scalability of the process, without a particularly high cost. A solution (or a dispersion) of the active ingredient in a suitable polymer solution is prepared. The atomization of this solution (or suspension) inside a hot air stream generates nanometric drops, from which the solvent evaporates rapidly with consequent formation of solid particles (He et al., 1999).

5.4.4 Drug release mechanisms The release of the drug from NPs occurs through a combination of diffusion and degradation mechanisms. Regarding the nanocapsules, release is controlled mainly by the diffusion of the drug through the polymeric membrane that surrounds it or the aqueous microporous network that interpenetrates the dense areas of the membrane. The active principle is released under the influence of a concentration gradient, according to a diffusion mechanism regulated by Fick’s first law: dM dC 5 2 DS dt dx

(5.1)

where dM/dt represents the diffusion rate, dC/dx is the concentration gradient through the membrane, D is the diffusion coefficient, and S is the overall surface area of the system. At equilibrium, drug release is constant over time and largely depends on the diffusion capacity of the drug, therefore on its affinity for the membrane, on the nature, on the thickness, and on the porosity of the membrane (Chien and Lin, 2007). In the case of nanospheres, the active ingredient is released through diffusion processes as well; however, the release rate and the concentration gradient at each point of the polymeric matrix vary as a function of time due to the continuous decrease in the quantity of drug contained within the system. In this case, the diffusion phenomenon is more correctly described by Fick’s second law: @ϕ Dð@^2ϕÞ 5 @t ð@x^2Þ

(5.2)

where ϕ is the concentration, ϕ 5 ϕ(x, t) is a function that depends on location x and time t, t is time, D is the diffusion coefficient, and x is the position (length). Furthermore due to interaction with biological fluids, the polymeric material can undergo hydrolytic and/or enzymatic degradation. This process contributes significantly to the release of the drug, also influencing rate of diffusion through the matrix or the polymeric membrane. The chemical composition of the polymer results therefore, decisive in the drug release kinetics making optimization of its activity profile possible (Vilar et al., 2012).

Biodegradable nanomaterials

Finally phenomena of solvation and swelling of the NPs may occur: the hydration of these systems within biological fluids causes distension and relaxation of polymeric chains, allowing the encapsulated agent to diffuse in the external environment (Peppas et al., 2000). This process is especially important in the case of hydrogels, that in the presence of water are able to increase up to five times their own volume. Theirs swelling can be induced by specific stimuli, such as changes in pH, temperature, or ionic strength, guaranteeing even more precise release control.

5.4.5 Targeting A crucial area of focus in drug delivery is the accurate targeting of a desired cell or tissue. In DDS, targeting refers to the drug carrier’s ability to deliver the cargo to the correct site at the correct time. An ideal and efficient nanocarrier should recognize a specific target cell or tissue, bind to it and deliver the drug all while avoiding unwanted drug induced side effects to healthy cells and tissues. Polymeric NPs may have engineered specificity through surface functionalization and manipulation of their tunable properties, allowing them to deliver a higher concentration of a drug to a desired site. DDS follow two major targeting mechanisms: Passive targeting and active targeting. In passive targeting, the NP reaches the target administration site passively, without the need for attaching additional ligands to serve this purpose. Active targeting, on the other hand, is a term that describes specific interactions between the drug carrier and the target cells, usually through specific ligand
receptor interactions. In active targeting, a ligand is used as a homing device through which the nanocarrier is conjugated to the target cell or tissue wherein the pharmaceutical agent is released after endocytosis (Varshosaz and Farzan, 2015). 5.4.5.1 Passive targeting The NP carrier reaches the target administration site without the aid of conjugated target-specific ligands to serve as homing devices. An example of passive targeting is the preferential accumulation of chemotherapeutic agents in solid tumors. This occurs as a result of the enhanced vascular permeability of tumor tissues compared with healthy tissue (Kaparissides et al., 2006). The NPs accumulate preferentially in the neoplastic tissues of the tumor as a result of the enhanced permeability and retention (EPR) phenomenon, first described by Maeda and Matsumura (Maeda, 2012; Matsumura and Maeda, 1986). The EPR is the result of the abnormal vasculature and impaired lymphatic drainage within neoplastic tissues. The size of NP drug carriers allows them to enter the tumor tissues through the EPR phenomenon (Bazak et al., 2014). Passive targeting holds disadvantages compared to active targeting. The NPs encounter obstacles on the route to the target site, namely, extracellular drug release and uptake by to nontarget cells.

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5.4.5.2 Active targeting In this targeting mechanism, the surface of the NP carrier is functionalized with ligands designed to act as homing devices by recognizing and binding to specific administration sites. These homing devices can be small organic molecules, peptides, antibodies, designed proteins, and nucleic acid aptamers. For example, specific antibodies can be used as homing devices for targeting of tumors. These antibodies recognize characteristic molecules on the tumor’s surface, which do not exist on healthy cells. By conjugating these antibodies to the NP, the DDS can then be led specifically to the tumor cells while leaving the normal cells unaffected (Abdolahpour et al., 2018). Since ligand
receptor interactions can be highly selective, this could allow a more effective targeting of the desired site, making active targeting a more appealing route than passive targeting. The advanced specificity due to active targeting also lowers toxicity and unwanted side effects that occur with nonspecific drug delivery to untargeted cells and tissues (Fig. 5.4).

Figure 5.4 Active targeting of tumor cells. The ligands on the nanoparticle surface recognize and bind to receptors on the tumor cells resulting in endocytosis Adopted by Zhang L., Li, Y., Yu, J. C., Chemical modification of inorganic nanostructures for targeted and controlled drug delivery in cancer treatment, 2014, J. Mater. Chem. B, 2, 452.

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5.4.5.3 Tumor targeting Utilizing polymeric NPs for tumor targeting is based on both passive targeting via the EPR effect and active targeting through surface ligands (Bhatia, 2016). According to Bae and Park (2011), the active targeting of tumors is assisted first by the EPR effect through which the drug-carrying NPs leave the blood stream and reach the general target area. There ligand
receptor interactions occur between the NP drug carrier and the targeted tumor cells. According to their perspective, the EPR effect and blood circulation time of the drug carrier are increased by the PEGylation of the NPs, that is, modification of the NPs with PEG. The improved EPR and circulation in the blood stream assists the drug carrier to reach the general target area more efficiently so that the active targeting may occur. Through using NPs as drug carriers, the drug’s distribution is limited to the target organ, and exposure to healthy tissues and organs is limited. Verdun et al. (1990) studied the effects of doxorubicin entrapped in polyisohexylcyanoacrylate-based NPs in mice compared to free doxorubicin. They found higher concentrations of the drug in the liver, spleen, and lungs of the mice treated with the encapsulated doxorubicin compared to the mice treated with the free doxorubicin. This shows the targeting ability of the nanocarrier and its recognition and preference of these specific organs. Distribution study cyclic arginylglycylaspartic acid (RGD)-doxorubicin-NP comprised of inulin multimethacrylate with a targeting peptide was carried out by Bibbly et al. in tumorbearing mice. Results showed that the drug concentration in the liver increased over time and decreased in the heart, lungs, kidneys, spleen, and plasma. This accumulation in the liver resulted in the maximum injected dose resulting in the liver and only a minute percentage reaching the tumor 48 hours after administration (Bibby et al., 2005). This study, as well as others, reveals the tendency of NPs to accumulate in the liver via uptake by the mononuclear phagocytic system (MPS). Therefore a challenge for tumor targeting with nanocarriers is to avoid their uptake by the MPS. However, this phenomenon can be manipulated to efficiently deliver chemotherapeutic agents via NPs to tumors in MPS-rich organs and tissues, for example, hepatocarcinoma, gynecological cancers, mall cell tumors, and others. Due to this effect, conventional anticancer drugloaded NPs are limited to MPS-rich organs (Bhatia, 2016; Fig. 5.5).

5.5 Clinical applications of biodegradable nanoparticles 5.5.1 Introduction The use of BPs is particularly appealing in clinical practice as these materials, due to their biodegradable nature, get eliminated from the body after fulfilling their purpose.

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Figure 5.5 Tumor sites can be reached either by passive targeting (EPR effect) or/and active targeting (Ligand - Receptor interaction).

Due to their vast versatility in nature and structure BPs can be tailored to possess specific, physical, chemical biological, functional, biomechanical, and degradation properties, which can serve the purpose of particular biomedical applications. The medical applications of BPs cover a rather heterogeneous field ranging from tissue engineering and tissue adhesives to implants and DDS. Several biodegradable polymeric materials have extensively been used with minor modifications for decades as sutures, orthopedic pins and nails, hemostatic sponges, fixation plates, and filaments. These materials bear a known safety profile, and their use in humans is approved by health authorities. Repurposing of these materials is a very cost-effective strategy, taking into account the time and effort needed to introduce a new material in clinical use. The plasticity of their polymeric nature allows modifications and varying compositions to be used to fit the purpose of new biomedical applications. Most of the already known natural and synthetic biodegradable polymeric materials can be formulated in the nanoscale matching the scale of biological systems, such as viruses, membranes, and protein complexes which are natural nanostructures. In this context, NPs with variable structure and properties have been formulated for various therapeutic applications (Song et al., 2018). The use of BPs as nanocarriers in drug delivery is particularly promising in drug development. Several studies have showed that BPs can be used to enhance biocompatibility, provide better encapsulation, and sustained release of drug molecules (Kumari et al., 2016). The potential of providing site-specific delivery for drugs, biomolecules, proteins, and peptides has led to a huge research output in this field.

Biodegradable nanomaterials

5.5.2 Regulatory aspects Despite the huge research efforts and investment by academia and pharmaceutical companies, regulatory approvals of novel nanomedicine products have not exceeded 10%, mostly because of failures in terms of efficacy and safety profiles during nonclinical and clinical studies. Regulatory guidance on nanomedicines is yet evolving and appears to be rather stratified and far from unambiguous across, the globe. A case by case approach is generally applied and request for product specific scientific advice by applicants is highly encouraged by regulatory agencies. Whatsoever there have been developed several concept papers by the European Medicines Agency (EMA) and a concise guidance document by the Food and Drug Administration (FDA) (Musazzi et al., 2017). Even if a BP is well characterized and extensively used in clinical practice in the macro- or microscale, its nanoscale applications undergo a separate and extensive regulatory review, due to the different toxicological profile of particles with dimensions less than 1 μm. Even minor changes in composition and/or physicochemical properties of NPs could result in clinically significant changes regarding pharmacodynamic, pharmacokinetics, and toxicity. Therefore, detailed characterization of drug products, identification of the critical attributes of the products, and the manufacturing process to achieve batch to batch consistency are a prerequisite along with in depth studies of how quality aspects of the product influence safety and efficacy profiles of the product. Safety concerns regarding NPs include potential infusion reactions, hypersensitivity reactions, oxidative stress, biodistribution and permanence, impact on immune system (Halamoda-Kenzaoui and Bremer-Hoffmann, 2018), unexpected toxicity effects due to increased reactivity, and permeability (Sufian et al., 2017).

5.5.3 Properties of biodegradable nanoparticles (BNPs) which impact clinical use The mechanical properties of BPs can easily be manipulated to fit the purpose of use. Strong mechanical properties are generally required for materials to be used as tissue adhesives or for tissue engineering, that is, the material should remain sufficiently strong until the surrounding tissue has healed, while for drug delivery more processable forms are required. Degradation time must match the time required for biomedical application. In vivo metabolism of the polymer matrix after fulfilling its purpose should invoke nontoxic constituents that can be easily eliminated to avoid a toxic response. Furthermore, material should be easily processable in the final product form with an acceptable shelf life and easily sterilized. The size and shape of NPs may affect solution dynamics in blood vessels. For example, it has been shown that NPs larger than 200 nm are collected in liver and

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spleen. Spherical geometry is typical for NPs designed for intravenous delivery to facilitate rheology and reduce immunogenicity (Ahlawat et al., 2018). Long rodshaped NPs on the other hand have been proved to show longer blood circulation and a highest bioavailability compared with short rod and spherical NPs (Zhao et al., 2017). Surface composition can also affect the course of elimination after delivery. Surface hydrophobicity enhances plasma protein adsorption and removal through opsonization and MPS. Surface charge can also affect the in vivo fate of NPs. Positively charged NPs are more quickly eliminated than negatively charged ones, while neutral NPs persist longer in blood. Surface characteristics can also affect sticking to vascular walls and uptake as well as cell removal by efflux pumps. NPs entering blood circulation tend to be favorably collected at sites of inflammation, that is, injuries, tumors, and infections. This constitutes a passive targeting due to the loose vasculature of endothelial cells, an effect known as EPR. This effect can be significantly enhanced by prolonging the presence of NPs in blood circulation by making NP invisible to plasma proteins. Stealth properties can be added by coating of NPs with hydrophilic or neutral groups after absorption or making copolymers of BPs with hydrophilic polymers such as polyvinyl alcohol (PVA), polyvinylpyrrolidone (PVP), PEG, or polysaccharides. The presence of hydrophilic and neutral chains at the surface can repel plasma proteins which results in prolonged presence in blood circulation and enhanced bioavailability (Suk et al., 2016). Modification of the NPs surface can also include conjugation with moieties that promote active targeting of tumor cells through NPs to overexpressed receptors of tumor cell membrane and phagocytosis or endocytosis mechanisms (Pillai, 2014).

5.5.4 Approved and investigational drugs with biodegradable polymeric nanoparticles of natural or synthetic origin In a recent analysis (D’Mello et al., 2017) of almost 350 submissions in FDA of nanomaterial-containing drug products since 1973, liposomes were the most prevalent category (33% drug products), followed by nanocrystal-containing drug products (23%). Emulsions formed 14% of the submissions, iron
polymer complexes 9% and micelles 6%. Drug products containing other nanomaterials, such as drug
protein complexes, drug
polymer complexes, and polymeric NPs, accounted for 14% of the overall applications. In another review (Bobo et al., 2016), a list of FDA-approved nanomedicines included among others, 2 drugs with protein NPs combined with drugs or biologics and 15 drugs containing polymer NPs with drugs or biologics. Most of these drugs were proteins PEGylated to improve stability, circulation time, or immunogenicity profile. Regarding material categories under investigation, it was shown that micellar, metallic, and protein-based particles entering the development process are increased in comparison to what has previously been approved.

Biodegradable nanomaterials

Natural and synthetic polymers have been explored for the synthesis of biodegradable NPs for drug delivery of anticancer, psychotic, antimicrobial and plant isolated, protein, peptides, and drugs. Natural polymers include albumin, gelatin, and chitosan, whereas synthetic polymers such as PLGA, PLA, PCL, and PACare drawing increasing attention for the last two decades (Kumari et al., 2016). The most common applications include development of improved chemotherapeutic delivery systems and improving oral bioavailability of hydrophobic and peptide/protein drugs. Future trends include application of biodegradable NPs in DNA/RNA delivery and gene therapy. 5.5.4.1 Anticancer drugs Tumor targeting with single or multiple chemotherapeutic agents to reduce drug doses and produce synergistic effects is a very appealing strategy in cancer research. Various polymers with a wide range of physicochemical properties have been utilized for delivery of cytotoxic drugs or other therapeutic agents (e.g., chemosensitizers, differentiation-inducing, and neovasculature disruption agents). Biodegradable materials used for this type of NPs include natural polymers (e.g., proteins, gelatin, chitosan) as well as synthetic polymers such as PLGA, poly(ethylenimine), poly(L-lysine), and PEG (Mokhtarzadeh et al., 2016; Afsharzadeh et al., 2018). Several approved and investigational drugs will be presented below. Protein NPs can encompass drugs attached to endogenous protein carriers, modified proteins where the active therapeutic is the protein itself, or composite functionalized platforms that rely on protein motifs for targeting delivery of the therapeutic agent (Bobo et al., 2016). Early protein NPs exploited the natural properties of serum proteins to facilitate dissolution and transport of drug moieties in blood circulation. An early example of a protein-based nanodrug is Abraxane, which contains albumin NPs (130 nm) conjugated with paclitaxel and was approved in 2005 by FDA. It was designed to improve paclitaxel chemoterapeutic applications by eliminating the need for use of the toxic solvent Kolliphor to solubilize paclitaxel (Bobo et al., 2016). Albuminbound paclitaxel NPs improved infusion time and eliminated the need to concomitantly administer antihistamines and dexamethasone to prevent an immune reaction to Kolliphor. Pazenir is the first generic of Abraxane in EU, another NP albumin-bound paclitaxel, which has been authorized in the EU since 2008. Studies have demonstrated the satisfactory quality of Pazenir. A bioequivalence study versus the reference product Abraxane was not required as Pazenir is administered intravenously and the NPs dissociate rapidly and because of the qualitative and quantitative compositions and the nature and behavior of the products (EMA/CHMP/147314/2019, 2019). After the success of Abraxane, several additional albumin-bound NPs (NABs) have entered clinical trials for the purpose of improving the efficacy and safety of other drugs (Patra et al., 2018). Among these are NAB-docetaxel, NAB-heat shock protein inhibitor, and NAB-rapamycin (Gonzalez-Angulo et al., 2013).

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Ontak (denileukin diftitox, Eisai, Inc.) is an example of protein NP drug combining active targeting proteins and cytotoxic molecules, which was approved in 2008 (Ventola, 2017). A major difference from Abraxane is that instead of using an unmodified protein it utilizes engineered particle complexes designed to enable active targeting (Havel, 2016). It is an interleukin (IL)-2 receptor antagonist that was initially designed to treat an aggressive form of non-Hodgkin’s peripheral T-cell lymphomas by targeting the cytocidal action of diphtheria toxin toward cells that overexpress the IL-2 receptor on T cells (Foss, 2006). Chitosan polymers: Chitosan-based NPs have been developed for the delivery of several combination chemotherapeutic schemes as well as gene carriers. Cationic polymeric NPs of camptothecin and curcumin have been prepared for synergistic colon cancer combination chemotherapy (Xiao, 2015). In another report, PEGylated chitosan NPs were loaded with both methotrexate as a targeting agent against folate receptors and mitomycin C, resulting in increased synergistic anticancer effect (Jia et al., 2014). Chitosan-based NPs have also been prepared for combinations of gefitinib agents with shMDR1 had the potential to overcome the multidrug resistance and improve cancer treatment efficacy, especially toward resistant cells (Yu et al., 2015) Gelatin has also been used as a NP carrier system for paclitaxel forming an amorphous water-soluble system allowing rapid drug releases at the target site (Lu et al., 2004). PLGA is listed as a safe material for human use by FDA, and research shows that PLGA NPs have great potential for the delivery of bioactive agents and applications in gene and vaccine delivery. It is commonly used blended with other polymers such as polypropylene fumarate, polyvinyl alcohol, or chitosan to moderate its acidic nature. Cisplatin (Avgoustakis et al., 2002), Docetaxel (Esmaeili et al., 2008), and Paclitaxel (Wang et al., 2011) are major chemotherapeutics that have been encapsulated in various PLGA nanosystems. Recently PLGA NPs surface engineered with hyaluronic acid for targeted delivery of paclitaxel to triple negative breast cancer cells were developed and showed improved cellular uptake (Cerqueira et al., 2017). Other applications are the encapsulation of Curcumin (Ranjan et al., 2012) and 9-NitroCamptothecin (Ahmadi et al., 2015). PCL is another promising polymer with great biodegradable properties at physiological pH. Various PCL-based NP systems have been investigated for enhances and better targeted delivery of Docetaxel (Zheng et al., 2009), Vinblastine (Prabu et al., 2008), Tamoxifen (Shenoy and Amiji, 2005), and Taxol (Ma, 2005). 5.5.4.2 Nanoparticles for oral delivery Improvement of oral bioavailability of drugs through NP formation is a widely investigated strategy. NPs formation can improve solubility of hydrophobic drugs and protect

Biodegradable nanomaterials

labile peptide and protein drugs and vaccines against the enzymatic and hydrolytic degradation in the gastrointestinal track. Uptake of NPs following oral administration has been shown to occur through apical sodium-dependent bile acid transportermediated cellular uptake, chylomicron transport pathways, and lymphatic uptake of the NPs by the Peyer’s patches in the gut-associated lymphoid tissue (Kim et al., 2018). Bioavailability of antifungal agents has been improved utilizing NPs delivery methods (Pandey et al., 2005) For example, itraconazole encapsulated in PLGA NPs showed enhanced intestinal permeability in an ex vivo study (Alhowyan et al., 2019). Another interesting goal remains the formulation of peptide and protein drugs such as insulin for oral delivery (Wong et al., 2017). In a recent paper, d-α-Tocopheryl polyethylene glycol 1000 succinate (TPGS)-emulsified PEG-capped PLGA NPs as a potential drug carrier for the oral delivery of insulin were synthesized and tested in diabetic rats by oral administration providing encouraging results (Alhowyan et al., 2019). 5.5.4.3 Future trends: nanoparticles for vaccines and gene therapy Polynucleotide and DNA vaccines present certain advantages over many conventional protein-based vaccines presenting better immunization efficiency (Hasson et al., 2015), lower production cost, and better stability profile and handling properties (Xu et al., 2014). BNPs can be effective carriers for DNA vaccines. A nano-chitosan-based DNA vaccine encoding T-cell epitopes of Esat-6 and FL was found to be effective against Mycobacterium tuberculosis infection in mice (Feng et al., 2013). In the field of gene therapy development of biodegradable nanomaterials for oligonucleotide delivery present an excellent opportunity to resolve safety problems usually encountered with either viral or other nonviral vectors. Improving gene transfer activity and functional quality of nanocarriers to optimize and target gene delivery are currently the biggest challenges in developing and applying effective strategies in the use of BPs as carriers for gene therapy (Mokhtarzadeh et al., 2016).

5.6 Future perspectives The use of polymeric NPs as nonviral DDS seems to be a critical area in future nanopharmaceutical research. It is evident that polymeric NPs may offer many advancements in drug delivery and will continue to attract attention from researchers of this field. There has been extensive research regarding nanosized drug carriers based on natural and synthetic polymers; however, there is need for this research to be extended

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into in vivo studying and clinical trials. The application of polymeric NPs as drug carriers has been expanded to serve the delivery of antibiotics, anticancers drugs, vaccines, and genes. This wide range of application is owed not only to the polymers’ intrinsic properties, that is, biocompatibility, biodegradability, nontoxicity, but also to the ability of polymeric NPs to be modified and functionalized to meet certain requirements. These particles can be engineered to have specific properties to aid them in efficient drug targeting and drug release, such as increase circulation time, active targeting of administration site through receptor recognition, and decreased recognition by the immune system. Looking at polymeric NPs from a chemical perspective, it is evident that there is much room for the generation of new and improved polymers. Chemical modification of polymers may lead to the engineering and fabrication of a new range of polymers with optimal chemical, physical, and biological properties. Through modification of polymeric NP drug carriers can be made to achieve improved targeting ability, stimuli-triggered drug release and targeted codelivery of multiple drugs. Much research has been carried out in the case of tumor targeting and anticancer drug delivery via polymeric nanocarriers. Among the future advancements in polymeric NP drug delivery technology, expanding their field of effect to allow treatment of even more pathologies, that is, immune system, and genetic diseases is very high on the list. There is a need for the development of large-scale fabrication processes for preparing large amounts of polymeric NPs. These fabrication methods should be economic, time efficient, and allow for the control of the particle size, the surface properties, and the drug release motives. Despite the advancements made in research of polymeric NP drug carriers, very few examples are currently in clinical trials. There is a need for in vivo testing to better overcome obstacles encountered during in vitro research and enrich the results and information collected thus far. It should be noted that even though these DDS show good results during in vitro studies, they fail during in vivo trials. Clinical trials and in vivo testing can aid in better understanding drug targeting and drug uptake mechanisms and overcoming problems such as this. In conclusion, it can be safely said that despite the great advancements exhibited in current research, there still exist several challenges and obstacles to be addressed and overcome.

Acknowledgments This research has been supported by the European Union and Greek national funds through the Operational Program Competitiveness, Entrepreneurship and Innovation, under the call RESEARCH
CREATE
INNOVATE (T1EDK-02024)

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7056. Younes, I., Rinaudo, M., 2015. Chitin and chitosan preparation from marine sources. Structure, properties and applications. Mar. Drugs 13 (3), 1133
1174. Zhao, Y., Wang, Y., Ran, F., Cui, Y., Liu, C., Zhao, Q., et al., 2017. A comparison between sphere and rod nanoparticles regarding their in vivo biological behavior and pharmacokinetics. Sci. Rep. 7, 4131. Zheng, D., Li, X., Xu, H., Lu, X., Hu, Y., Fan, W., 2009. Study on docetaxel-loaded nanoparticles with high antitumor efficacy against malignant melanoma. Acta Biochim. Biophys. Sin. (Shanghai) 41 (7), 578
587.

Further reading Mao, H.Q., Roy, K., Troung-Le, Vu. L., Janes, K.A., Lin, K.Y., Wang, Y., et al., 2001. Chitosan-DNA nanoparticles as gene carriers: synthesis, characterization and transfection efficiency. J. Control. Release 70 (3), 399
421. Midoux, P., Breuzard, G., Gomez, J.P., Pichon, C., 2008. Polymer-based gene delivery: a current review on the uptake and intracellular trafficking of polyplexes. Curr. Gene Therapy 8 (5), 335
352.

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CHAPTER SIX

Modulating the immune response with liposomal delivery David Nardo1, David Henson1, Joe E. Springer2 and Vincent J. Venditto1 1

Department of Pharmaceutical Sciences, University of Kentucky, Lexington, KY, United States Spinal Cord and Brain Injury Research Center, University of Kentucky, Lexington, KY, United States

2

6.1 Introduction Immunotherapy using lipid-based nanoparticle (LNPs) was recognized over 40 years ago with the introduction of liposomal vaccine strategies by Gregoriadis (Schwendener, 2014; Allison and Gregoriadis, 1974; Gregoriadis, 2016). As our understanding of the immune system, the pathogenesis of disease, and mechanistic understanding of therapeutics continue to evolve, LNPs provide a platform with significant benefit in clinical applications to improve patient care. A number of liposomal formulations have been approved for anticancer, antifungal, and antiangiogenic applications, paving the way for their continued utility in modulating immune responses in disease. The compositional diversity of LNPs provides strategies to manipulate formulation characteristics (i.e., size, charge), encapsulate diverse therapeutics in aqueous (i.e., doxorubicin) or lipophilic (i.e., amphotericin B) compartments, and display a variety of targeting moieties (i.e., carbohydrates, antibodies) to modulate well-defined immune mechanisms. The utility of LNPs is highlighted by significant investment in these technologies, the resulting clinical trials, and approval of various LNP therapeutics by regulatory agencies. The utility of liposomes for drug delivery continues to evolve and the components, formulations, and preclinical evaluation of a multitude of formulations have been reviewed extensively elsewhere (Cullis and Hope, 2017; Bulbake et al., 2017; Watson et al., 2012; Semple et al., 2010). The precedent for clinical translation positions the field of liposomal drug and gene delivery to have a significant impact in the future of immune modulation to reduce the burden of numerous diseases in patients. This chapter is organized to provide a brief overview of liposomes and the immune system followed by substantial discussion of clinically approved liposome-based therapeutics and what lessons can be garnered from these efforts to advance the field of liposome-based immune modulation.

Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00006-0

© 2020 Elsevier Inc. All rights reserved.

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6.1.1 Principles of lipid-based nanoparticles LNPs include micelles, oil-in-water emulsions, drug
lipid complexes, cochleates, and liposomes, which will be the focus of this chapter (Li et al., 2015). Liposomes are spherical vesicles composed of a single or multiple lipid bilayers that encapsulate an aqueous core. Since they were first described by Bangham and Horne in 1964, liposomes have been extensively studied for various applications including drug delivery, gene transfection, imaging, and immunizations (Daraee et al., 2016; Bangham and Horne, 1964). Their structural and functional similarity to cellular membranes also provide a platform to characterize the behavior of various biological processes including immunological reactions (Bangham et al., 1974). The primary subunits of liposomes are either natural or synthetic lipids organized into a lipid bilayer through self-assembly, often with cholesterol to help stabilize the formulation (Daraee et al., 2016). Lipids are made of a hydrophilic head group and a hydrophobic lipid tail and can form various structures depending on the three-dimensional space occupied by the head group and tails (Fig. 6.1) (Li et al., 2015). Lipids with head groups and tails of similar width (cylindrical) form bilayers and liposomal structures, while lipids with wider head groups than tails (inverted conical) form micelles, and lipids with wider tails than head groups (conical) form inverted micelles that lead to hexagonal HII structures. Importantly mixtures of lipids with different structures or characteristics (i.e., anionic and cationic pairs) can access alternative conformations when formulated together due to lipid interactions (Cullis and Hope, 2017; Li et al., 2015). The diversity of natural and synthetic lipids provides a platform for application-specific rational design. Naturally occurring head groups include phosphatidylcholine, phosphatidylethanolamine, phosphatidylserine, phosphatidic acid, phosphatidylinositol, phosphatidylglycerol, and cardiolipin, among others (Fig. 6.2) (Li et al., 2015). The chemical properties of the head groups provide liposomes with many of their clinical properties, such as immunogenicity, stability, charge, and reactivity with cell membranes (Kohli et al., 2014). The hydrocarbon tails vary in length and saturation, which influences the stability and fluidity of the liposomes, and impacts the biological activity of the formulations. For example, longer saturated lipid tails exhibit higher transition temperatures due to increased van der Waals interactions in the bilayer and increased stability of the formulation (Venditto et al., 2014b). Rational design of liposomes utilizing lipids with desirable characteristics provides a unique strategy to tailor the drug delivery properties of the nanoparticle, while providing a platform to engage with specific components of the immune system. Liposomes are characterized by various properties, including the number of lipid bilayers (lamellae), particle size, membrane fluidity, and charge. These properties, resulting from the lipids used in the formulation and the preparation technique, determine the interaction of the liposomes with biological systems, which dictates the

Modulating the immune response with liposomal delivery

Figure 6.1 Overview of lipid and liposome compositional landscape. Lipid nanoparticle characteristics and biological activity are based on the compositional makeup of the formulation. (A) Lipid structure (cylindrical, inverted conical, conical) dictates LNP architecture (bilayer, micelle, inverted micelles), respectively. Formulations composed of mixed lipid systems (e.g., cationic and anionic lipids) can alter the LNP architecture. (B) Small-molecule delivery. Aqueous soluble dugs (hexagons) are encapsulated in the interior of the formulation, while lipophilic drugs (ovals) are incorporated in the bilayer. Lipid-conjugated polymer (i.e., polyethylene glycol) and targeting moieties (i.e., antibodies) are presented on the periphery to alter the pharmacokinetic and pharmacodynamics profiles of the LNP. (C) Gene delivery. Anionic nucleic acid molecules engage with cationic lipids to form lipoplexes in multilamellar vesicles and other complex architectures. (D) Vaccine delivery. Immunogens (helical structures) and adjuvants (triangles) are combined in liposomal architectures to modulate epitope-specific immune responses. LNP, Lipid-based nanoparticle.

immunogenicity of the liposomal formulation (Daraee et al., 2016; Torchilin and Weissig, 2003). For example, the size of the formulation can determine the efficiency and immunogenicity of liposomal uptake by cells. Dendritic cells will generally take up smaller, unilamellar liposomes, while macrophages tend to take up larger particles.

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Figure 6.2 Structures of lipids commonly used in liposomal formulations. Naturally occurring phospholipids displayed with 1,2-dioleoyl-sn-glycero (DO) lipid tails. Additional forms of these lipids utilized in lipid nanoparticle formulations include saturated lipid tails of increasing carbon length, C14—dimyristoyl (DM), C16—dipalmitoyl (DP), and C18—distearoyl (DS). Cholesterol is included in many formulations to improve the fluidity of the LNPs. LNPs, Lipid-based nanoparticles.

When used as vaccines, liposomes larger than 100 nm skew the response toward cellular immunity (TH1-dependent responses), while smaller liposomes, as well as multilamellar liposomes skew the immune response toward humoral immunity (TH2-dependent responses) (Watson et al., 2012). In drug delivery, the size of the formulations also determines the pharmacokinetics of liposomal molecules, as smaller particles are able to move freely between compartments, while larger ones can be used as depots (Marasini et al., 2017). The fluidity of liposomal vesicles, which can be increased by using smaller or unsaturated lipid tails or by addition of cholesterol to

Modulating the immune response with liposomal delivery

formulations, elicits stronger responses (Watson et al., 2012). Surface charge also plays a role in immunogenicity. Cationic particles can interact more easily with cellular surfaces to induce stronger immune responses and enhance the depot effect of liposomes (Marasini et al., 2017; Christensen et al., 2011). It should be noted that bioactive lipids are also capable of activating specific immune responses, such as sphingosine-1phosphate (S1P), an extracellular signaling molecule that engages with S1P receptor to mediate vascular and immune function, and eicosanoids, which are involved in the regulation of various physiological processes (Watson et al., 2012; Alving et al., 2006; Ferguson and Nguyen, 2016; Nisini et al., 2018). However, their utility in liposomal formulations is limited due to the broad activity of various tissues in the body and will not be discussed further.

6.1.2 Principles of the immune system Modulation of immune responses using liposomal delivery relies on our understanding of the immune system and the mechanistic underpinnings associated with disease progression. The most fundamental aspect of the immune system is the ability to distinguish self from nonself. Recognition of nonself is achieved through receptor-mediated detection of molecular patterns (innate immunity) and receptor editing of cellular and secreted proteins (adaptive immunity) to recognize nonself with high affinity. All foreign substances that enter the body are capable of mounting an immune response leading to their clearance from circulation. The rational design of LNPs to harness specific immunologic pathways enhances the applicability of these systems for use in the clinical setting. Inclusion of small-molecule therapeutics, nucleic acid-based agents, and immunogens offers strategies to modulate-specific immune responses (Semple et al., 2010; Karumanchi et al., 2018; Subramanian et al., 2017; Gause et al., 2017). Initial response to pathogens is achieved via activation of pattern recognition receptors (PRRs), such as toll-like receptors (TLRs) or NOD-like receptors, on antigenpresenting cells (APCs), such as macrophages and dendritic cells. PRRs recognize pathogen-associated molecular patterns (PAMPs), which consist of proteins, lipids, glycolipids, or nucleic acids, that are characteristic of microbes (Murphy and Weaver, 2016). Activation of PRRs initiates a signaling cascade that leads to cytokine and chemokine expression resulting in a productive innate and adaptive immune response. Following activation, APCs upregulate the presentation of peptide fragments on human leukocyte antigen (HLA), also known as major histocompatibility complex (MHC) in mice. HLA presentation can lead to activation of both helper T (TH) cells and cytotoxic T lymphocytes (CTLs), which play crucial roles in adaptive immunity. While activated CTLs induce apoptosis in infected and defective (i.e., oncogenic) cells, TH cells engage with activated B cells to increase antibody production, promote affinity maturation of the antibody variable region, and antibody class switching from IgM and IgD to IgA,

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IgE, and IgG (Murphy and Weaver, 2016). The immune system also has counteracting mechanisms to inhibit responses against self-antigens and resolve immune responses once pathogens are cleared. The regulatory mechanisms established by the immune system to maintain homeostasis include cytokines (i.e., IL-10), cells [i.e., regulatory T cells (Treg)], and specific antibody subclasses (i.e., IgG4) (Murphy and Weaver, 2016). Rational design of LNPs takes advantage of the relatively immunologically inert nature of LNPs to enhance therapeutics as a means to activate stimulatory or inhibitory pathways to achieve desirable responses to treat and prevent disease. As continued advances in immune function and mechanisms of immunity are uncovered, the field of immunomodulatory LNPs will continue to evolve. This chapter will highlight advances made in the field of LNP development focused on modulating the immune system with three classes of LNP therapeutics: (1) LNPs as small-molecule drug carriers; (2) LNPs as gene delivery vehicles; and (3) LNPs as vaccines. Special attention will be given to the clinical utility of LNPs with lessons learned from currently approved LNP-based therapeutics that can guide future development of LNPs with targeted immunomodulatory properties.

6.2 Liposomal immune modulation with small-molecule therapeutics Small-molecule therapeutics are incorporated in liposomal carriers through encapsulation in the aqueous interior for water-soluble drugs, or incorporation in the bilayer for lipophilic drugs (Fig. 6.3) (Tazina et al., 2011). Preclinical evaluation of liposomal nanoparticles has generated tools for mouse models and has led to important discoveries that set a groundwork for clinical approval of a number of liposomal therapeutics (Ventola, 2017). Liposomes enhance the utility of small-molecule therapeutics by improving the solubility and bioavailability of the drug, while often reducing drugmediated toxicities (Song et al., 2012). Liposomes with small-molecule therapeutics have been approved for cancer therapy (e.g., Doxil, Marqibo), fungal disease (e.g., Amphotec, Ambisome), analgesia (e.g., DepoDur, Exparel), and photodynamic therapy (e.g., Visudyne) (Bulbake et al., 2017). Furthermore a number of small-molecule liposomal formulations are currently in clinical trials (Table 6.1). The data garnered from clinical and preclinical studies provide the basis for further investigation of additional therapeutics with immunomodulatory capacity. The principles of liposomal drug delivery have been reviewed considerably elsewhere (Tazina et al., 2011; Song et al., 2012; Barenholz, 2012) and will not be extensively discussed within the confines of this chapter. However, the utility and data garnered from specific formulations are discussed in the context of immune modulation.

Modulating the immune response with liposomal delivery

Figure 6.3 Small-molecule immune modulation drugs with potential clinical utility in liposomal formulations. Many therapeutics have been incorporated in liposomal formulations, but few have been utilized in the clinical setting. Liposomes provide a platform for continued development and evaluation of LNP-based immunomodulation with a diversity of therapeutics with well-defined target specificity. LNP, Lipid-based nanoparticle.

6.2.1 Principles of liposomal pharmacology based on Doxil Doxil was the first liposomal therapeutic approved for clinical use. It consists of doxorubicin encapsulated in a hydrogenated soy phosphatidylcholine lipid bilayer that is surrounded by a PEG corona (Barenholz, 2012). The adverse effects and pharmacokinetic profile of doxorubicin made it an ideal candidate for inclusion in a liposomal

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Table 6.1 Clinical pipeline of small-molecule therapeutics delivered using lipid nanoparticles. Agent Stage Therapeutic class Therapeutic target

Indication

References

Amphotericin B Anthralin

FDA approved FDA approved

Antifungal Anthracene

Fungal infection Psoriasis

Stone et al. (2016) Saraswat et al. (2007)

Bupivacaine Cytarabine

FDA approved FDA apporved

Opioid Antineoplastic

Pain relief Neoplastic meningitis

Dasta et al. (2012) Glantz et al. (1999)

Daunorubicin

FDA approved

Antineoplastic

Gill et al. (1996)

Doxorubicin

FDA approved

Antineoplastic

Irinotecan

FDA approved

Antineoplastic

AIDS-related karposi sarcoma Various oncologic conditions Pancreatic cancer

Mifamurtide

EMA approved phase II/III

Muramyl tripeptide

Osteosarcoma

Morphine sulfate Paclitaxel

Opioid Taxane

Poractant alfa

FDA approved Approved in China FDA approved

Tacyildiz et al. (2018) and Clinicaltrials.gov (2011a, 2018a) Gambling et al. (2005) ernabeu et al. (2017)

Pulmonary surfactant

Verteporphin

FDA approved

Photosensitizer

Vincristine

FDA approved

Antineoplastic

Iron

Phase IV

Metalic ion

Ergosterol Inhibition of cell proliferation Opioid receptor Nuecleoside antimetabolite Topoisomerase II inhibition Topoisomerase II inhibition Topoisomerase I inhibition Tumor monocytes and macrophages Opioid receptor Microtubial dysfunction Surface tension reduction in alveoli ROS production, vessel occlusion Microtubules

Pain relief Cancers

Barenholz (2012) Wang-Gillam et al. (2016)

Respiratory distress syndrome

Clinicaltrials.gov (2012b)

Choroidal neovascularization Acute lymphoblasic leukemia Chronic kidney disease, iron deficiency anemia, and inflammatory bowel disease

Group ToA-rMDWPTTS (1999) O’Brien et al. (2013) Clinicaltrials.gov (2013a, 2016a)

Lidocaine

Phase VI

Antiarrhythmic

Amikacin

Phase III

Antibiotic

Ciprofloxacin

Phase III

Fluoroquinolones

Cytarabine and daunorubicin

Phase III

Nucleotide antimetabolite and antineoplastic

T4N5 Liposomes

Phase III

Enzyme

Alendronic acid

Phase II

Bisphosphonate

All-trans retinoic acid

Phase II

Lipid-souble vitamin

Alprostadil

Phase II

Prostaglandin

Botulinum toxin

Phase II

Neurotoxic protein

Cobamamide

Phase II

Cetirizine Hydrocloride

Phase II

Active form of vitamin B-12 Antihistamine

Docetaxel

Phase II

Taxane

Class Ib sodium channel blocker Bacterial ribosome

Topoisomerase IV inhibition DNA repair inhibition and topoisomerase II inhibition DNA repair

Monocyte modulation Cell cycle arrest Blood vessel expantion Inhibits acetylcholine release Coenzyme

Local anesthetic

Clinicaltrials.gov (2012c)

Mycobacterium infection, nontuberculosis and chronic pseudomonas Noncystic fibrosis bronchiectasis High-risk acute myloid leukemia

Clinicaltrials.gov (2015a, 2011b)

Xeroderma pigmentosum, and nonmelanoma skin cancer Restenosis prevention

Clinicaltrials.gov (2004a,b)

Interstitial cystitis

Clinicaltrials.gov (2014b)

Atopic dermatitis

Clinicaltrials.gov (2012e)

Clinicaltrials.gov (2014a) Clinicaltrials.gov (2012d)

Clinicaltrials.gov (2008a, 2016b) Cancer therapeutics Clinicaltrials.gov (2003, 2004c) Arteriosclerosis obliterans Clinicaltrials.gov (2016c)

Inhibition of H-1 Allergic rhinitis receptor Microtubial Cancer therapeutics dysfunction

Clinicaltrials.gov (2007) Clinicaltrials.gov (2010a) (continued)

Table 6.1 (Continued) Agent

Stage

Therapeutic class

Therapeutic target

Indication

References

Factor VIII

Phase II

Clotting factor

Hemophilia A

Clinicaltrials.gov (2011c)

Fentanyl

Phase II

Opioid

Coenzyme of Factor IX Opioid receptors

Pain relief after ACL knee surgery Colorectal neoplasm

Clinicaltrials.gov (2008b)

Floxuridine and Irinotecan Phase II

Antineoplastic and nucleotide antimetabolite

Lurtotecan

Phase II

Mitoxantrone

Phase II

Camptothecin analogue Anthracendione

NDDP

Phase II

Platinum

9-Nitrocamptothecin

Phase II

Prednisolone

Phase II

Camptothecin analogue Glucocorticoid

SN-38

Phase II

Annamycin

Phase I/II

Active metabolite of irinotecan Anthracycline

Cisplatin

Phase I/II

Platinum

Topoisomerase I inhibition and DNA repair inhibition Topoisomerase Ovarian neoplasm inhibitor Topoisomerase II Diffuse large B cell inhibition lymphoma, peripheral T cell lymphoma and natural killer lymphoma Cross linking of Malignant mesothelioma DNA Topoisomerase Non-small-cell lung inhibitor cancer Glucocorticoid Acute relapsing remitting receptor multiple sclerosis Topoisomerase I Cancer therapeutics inhibition Topoisomerase II Cancer therapeutics inhibition Cross linking of Advanced or refractory DNA solid tumor, metastatic breast cancer, and skin cancer

Clinicaltrials.gov (2006a)

Clinicaltrials.gov (2002a) Clinicaltrials.gov (2015b)

Clinicaltrials.gov (2004d) Clinicaltrials.gov (2005a) Clinicaltrials.gov (2009a) Clinicaltrials.gov, (2005b, 2006b) Clinicaltrials.gov (2017a) Clinicaltrials.gov (2013b)

Curcumin

Phase I/II

Curcuminoid

NF-κB inhibition, antineoplastic

Cyclosporin A

Phase I/II

Immunosuppressant

Calcineurin

Dexamethasone

Phase I/II

Glucocorticoid

KRN7000

Phase I/II

Lipotecan

Phase I/II

Oxaliplatin

Phase I/II

α-Galactosylceramide Treg induction analog Camptothecin Topoisomerase analogue inhibitor Platinum Cross linking of DNA Enzyme Antioxidant

Recombinant human Cu/ Phase I/II Zn superoxide dismutase Ropivacaine Phase I/II TLK199 Phase I/II

Amino amide Peptide

Vinorelbine

Phase I/II

Vinca alkaloid

CAL02

Phase I

Antitoxin

Local anesthetic Inhibitor of Glutathione S-transferases P1-1 Microtubial dysfunction Inhibition of Streptococcus pneumonia virulence factors

Cancer therapeutics

Singh and Aggarwal (1995), Storka et al. (2015), Feng et al. (2017), and Clinicaltrials.gov (2014c) Chronic rejection in lung Clinicaltrials.gov (2012a) transplant Multiple myeloma and Clinicaltrials.gov (2017b, osteoarthritis 2018b) Allogenic bone marrow Chen et al. (2017), and Clinicaltrials.gov (2011d) transplant Liver cancer Clinicaltrials.gov (2017c) GI cancer

Clinicaltrials.gov (2009b)

Breast cancer

Clinicaltrials.gov (2012f)

Inguinal hernia Mylodisplastic syndrome

Clinicaltrials.gov (2018c) Clinicaltrials.gov (2002b)

Advanced malignancy

Clinicaltrials.gov (2016d)

Pneumonia and pneumococcal infection

Clinicaltrials.gov (2015c)

(continued)

Table 6.1 (Continued) Agent

Stage

Therapeutic class

Therapeutic target

Indication

References

E7389

Phase I

Macrocyclic ketone

Solid tumor

Clinicaltrials.gov (2017d)

Gemcitabine

Phase I

Nucleoside antimetabolite

Advanced solid tumors

Clinicaltrials.gov (2018d)

Iodixanol

Phase I

Radiocontrast agent

Computed tomography scans

Clinicaltrials.gov (2014d)

Prilocaine Rhenium-188 Topotecan

Phase I Phase I Phase I

Amino amide Radionuclide Camptothecin analogue

Tubulin interference Inhibition of DNA replication Accumulation of radiographic density Local anesthetic Radiation Topoisomerase inhibitor

Clinicaltrials.gov (2010b) Clinicaltrials.gov (2014e) Clinicaltrials.gov (2008c)

Glutathione

Not applicable

Antioxidant protein

Dental anesthesia efficacy Solid tumor Small cell lung cancer, ovarian cancer, solid tumors Low glutathione levels

Methotrexate

Not stated

Severe psoriasis

Ali et al. (2008)

Reduction of ROS Antineoplastic/ Dihydrofolate immunosuppressant reductase

FDA, Food and Drug Administration; NF-κB, nuclear factor κ B; ROS, reactive oxygen species.

Clinicaltrials.gov (2017e)

Modulating the immune response with liposomal delivery

formulation. In early phase trials, liposomal doxorubicin led to an increase in area under the curve of 609 mg/h/L compared to 1 mg/h/L when administered at an equivalent dose of free drug (Gabizon et al., 1994). In the first successful Doxil trial for treatment of AIDS-related Kaposi sarcoma, Doxil administration resulted in a 5- to 11-fold increase in tumor doxorubicin levels compared with free doxorubicin. It also led to significantly reduced toxicity and patient tolerance (Northfelt et al., 1996). These results translated to greater clinical efficacy compared to the standard of care at the time, which consisted of a combination of free doxorubicin, bleomycin, and vincristine (Northfelt et al., 1998). In later trials with patients diagnosed with breast, ovarian and lung cancers, liposomal doxorubicin also demonstrated increased drug accumulation within solid tumors compared with free doxorubicin (Gabizon et al., 1994). The inclusion of PEG in Doxil provides an intriguing paradigm for immune suppression. PEG prevents liposomal recognition by the reticuloendothelial system, which extends drug circulation half-life of the drug from hours to days. However, humans and laboratory animals are able to mount an antibody response toward PEG, which results in opsonization and accelerated blood clearance of subsequent doses of liposomes (Kierstead et al., 2015; Dams et al., 2000; Ishida et al., 2006). Intriguingly PEGylated liposomes containing doxorubicin fail to induce a productive immune response due to the cytotoxic/immunosuppressive effects of doxorubicin (Laverman et al., 2001). This strategy for immunosuppression has been explored using ovalbumin as an antigen, with successful inhibition of antibodies to the immunogenic protein (Oja et al., 2000; Tardi et al., 1997), as well as methotrexate as the therapeutic (Shek et al., 1986). Methotrexate alone exhibits cytotoxic and immunosuppressive properties and is administered to patients with cancer and autoimmunity. Liposomal formulations of methotrexate have been used in patients to suppress immune responses in patients with severe psoriasis (Watson et al., 2012; Torchilin and Weissig, 2003; Marasini et al., 2017). Similar to the Doxil formulation, methotrexate formulations administered to laboratory animals in conjunction with bovine serum albumin suppress the proteinspecific responses (Shek et al., 1986). However, a curious dose-dependent response was observed with low-dose methotrexate formulations activating an immune response (Shek et al., 1986). Altogether Doxil and other liposomes containing cytotoxic therapeutics established a foundation for clinical utility of liposomal therapeutics as well as a paradigm to modulate targeted immune responses in patients and laboratory animals.

6.2.2 Immune modulation using small-molecule therapeutics Prior to the approval of Doxil and exploration of the associated immunosuppressive therapeutics, dichloromethylene bisphosphonate (clodronate) was prepared in a

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liposomal formulation to eliminate phagocytes from the spleen (van Rooijen and van Nieuwmegen, 1984). Clodronate is a bisphosphonate that inhibits bone resorption and can be used to treat osteoporosis (Ghinoi and Brandi, 2002). Formulations containing clodronate, and lacking PEG-lipid, are rapidly engulfed by phagocytes leading to cell death when intracellular clodronate concentrations reach cytotoxic levels (van Rooijen and van Nieuwmegen, 1984; Van Rooijen and Sanders, 1994). The selective elimination of phagocytic cells provides a tool for use in laboratory animals with potential utility in clinical applications. Clodronate liposomes were shown to reduce the severity of thrombocytopenic purpera (Alves-Rosa et al., 2000), autoimmune hemolytic anemia (Jordan et al., 2003), and arthritis in mouse and rat models (Richards et al., 1999, 2001). While permanent depletion of macrophages in chronic autoimmune disease would likely lead to adverse effects, including infections, the ability of short-term depletion in an acute autoimmune disease may prove advantageous. Studies of clodronate liposomes in a mouse model of hemolytic anemia provide evidence that liposomes suppress clearance of opsonized red blood cells when given simultaneously but had no effect 2 weeks after administration, demonstrating the temporary nature of this effect (Jordan et al., 2003). The rapid onset of action and limited duration indicates that liposomal clodronate can provide an effect similar to a splenectomy on a temporary basis while avoiding the lifelong consequences of spleen removal (Jordan et al., 2003). Targeting of clodronate containing liposomes was explored by the addition of mannose to the liposomes, which enabled passage of the liposomes across the blood
brain barrier and uptake of these liposomes within glial cells (Umezawa and Eto, 1988). Mannosylation of clodronate liposomes led to decreased severity of symptoms in a experimental autoimmune encephalomyelitis (EAE) rat model of multiple sclerosis through targeted reduction of macrophages (Huitinga et al., 1990). Liposome targeting utilizing carbohydrate, small molecules, peptides, and antibodies has been explored extensively in the context of drug delivery for tumors (Deshpande et al., 2013). Targeted delivery can also be achieved by altering the route of administration of the liposomal therapeutic (Allen et al., 1993). Most liposomal therapeutics have been investigated using intravenous or intraperitoneal injection. However, inhalation of aerosolized liposomes using a nebulizer provides a strategy to deliver therapeutics directly into the lung (Waldrep et al., 1993). This strategy has been implemented with liposomes containing cyclosporine A administered to patients following lung transplant (Behr et al., 2009). Free cyclosporine A is currently prescribed as a component of the posttransplant cocktail to suppress the phosphatase activity of calcineurin and suppress lymphocyte activity (Matsuda and Koyasu, 2000). Behr and colleagues found that 40% of a 10 mg dose of inhaled liposomal cyclosporine A was deposited in the lung (Behr et al., 2009). A clinical trial of 21 randomized patients demonstrated that treatment with aerosolized liposomal cyclosporine in combination with conventional oral

Modulating the immune response with liposomal delivery

immune suppression was superior to conventional oral immune suppression with a primary endpoint of chronic rejection with bronchiolitis obliterans, retransplant, or death (Clinicaltrials.gov, 2012a). Glucocorticoids are a class of immunosuppressive therapeutics that play an essential role in the management of various diseases from transplant to rheumatic diseases. Unlike cytotoxic agents, such as doxorubicin, methotrexate, and clodronate, glucocorticoids suppress immune responses through transactivation of the glucocorticoid receptor to induce transcription of antiinflammatory target genes and suppression of proinflammatory cell effects (van der Velden, 1998). Due to the utility of glucocorticoids in patients with multiple sclerosis and other autoimmune disorders glucocorticoid liposomes were tested in animal models of autoimmune disease. PEGylated liposomal formulations of prednisolone showed selective accumulation in the inflamed central nervous system (CNS) of rats with EAE (Schmidt et al., 2003). In an adaptive transfer model of EAE in rats, liposomal prednisolone achieved greater reduction of disease pathology compared to free methylprednisolone at 20% the concentration of the free drug (Schmidt et al., 2003). The activity of liposomal glucocorticoids has also been replicated in models of arthritis (Metselaar et al., 2003, 2004; Avnir et al., 2007), and more recently shown to be superior to free drug in a model of adjuvant-induced arthritis using a non-PEGylated formulation of dexamethasone (Anderson et al., 2010). The preclinical work with liposomal glucocorticoids has advanced into clinical trials using Nanocort, a PEGylated formulation of prednisolone, developed by Galapagos NV. This agent has been tested in several phase I and phase II trials in patients with multiple sclerosis (Clinicaltrials.gov, 2009a) and ulcerative colitis (van Assche et al., 2016).

6.2.3 Future directions in liposomal immune modulation Inclusion of well-established therapeutics, such as glucocorticoids, or novel therapeutics into liposomal formulations that target immunological signaling pathways provides a clear advantage over broad acting agents. Clinically approved immunomodulatory agents available as liposomal formulations include anthralin for psoriasis (Saraswat et al., 2007) and α-galactosylceramide KRN7000 for graft-versus-host disease (Chen et al., 2017). Additionally, the tetracyclic antibiotic minocycline is currently being evaluated in a preclinical EAE model (Hu et al., 2009). Additional therapeutics capable of suppressing immune responses have been identified, including mammalian target of rapamycin (mTOR) (Dumont and Su, 1996), ibrutinib [bruton protein-tyrosine kinase (BTK)] (Roskoski, 2016), fenebrutinib (BTK) (Crawford et al., 2018), PRT062607 (Syk) (Spurgeon et al., 2013), curcumin (NF-κB) (Singh and Aggarwal, 1995), and tofacitinib (JAK3) (Meyer et al., 2010). Inclusion of these agents and other novel therapeutics or prodrugs of the active molecules warrant exploration for their ability to modulate immune responses. For

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example, NF-κB inhibitors including curcumin, quercetin, and Bay11-7082 are lipophilic molecules that have been incorporated into liposomes and administered to mice expressing an ovalbumin-specific T cell receptor (Capini et al., 2009). After immunization with ovalbumin in complete Freund’s adjuvant, splenocytes from mice that received NF-κB inhibitors exhibited decreased T cell proliferation in vitro compared with untreated controls. In a methylated bovine serum albumin (mBSA)-induced inflammatory arthritis model, formulations of mBSA with NF-κB inhibitors suppressed the development of arthritis as compared to controls lacking either mBSA or NF-κB inhibitor (Capini et al., 2009). Other currently approved therapeutics, including the macrolide antibiotic azithromycin, also warrant investigation in liposomes (Oh et al., 1995). In addition to its antimicrobial activity, there is evidence of the immunosuppressive activity of azithromycin through inhibition of p65 translocation and NF-κB signaling (Murphy et al., 2008; Al-Darraji et al., 2018; Feola et al., 2010). The immunosuppressive activity of free azithromycin has been investigated in the context of lung infection (Feola et al., 2010), myocardial infarction (Al-Darraji et al., 2018), and spinal cord injury (Zhang et al., 2015) for its ability to induce alternatively activated macrophages that limit inflammation induced lesion volume in these models. Development of liposomal formulations containing azithromycin and other approved immunosuppressive therapeutics may provide further opportunities to selectively modulate immune responses by harnessing the benefits associated with liposomal delivery of these drugs.

6.3 Liposomal immune modulation with liposomal gene vectors The objective of gene therapy is to introduce missing or faulty gene products or remove/decrease faulty genes directly or through RNA interference (RNAi) (Riley and Vermerris, 2017). LNPs are currently among the leading nonviral gene delivery system (Chen et al., 2016). In August 2018 after demonstrating efficacy in phase III clinical trials, the Food and Drug Administration (FDA)-approved Patisiran, an LNPbased therapy for the treatment of transthyretin induced amyloidosis (Table 6.2) (Wood, 2018). Patisiran delivers double-stranded small-interfering RNA (siRNA) targeting transthyretin messenger RNA (mRNA) to prevent translation, ultimately reducing levels of transthyretin protein in the liver and the resultant amyloid deposits in tissues (Adams et al., 2018). Numerous other LNPs have been studied as gene therapy vectors for the treatment of cancer, hepatitis, atherosclerosis, and other systemic conditions (Cullis and Hope, 2017; Yin et al., 2014; Kanasty et al., 2013; Zhao and Huang, 2014; Chakraborty et al., 2017). Currently there are four LNPs in clinical

Table 6.2 Clinical pipeline of nucleic acid therapeutics delivered using lipid nanoparticles. Agent Stage Therapeutic Therapeutic target Indication class

References

Patisiran

Adams et al. (2018)

ARB-1467

FDA siRNA approved Phase II siRNA

Transthyretin

Transthyretin-related hereditary amyloidosis Hepatitis B infection

Hepatitis B virus genome

DCR-MYC

Phase II

siRNA

c-Myc oncogene

ND-L02s0201

Phase II

siRNA

Heat shock protein 47

TKM-100802 Phase II

siRNA

Prexigebersen Phase I/II (BP 1001)

ASO

Zaire ebolavirus L. polymerase and viral protein 35 Growth factor receptor bound protein 2

TKM 080301 Phase I/II

siRNA

Polo-like kinase 1

ALN-PCS02

Phase I

siRNA

ALN-VSP02

Phase I

siRNA

Atu027 DCR-PH1

Phase I Phase I

siRNA siRNA

Proprotein convertase subtilisin/kexin type 9 Kinesin spindle protein Solid tumors and vascular endothelial growth factor Protein kinase N3 Advanced solid tumors Hydroxyacid oxidase 1 Primary hyperoxaluria type 1

Hepatocellular carcinoma and various oncologic conditions Idiopathic pulmonary fibrosis Ebola virus

Acute myeloid leukemia and chronic myelogenous leukemia Liver cancer and adrenocortical carcinoma Hypercholesterolemia

Mani et al. (2018), and Clinicaltrials.gov (2015d) Kaczmarek et al. (2017)

Bansal et al. (2016), Zabludoff et al. (2017), and Clinicaltrials. gov (2018e) Dunning et al. (2016)

Ohanian et al. (2018), and Clinicaltrials.gov (2016e,f) Demeure et al. (2016), and Titzede-Almeida et al. (2017) Fitzgerald et al. (2014) Xu and Wang (2015)

Schultheis et al. (2014) Kaczmarek et al. (2017) (continued)

Table 6.2 (Continued) Agent Stage

EPHARNA

Phase I

Therapeutic Therapeutic target class

Indication

References

siRNA

Ephrin type A receptor 2

Various oncologic conditions

ASO

K-Ras and MAPK signaling Tumor-associated antigenpositive cells Apolipoprotein B Zaire ebolavirus viral proteins 24 and 35 Tumor-associated antigenpositive cells

Advanced cancers

Wagner et al. (2017), Naing et al. (2017), and Clinicaltrials.gov (2012g, 2017f) Steinberg et al. (2005)

LErafAONPhase I ETU Lipo-MERIT Phase I

MERIT

PRO-040201 Phase I TKM-100201 Phase I

siRNA siRNA

TNBCMERIT

MERIT

Phase I

Malignant melanoma Hyperlipidemia Ebola virus

Jabulowsky et al. (2018), and Clinicaltrials.gov (2015e) Xu and Wang (2015) Xu and Wang (2015)

Triple-negative breast cancer

Frenzel et al. (2017), and Clinicaltrials.gov (2014f)

ASO, Antisense oligonucleotide; MERIT, mutagenome engineered RNA immune therapy; siRNA, small-interfering RNA.

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trials, and as the field continues to evolve, additional LNPs will undoubtedly progress into the clinic. The principles of liposomal gene delivery used in these clinically relevant formulations can be readily translated to immune targets to alter the pro- and antiinflammatory profiles associated with disease progression.

6.3.1 Principles of liposomal gene delivery There are two approaches to alter genetic expression: (1) introduction of functional genes in patients who have faulty gene production or a protein deficiency, usually in the form of a plasmid or mRNA and (2) gene interference in patients with dysregulated gene expression using microRNA (miRNA), short hairpin RNA (shRNA), siRNA, and antisense oligonucleotides (ASOs). miRNA, shRNA, and siRNA function through RNAi by inhibiting translation of their target mRNA when loaded onto RNA-induced silencing complex (RISC) (Hannon, 2002; Wilson and Doudna, 2013). ASOs bind directly to complementary mRNA sequences and promote mRNA degradation by RNase H (Riley and Vermerris, 2017; Kaczmarek et al., 2017). The simplest strategy for achieving gene therapy is to introduce genes systemically, with the goal that they will be taken up by cells in vivo, similar to the success achieved in vitro. Although this strategy has yielded promising results in vitro, naked nucleic acids are usually degraded in circulation and ultimately yield poor outcomes when delivered into live organisms (Kaczmarek et al., 2017; Lv et al., 2006; Garber, 2017). As a result, viral and nonviral vectors have been used to enhance delivery of nucleic acids into animals. Viral gene delivery uses the natural strategies of viruses to insert specific genes introduced into the viral genome. Viral particles like adenoassociated virus have acceptable efficacy and safety parameters but also have significant limitations (Zhang et al., 2004; Zabner, 1997). One of the major difficulties with viral gene vectors is that they are difficult to manufacture. In addition, because they utilize viral vectors for delivery they can result in systemic infections. Additionally previous exposure to the viral vector can result in significant immune response against viruses and hinder their ability to transfect cells (Lv et al., 2006). Nonviral gene delivery is a relatively broad term to describe delivery using various strategies to improve nucleic acid delivery. Among these are the use of vectors including LNPs, peptides, polymers, dendrimers, and other nanoparticle strategies (Al-Dosari and Gao, 2009; Ramamoorth and Narvekar, 2015). LNPs interact with anionic phosphates in nucleic acids via cationic amines on the nanoparticle surface. Liposome/ nucleic acid complexes (lipoplexes) readily form by incubating the liposomes with nucleic acids, allowing them to self-assemble. Nucleic acids can also be encapsulated in the interior of a liposome to emulate viral particles by hydrating lipids in a nucleic acid solution. However, encapsulation requires high amounts of nucleic acids to achieve adequate nucleic acid concentrations inside the liposomes. Some cationic lipids

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(Fig. 6.4) that initially demonstrated strong interactions with nucleic acids proved useful for in vitro transfections but became limited in vivo due to large particle size ( . 100 nm), instability, positive surface charge, and dose-limiting toxicities (Cullis and Hope, 2017; Zhang et al., 2004; Candiani, 2016). Ionizable cationic lipids, such as 1,2-dioleoyl-3-dimethylammonium propane, along with the ethanol loading procedure, improved the viability of LNPs, yielding higher loading efficiency and smaller vesicles (Cullis and Hope, 2017). In 2005 a novel lipoplex preparation method called the T-tube mixing technique was described (Jeffs et al., 2005). In this procedure, lipids are dissolved in ethanol and rapidly mixed with nucleic acids in a low-pH buffer to achieve a small diameter nanoparticle of ,100 nm. This technique was enhanced in 2012 to allow for more rapid and controllable mixing, achieving more uniform systems (Cullis and Hope, 2017). This method results in complex structures that have a hydrophilic core made of inverted micelles of lipid encapsulating nucleic acids

Figure 6.4 Cationic lipids for gene delivery. Cationic lipids used for gene delivery include six general classes of molecules. Examples of each class are shown above: monovalent lipids (DOTMA, DOTAP), multivalent lipids (DOGS, DOSPA), guanidinium-containing lipids (UGG), ionizable lipids (DODMA, DLinDMA, DLinMC3DMA), gemini surfactants (C18-3-18), and cholesterol analogues (DC-cholesterol). DC-cholesterol, 3β-[N-(Nʹ,Nʹ-Dimethylaminoethane)-carbamoyl]cholesterol hydrochloride; DLinDMA, 1,2-dilinoleyloxy-3-dimethylaminopropane; DLinMC3DMA, dilinoleylmethyl4-dimethylaminobutyrate; DODMA, 1,2-dioleyloxy-3-dimethylaminopropane; DOGS, dioctadecylamidoglycylspermine; DOSPA, 2,3-dioleyloxy-N-[2(sperminecarboxamido) ethyl]-N,N-dimethyl-1-propanaminium; DOTAP, 1,2-dioleoyl-3-trimethylammonium-propane chloride; DOTMA, 1,2-di-Ooctadecenyl-3-trimethylammonium propane chloride; UGG, unsaturated guanidinium glycoside.

Modulating the immune response with liposomal delivery

surrounded by a coating of lipids (Cullis and Hope, 2017). Other strategies to enhance nanoparticle delivery of gene targeting molecules have focused on altering the nucleic acid molecules. For example, Geall et al. demonstrated the capability of selfamplifying mRNA, based on in vivo viral replication, to overcome the limitations of vector molecules by increasing the bioavailability of mRNA within cells (Ulmer and Geall, 2016). Various groups have also focused on applying chemical modifications to siRNA to improve the activity, half-life and specificity of these molecules (Song et al., 2017; Deleavey et al., 2009). In addition to size, the charge of the cationic molecules plays a essential role in their efficacy. The ratio of cationic lipid amines to nucleic acid phosphates (N/P ratio) is crucial for liposomal gene delivery as it determines how tightly bound nucleic acids are to the liposome, as well as the ability to adhere to cellular surfaces (Woodle and Scaria, 2001). A common procedure employed to assess the binding capacity of lipoplex systems includes testing various N/P ratios on an electrophoresis gel retardation assay (Candiani, 2016; Hughes et al., 2010). In addition, the type and amount of PEGylated lipids used creates a difference in the LNP size and influences the potency of the system. In general addition of PEGylated lipids results in larger LNPs and can result in a reduced capacity for gene delivery, especially with long-lived PEGylated molecules that reduce the ability of LNPs to adhere to cells or be taken up by hepatocytes. Optimal gene delivery efficiency with siRNA has been observed with PEGylated lipids containing short (C14) acyl chains that can dissociate from the liposomal structure in vivo with a half time of ,30 minutes (Cullis and Hope, 2017).

6.3.2 In vitro liposomal gene delivery Lipoplexes are taken into cells through various mechanisms, including macropinocytosis, phagocytosis, receptor-mediated endocytosis (in the case of liposomes with targeting molecules), and clathrin-mediated and caveolae-mediated endocytosis (Zabner, 1997; Huth et al., 2006; Allen and Cullis, 2013; Heath et al., 1985; Tandrup Schmidt et al., 2016). Additionally studies have shown that some LNPs may also enter cells through fusion with the cell cytoplasm (Zabner, 1997; Huth et al., 2006). While clathrin- and caveolae-mediated endocytosis are the primary mechanisms of lipoplex uptake, other mechanisms can vary between different cell types, which can affect the nature and efficacy of the delivery system (Huth et al., 2006; Wasungu and Hoekstra, 2006). Following endocytosis, nucleic acids can be degraded once the endosome fuses with a lysosome, especially in the case of lipoplexes formed through incubation of preformed liposomes and nucleic acids. However, compartmentalization of nucleic acids within lipoplexes can help protect nucleic acids from degradation (Cullis and Hope, 2017). A notable characteristic of lipoplexes is the formation of complex multilamellar and inverted micellar structures, known as hexagonal HII phase structures, around the

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nucleic acid (Fig. 6.1) (Zabner, 1997; Dan and Danino, 2014; Koynova, 2010; Gershon et al., 1993). The hexagonal HII phase structure protects the nucleic acids from degradation within the lysosome. It is also hypothesized that the formation of these structures allows for interactions between LNPs and cellular membrane lipids, resulting in disruptive nonlamellar structures that allow nucleic acids to escape into the cytoplasm where they can interact with effector proteins such as RISC, or be transported to their final destinations in the nucleus or rough endoplasmic reticulum (Cullis and Hope, 2017; Semple et al., 2010). Once inside the lysosome, the lipoplex must escape into the cytosol to execute its function. A proposed mechanism for its escape includes interaction of the endosomal lipid phosphatidylserine in the outer leaflet of the lysosome with the cationic lipoplex (Zabner, 1997). As phosphatidylserine flips to the inner leaflet of the endosome, it coordinates with the cationic lipoplex and orchestrates its escape from the lysosome (Zabner, 1997). Phosphatidylserine may also stimulate hexagonal phase formation within the endosome to protect nucleic acids from degradation (Zabner, 1997; Wasungu and Hoekstra, 2006). Escape of lipoplexes from endosomes can also be aided by dioleoylphosphatidylethanolamine (DOPE), which can destabilize lysosomal lipid bilayers, or by addition of adenoviral protein to the lipoplex, which is believed to disrupt the endosomal structure by interacting with cellular integrins (Zabner, 1997; Meunier-Durmort et al., 1997; Nemerow and Stewart, 2016). Based on this mechanism, effective nucleic acid release into the cytosol requires a tight intermembrane interaction between the lipoplex surface and endosomal membrane. Therefore PEGylated lipoplexes benefit from the use of exchangeable PEG-lipid analogs or cleavable pH-sensitive PEG analogs (Wasungu and Hoekstra, 2006). The activity of interfering RNAs and ASOs occurs once they escape the endosome and enter the cytoplasm to engage with their cytoplasmic targets (Hannon, 2002; Leonhardt et al., 2014), while plasmid DNA must enter the nucleus to undergo transcription (Dean et al., 2005). Nuclear migration of plasmid DNA from the cytoplasm to the nucleus after escaping the lysosome is necessary for appropriate transcription and has been described (Cohen et al., 2009; Saffari et al., 2016). The leading theory suggests that migration of DNA into the nucleus occurs through passive diffusion, especially in the case of smaller molecules (less than B300 base pairs) (Saffari et al., 2016; Coppola et al., 2012). Entry of larger molecules into the nucleus can be aided by modifications in the nucleic acid that assist in homing to the nucleus (Zhang et al., 2004; Zabner, 1997; Saffari et al., 2016; Coppola et al., 2012; van Gaal et al., 2011).

6.3.3 In vivo liposomal gene delivery The complexity of cellular uptake and intracellular delivery is further complicated using in vivo systems, which must navigate immune responses, protein adsorption, and

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biological processes directed toward the formulation. The propensity of LNPs to associate with lipid trafficking proteins, such as apolipoprotein E (ApoE), causes LNPs to be targeted to hepatocytes (Cullis and Hope, 2017). ApoE also plays a central role in delivery to neuronal tissues (Cullis and Hope, 2017). While circulating cholesterol cannot cross the blood
brain barrier, ApoE is capable of transporting brain-derived cholesterol to neurons. When administered intracortically, LNPs can be transported into neuronal tissues in the same manner as they do in the liver (Cullis and Hope, 2017). Another major target of LNPs are the phagocytic cells of the immune system, suggesting the potential therapeutic utility of LNP therapies in inflammatory and immune-mediated diseases (Cullis and Hope, 2017; Claassen, 1992). Targeting of specific tissues and/or cell types can be achieved using antibodies or other moieties on the liposomal surface. For example, in one study, cyclic arginine-glycine-aspartic acid was prepared in a formulation composed of DOPE, dipalmitoylphosphatidylcholine (DPPC), cholesterol, and PEGylated distearoylphosphoethanolamine (DSPE) complexed with siRNA, to target αVβ3 integrin on A549 lung cancer cells, which resulted in more effective inhibition of gene expression than non-targeted liposomes (Byrne et al., 2008). Targeted delivery of lipoplexes has also been achieved using Listeriolysin O to target HER2 expressing cells using the luciferase reporter gene on a DNA plasmid (Kullberg et al., 2014). Despite their relative safety, a major downside to using cationic lipids is their ability to damage cells by rupturing the anionic cell membrane. In order to reduce their cytotoxicity, cationic lipids are often mixed with helper lipids, such as cholesterol and other natural lipids, which also increase the stability of these formulations (Hafez et al., 2001). The damage resulting from cellular lysis during LNP gene delivery can ultimately lead to activation of the immune system upon recognition of cellular damage (Lv et al., 2006). To attenuate immune reactions related to cell lysis and to inhibit the removal of LNPs from circulation, immunosuppressants like dexamethasone are often used in the clinical setting. This strategy has been shown to improve the efficacy of these therapies in clinical settings (Cullis and Hope, 2017; Zabner, 1997).

6.3.4 Immune modulation using liposomal gene vectors Immune modulation with gene delivery vectors provides a platform to induce or suppress expression of immune mediators or proteins involved in pro and antiinflammatory pathways enabling therapeutic discovery in multiple disease states. One example of such a platform involves suppression of tumor necrosis factor α (TNFα), a major cytokine involved in various inflammatory processes. Anti-TNFα siRNA was investigated in the setting of LPS-induced sepsis in mouse models using DOTMA-based LNPs (Sørensen et al., 2003), and in the setting of inflammatory bowel disease (Zhang et al., 2006). In both instances, the downregulation of TNFα resulted in improved

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outcomes in mice. More recently Aldayel et al. (2018) published a study demonstrating the ability of DOTAP-based LNPs to deliver TNFα siRNA to reduce inflammation and disease symptoms in mouse models of collagen-induced arthritis and in methotrexate resistant, anticollagen-induced arthritis. Strategies to modulate immune responses by targeting cell cycle proteins, which are implicated in inflammatory immune responses, have also been employed. Cyclin D1 is a regulatory molecule involved in the proliferation of lymphocytes during inflammation. In 2008, Peer et al. developed a neutral liposome formulation containing anticyclin D1 siRNA with a guiding monoclonal antibody to target β7 integrin in mouse models of colitis. The reduction of cyclin D1 resulted in significant reduction in inflammatory damage in the colonic tissue of mice (Peer et al., 2008). Cyclin D1 has also been the target of studies in ex vivo human tissues by Piazza et al., which demonstrated marked reduction of cyclin D1 in various tissue cultures (Chen et al., 2016; Piazza et al., 2017; Xue et al., 2015). More recently using an innovative approach, Katakowski et al. (2016) delivered siRNA targeting the costimulatory molecules CD40, CD80, and CD86 in dendritic cells to inhibit T and B cell activations. The authors took advantage of single chain antibodies against dendritic cell marker DEC205 to guide the therapy. Their results demonstrate the ability of a DLinMC3DMA-based LNP formulation (Fig. 6.4) to downregulate specific costimulatory molecules and provides a mechanism to tailor immune responses to inhibit mixed lymphocyte responses (Katakowski et al., 2016). The immunomodulatory potential of miRNAs has been described in multiple disease states from cancer to cardiovascular to neuronal disease (Chakraborty et al., 2017; Feinberg and Moore, 2016; Tessitore et al., 2016; Wang et al., 2017; Wang and Springer, 2015). Eulalio et al. (2012) identified two miRNA molecules, hsa-miR199a-3p and hsa-miR590-3p, which were introduced with AAV vectors and shown to promote cell cycle reentry of adult cardiomyocytes ex vivo and stimulate cardiac regeneration following myocardial infarction. Lipofectamine RNAiMAX was used by Lesizza et al. to deliver hsa-miR-199a-3p and hsa-miR590-3p to mice following myocardial infarction (Lesizza et al., 2017). Both miRNAs demonstrated improved recovery from myocardial infarction compared with untreated controls (Lesizza et al., 2017). miR-210 is associated with repression of mitochondrial metabolism and attenuation of keratinocyte proliferation, leading to decreased wound healing (Ghatak et al., 2016). Ghatak et al. (2016) tested antihypoxamiR functionalized gramicidin lipid nanoparticles in a mouse model of ischemic injury to deliver antimiR-210. The improvement in wound closure observed in their study demonstrates the significance of this miRNA target in wound healing. LNPs have also been studied for DNA delivery in the context of immunotherapy. Dinh et al. (2011) demonstrated inhibition of osteoclastogenesis and proinflammatory cytokine production in primary culture macrophages using NF-κB oligonucleotide

Modulating the immune response with liposomal delivery

decoys complexed with mannose or galactosyl liposomes. A similar approach was studied by Wijagkanalan et al. (2011) in the setting of lung inflammation. The authors found reduced levels of TNFα, IL-1β, CINC-1, and decreased neutrophil infiltration and NF-κB expression in rats treated with this LNP-based NF-κB oligonucleotide decoy therapy. A major concern following solid organ transplantation is ischemia reperfusion injury, where cellular damage to the transplanted organ is accentuated following organ reperfusion due to metabolic stress and inflammation (Chen and Date, 2015; Salvadori et al., 2015; Guan et al., 2014). Production of nitric oxide from endothelial cells has been found to be protective in this setting by increasing organ reperfusion and inhibiting expression of cell adhesion molecules and proinflammatory cytokines through induction and activation of IκBα, which blocks transcription of the proinflammatory transcription factor NF-κB (Chen and Date, 2015; Guan et al., 2014; Iwata et al., 2001). In 2001, Iwata et al. utilized an endothelial nitric oxide synthase (eNOS) DNA plasmid delivered with cationic LNPs and demonstrated that eNOS upregulation led to reduced ischemic damage to transplant allografts in rats, secondary to NF-κB inhibition and reduction of leukocyte infiltration to the graft (Iwata et al., 2001). The antiinflammatory cytokine IL-10 plays a crucial role in many inflammatory conditions. Pascual et al. used an elastic liposome formulation containing DOTAP to deliver an IL-10 plasmid into human corneal epithelial cells and enucleated rabbit eyes (Vicente-Pascual et al., 2018). Multiple studies have demonstrated the capacity of IL10 and IL-4 delivery to reduce inflammation, improve function, and prolong survival of cardiac allografts in cellular and animal models of cardiac transplant (Furukawa et al., 2005, 2008; Tung et al., 2003; Sen et al., 2001; Oshima et al., 2002). IL-10 was also shown to improve outcomes in rat models of liver transplant using DOTAPbased liposome formulations (Fabrega et al., 1996). Another target for regulating inflammatory conditions was studied by Duguay et al. who demonstrated the ability of LNPs formulated with ionizable cationic lipids using microfluidic mixing to transfect mast cell lines with green fluorescent protein. Mast cells, which are often refractory to LNP gene delivery using conventional approaches, are crucial to the development of various allergic reactions. These results provide a novel target for studying LNPs in the setting of allergic and rheumatic conditions (Duguay et al., 2018).

6.3.5 Future directions for liposomal gene vectors As ongoing and future research delineates the molecular mechanism of different disease processes, LNP gene therapy vectors are a promising treatment option. Unfortunately, there are challenges that face broad utilization of LNPs. Much of the research up until this point has centered primarily on the delivery of siRNA, continued efforts will need to focus on optimizing lipid particles for the delivery of alternate therapeutic

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molecules such as miRNA and DNA plasmids. Additionally, the development and evaluation of miRNA inhibitors, such as antiinflammatory miR-155, which have been shown to act on silencing genes to regulate expression (Wang et al., 2015). An additional challenge to the use of LNPs is their propensity to be taken up by liver and phagocytic cells. As the field continues to advance, novel methods such as the addition of targeting antibodies will need to be studied to improve delivery to specific cells of interest (Woodle and Scaria, 2001). Furthermore the delivery of cationic molecules has been linked to inflammation in clinical studies, partly due to lysis of cells by cationic lipids, which releases various inflammatory molecules, such as IL-6 (Ruiz et al., 2001; Zabner et al., 1997; Scheule et al., 1997). Cationic nanoparticles can also activate the alternative pathway of the complement system through electrostatic interactions with complement proteins, leading to accelerated clearance (Szebeni et al., 2011). Some studies have attempted to overcome this through PEGylation or reduction of the N:P ratio on nanoparticle surfaces with varied effect (Filion and Phillips, 1997; Szebeni and Moghimi, 2009). Broering et al. (2014) showed that 20 -methylation of siRNA could help LNPs avoid PRR-mediated immune activation, implying that posttranslational modifications on nucleic acids may reduce the immunogenicity of lipoplexes. Unfortunately some of these solutions have been shown to reduce the transfection efficacy of LNPs, which warrants the investigation of novel methods to reduce inflammatory responses to LNP-mediated gene vector delivery. The field of gene delivery still lacks concrete understanding of the physiological pathways that underlie LNP-based gene delivery. While mechanisms of delivery have been described, a critical area of continued focus and discovery is how lipoplexes deliver their payload into cells and how inclusion of different lipid components affect the conformation of lipoplexes to achieve optimal delivery into cells and efficacious immune modulation (Tiffany and Szoka, 2016).

6.4 Immune stimulation with liposomal vaccines Liposomal vaccines targeting diphtheria toxin and mycobacterium were first reported in 1974 (Allison and Gregoriadis, 1974). Since this initial report, multiple studies on the capabilities of liposomes in this context have enabled the field to grow in the clinical setting (Schwendener, 2014). These efforts have resulted in two commercially approved virosomal vaccines against influenza and hepatitis A, as well as several others currently in clinical trials (Table 6.3) (Schwendener, 2014; De Serrano and Burkhart, 2017; Alving et al., 2016). In addition to viruses, liposomal vaccines have

Table 6.3 Clinical pipeline of LNP-based vaccines. Agent Stage Therapeutic class

Therapeutic target

Indication

Epaxal

EMA approved

Virosome

Hepatitis A virus

Hepatitis A

Inflexal V

EMA approved

Virosome

RTS.S/AS01

Phase III

MPL and QS21

Tecemotide

Phase III terminated

MPLA

dHER2

Phase II

AS15

GSK692342

Phase II

M72 and AS01/AS02

JVRS-100

Phase II

Noncoding plasmid RNA Plasmid DNA

Human simplex virus type 2 proteins

Herpex simplex virus type 2

Shlapobersky et al. (2012)

Influenza virus proteins HIV glycoprotein 120

Influenza HIV virus

Watson et al. (2012) Clinicaltrials.gov (2017h)

VCL-HB01/ Phase II HCLHM01 Vaxisome Phase II HIV clade C Phase I/II Env protein

Virosome MF59 and AS01

References

Lim et al. (2014) and Chappuis et al. (2017) Hemagglutinin and Influenza Herzog et al. (2009), Gasparini et al. (2013), neuraminidase Conne et al. (1997) Plasmodium falciparum Plasmodium falciparum and Watson et al. (2012), Mahmoudi and circumsporozoite plasmodium vivax Keshavarz (2017), and protein and hepatitis Clinicaltrials.gov B surface antigen (2017g) Mucin 1 lipopeptide Multiple myeloma, non- Watson et al. (2012) BPL25 small-cell lung cancer, prostate, rectal, and breast cancer Receptor tyrosineReceptor tyrosineHamilton et al. (2012) protein kinase erbB-2 protein kinase erbB-2 positive breast cancer Mycobacterium tuberculosis Mycobacterium tuberculosis Montoya et al. (2013) and Clinicaltrials.gov (2006c) Influenza virus proteins Influenza Watson et al. (2012)

(continued)

Table 6.3 (Continued) Agent Stage

Therapeutic class

Therapeutic target

Indication

References

RH5.1/ AS01 Phase I/II

AS01

Plasmodium falciparum

Clinicaltrials.gov (2016g)

AFO-18

Phase I

CAF01

Plasmodium falciparum RH5 protein HIV-1 peptide antigens

HIV-1 virus

Ag85B-ESAT- Phase I 6 CM04-01 Phase I

CAF01

Human cytomegalovirus

Tandrup Schmidt et al. (2016) Watson et al. (2012), and van Dissel et al. (2014) Clinicaltrials.gov (2016g)

CN54gp140 MPLA

Phase I

MPLA

HIV-1 virus

Clinicaltrials.gov (2018f)

CTH522

Phase I

CAF01

Chlamydia trachomatis

Wern et al. (2017)

DPX-0907

Phase I Phase I

Ovarian, breast, and prostatic neoplasms HIV virus

Karkada et al. (2014)

eOD-GT8 60mer ID93

Polynucleotide-based adjuvant MPLA

Phase I

Glycopyranosyl lipid A

Mycobacterium tuberculosis

LipovaxinMM

Phase I

DC-SIGN antibody and interferon-γ

Metastatic melanoma

Baldwin et al. (2016), and Clinicaltrials.gov (2015f) Gargett et al. (2018)

L8-5 dLOS

Phase I

Deacylated l ipooligosaccharide

Neisseria meningitidis

Zollinger et al. (2012)

mRNA-1325

Phase I

mRNA

Zika virus

Kaczmarek et al. (2017)

Plasmid DNA

Tuberculosis antigen Ag85B-ESAT-6 Cytomegalovirus proteins Recombinant HIV-1 glycoprotein isolate 97CN54 Chlamydia trachomatis outer membrane protein CTH522 Tumor-specific HLAA2-restricted peptides HIV virus proteins Mycobacterium tuberculosis proteins Human melanoma cell line membrane vesicles Outer membrane proteins of Neisseria meningitides Zika virus peptides

Mycobacterium tuberculosis

Clinicaltrials.gov (2018e)

ONT-10

Phase I

Pentaerythritol-based lipid A

Hyperglycosylated MUC1 protein

Advanced breast and ovarian carcinoma

rCSP

Phase I

Glucopyranosyl lipid A

Plasmodium falciparum

TDENV-PIV TVDV C-GTF

Phase I Phase I Not stated

Dengue virus Dengue virus Streptococcus mutans

Clinicaltrials.gov (2012h) Danko et al. (2018) Childers et al. (1999)

Dpt

Not stated

AS01 and AS03 Plasmid DNA Liposomal Streptococcus mutans glucosyltransferase Dust mite body extract

Plasmodium falciparum circumsporozoite protein Inactivated dengue virus Dengue proteins Streptococcus mutans glucosyltransferase

Nemunaitis et al. (2013) and Clinicaltrials.gov (2013c) Clinicaltrials.gov (2018g)

Mite allergy

Alvarez et al. (2002)

Dust mite antigens

HLA, human leukocyte antigen; LNP, lipid-based nanoparticle; mRNA, messenger RNA; MPLA, monophosphoryl lipid A.

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also been tested against bacterial, parasitic, and fungal infections using protein components from these pathogens (Schwendener, 2014; De Serrano and Burkhart, 2017). Additionally as liposomal gene delivery has advanced, cationic liposomes have been developed as vectors for nucleic acid vaccines (Schwendener, 2014; Kramps and Elbers, 2017; Richner et al., 2017; Meyer et al., 2018; Scorza and Pardi, 2018). Liposomal vaccines provide an ideal platform to modulate immune responses by incorporating lessons from both small-molecule delivery and gene delivery to mediate epitope-specific immune responses for two reasons. One reason is the versatility of liposomes in incorporating a variety of bioactive molecules that can modulate the immune response (Schwendener, 2014; De Serrano and Burkhart, 2017; Alving et al., 2016). The other reason is the targeting of liposomal vectors for recognition by phagocytic cells, which allows for efficient processing and presentation to adaptive immune cells (Schwendener, 2014; De Serrano and Burkhart, 2017; Alving et al., 2016). The modular approach to vaccine development utilizing lipid nanoparticles provides opportunities to tailor epitope-specific responses to a variety of immunogens by incorporating adjuvants (see Section 6.4.2) and immunogens within the same particle.

6.4.1 Principles of Liposomal Vaccines APCs rapidly ingest liposomes after injection and process their contents for continued immune activation (Alving, 1991). This processing results in PRR-mediated activation by liposomal adjuvants (see Section 6.4.2) and presentation of antigens on MHC-II, leading to the induction of an immune response against the antigen. Antigens and adjuvants can be bound to liposomes via electrostatic interaction to the lipid surface, covalent and noncovalent anchoring to lipids, and encapsulation within the lipid bilayer (Watson et al., 2012). While each of these methods can be optimized to achieve adequate delivery, the method used can affect the efficacy of the formulation. Encapsulation provides a means to protect molecules from degradation by carrying liposomal therapies within the aqueous interior of the liposome. However, encapsulation efficiency can vary depending on the encapsulation method and antigen used, ranging from 25% to 90% (Tandrup Schmidt et al., 2016; Alving, 1991; Shariat et al., 2014; Eloy et al., 2014; Sur et al., 2014). Upon administration, encapsulated antigens are taken up by lymphocytes and APCs, which disrupt the formulations and process the antigens for presentation. Antigen presentation relies primarily on professional APCs, such as macrophages and dendritic cells. While B cells can also act as APCs for encapsulated antigens, early studies showed that liposomal vaccine strategies were less effective when APCs are not present, possibly because B cells have less phagocytic activity than do APCs (Watson et al., 2012; Alving, 1991) and the lack of epitope recognition using encapsulated antigens. Covalently anchoring epitopes to lipids for

Modulating the immune response with liposomal delivery

presentation in liposomes has been shown to enhance their immunogenicity (Watson et al., 2012). Despite a lack of direct evidence for direct B cell engagement, the enhanced immunogenicity is possibly due to the fact that surface presented antigens can be taken up by B cells through receptor-mediated endocytosis, which allows B cells to present antigens more effectively to TH cells. In general surface-conjugated antigens elicit greater antibody responses than encapsulated ones, although CTL responses are similar with either method (Watson et al., 2012). With the advent of liposomal gene delivery technologies, DNA and RNA vaccines have become a novel possibility for creating vaccines against viruses and cancer cells (Hasson et al., 2015; Lee et al., 2018). Unlike other liposomal vaccines that rely on peptide antigens, these vaccines deliver genetic material to encode antigens within cells (Pardi et al., 2018). Antigens are then presented on MHC-I to induce an immune response to intracellular antigens similar to a viral infection or tumor cell. DNA vaccines using various cationic liposome vectors have been engineered against herpes simplex virus 1 and influenza A virus (Schwendener, 2014; Lee et al., 2018; Liu et al., 2014; Rodriguez et al., 2013). Additionally, the liposomal system Vaxfectin has been used in various animal models to enhance the immune response against herpes simplex type 2, measles, influenza, malaria, and simian immunodeficiency virus (Shlapobersky et al., 2012; Lin et al., 2013; Luo et al., 2014; Kulkarni et al., 2013; Smith et al., 2013). RNA formulations have been tested in various animal models of melanoma, pancreatic cancer, and lung cancer, as well as various viral infections (Scorza and Pardi, 2018; Pardi et al., 2018; Li et al., 2011). Recently, a lipid nanoparticle-based mRNA vaccine was tested against Ebola virus and demonstrated efficacy and safety in guinea pigs using two mRNA vectors of viral glycoprotein delivered with a proprietary liposomal formulation of phosphatidylcholine, cholesterol, and PEGylated lipid (Meyer et al., 2018). Review of various other examples of these novel vaccine systems has been published recently (Schwendener, 2014; De Serrano and Burkhart, 2017; Pardi et al., 2018).

6.4.2 Liposomal adjuvants Adjuvants are vaccine components that increase the immunogenicity of antigens by stimulating immune activity to enhance the response against the target antigen or by improving antigen delivery to lymphocytes (Murphy and Weaver, 2016). Liposomes alone serve as vaccine adjuvants due to their ability to stabilize antigens, increase their bioavailability, reduce toxicity, and tailor the immune response achieved during immunization (Schwendener, 2014; Watson et al., 2012; Christensen et al., 2011; Alving et al., 2016). Furthermore the physical characteristics of liposomes, such as size and circulation half-life, can influence their immunogenicity. For example, larger liposomes, as well as PEGylated liposomes, have longer circulation half-life, which leads to prolonged antigen exposure (Schwendener, 2014). Larger liposomes have also been

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shown to induce a more inflammatory TH1 response compared to smaller liposomes with the same adjuvant (Badiee et al., 2012). The lipid components of liposomes can also alter the immune response. Saturated lipid tails have a tendency to activate helper T cells (TH), while unsaturated fatty acids tend to activate both TH cells and CTLs (Marasini et al., 2017). Despite improving the immune response against antigens, liposomes are relatively immunologically inert. In most instances, liposomal formulations are not sufficient to induce immunity toward antigens and must be formulated with adjuvants to induce an immune response (Murphy and Weaver, 2016). Several adjuvants have been developed and investigated, but not all adjuvants are compatible with liposomes and others have not been fully vetted for clinical applications (Fig. 6.5). Currently the US FDA

Figure 6.5 Liposomal adjuvants. Adjuvants evaluated in conjunction with liposomal formulations target a variety of receptors to activate cellular and humoral immune responses. The versatility of liposomes enable facile incorporation of lipid-based adjuvants into the liposomal bilayer, or aqueous soluble adjuvants to be incorporated in the liposome interior or as a mixture after the liposomes are formed. The modular approach to liposomal formualtions allows for diverse mixtures of antigens and adjuvants to be explored to achieve desired immune outcomes.

Modulating the immune response with liposomal delivery

has approved use of aluminum salts (or alum) and the combination of alum and monophosphorylated lipid A, AS04, to be used in human trials. Additionally European authorities authorize alum, AS04 and MF59, a squalene-based oil in water emulsion (Murphy and Weaver, 2016; Lee and Nguyen, 2015). While all liposomal adjuvants seek to enhance immunity against target antigens, the complexity of the immune system precludes utility of a single adjuvant for all vaccine applications. Different antigens require different immunogenic adjuvants to induce a robust immune response (Yasuda et al., 1977; Kinsky and Nicolotti, 1977). As with antigens, adjuvant association with liposomes varies depending on the properties of the molecule and requires optimization. Freund’s adjuvant, an oil-in-water emulsion of inactivated mycobacteria, is one of the most commonly used adjuvants in preclinical vaccine research and one of the first adjuvants used to enhance immunity against pathogenic antigens (Allison and Gregoriadis, 1974; Murphy and Weaver, 2016; Alving et al., 2016; Alving, 1991). The peptidoglycan muramyl dipeptide is responsible for the activity of Freund’s adjuvant, which can elicit a TH1 and TH2 response after activation of the NOD2 receptor in APCs and has been tested in various liposomal vaccine formulations targeting pathogens and cancer (Alving, 1991). Other liposomal adjuvants include monophosphoryl lipid A (MPLA) and PAM3CysSK4 (Giddam et al., 2012; Venditto et al., 2013; Venditto et al., 2014a). Lipid A is the lipid portion of lipopolysaccharide (LPS) and elicits a similar response to LPS, by activating TLR4. MPLA, which lacks one of the two phosphate groups in the parent molecule, exhibits diminished inflammation upon vaccination. Additionally incorporation of MPL into liposomes reduces the systemic adverse effects of MPLA (Alving et al., 2016). Due to its efficacy and safety profile, MPLA has been evaluated in several preclinical models and is a component of AS01, AS02, AS04, and which have progressed to clinical trials (Schwendener, 2014; Alving et al., 2016). Synthetic versions of MPLA have also been investigated including glycopyranosyl lipid adjuvant, a hexaacylated derivative of lipid A that activates TLR4 at the same order of magnitude as MPLA and LPS (Coler et al., 2011). More recent studies of this adjuvant have shown that it may induce a stronger TH1 response than MPLA (McKee and Marrack, 2017). PAM3CysSK4 is a TLR2 agonist, which also readily incorporates into liposomes and has been found to elicit a robust immune response and can shift the IgG1/IgG2a balance toward a less inflammatory TH2 response in mice (Giddam et al., 2012; Bal et al., 2011). CpG oligonucleotides also serve as an important adjuvant for use in liposomal vaccines. CpG oligonucleotides are similar to bacterial DNA and can trigger TLR9 to increase antigen expression and TH1 responses to antigens (Schwendener, 2014; Bal et al., 2011; Rao et al., 2004). Cyclic dinucleotides targeting the cytosolic receptor, stimulator of interferon genes, induces similar responses as CpG against antigens. Preclinical studies have demonstrated their potential as adjuvants in liposomal vaccines

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(Van Dis et al., 2018; Dubensky et al., 2013). In addition to their utility as vectors, cationic liposomes alone can serve as adjuvants due to their increased immunogenicity compared to classical liposomes and can induce robust TH1 and TH17 responses (Christensen et al., 2011; Alving et al., 1977). CAF01 is liposomal vaccine vector composed of the cationic lipid dimethyldioctadecylammonium with the glycolipid trehalose 6,6-dibehenate, an analog of cord factor from Mycobacterium tuberculosis. It has been shown to induce a TH1 type response against M. tuberculosis and other bacterial species (Schwendener, 2014; Alving et al., 2016). The mechanisms by which cationic liposomes enhance immune responses are not fully understood but may be related to increased uptake of immunogens into cells, the direct effects of the cationic lipids on immune cells, or the cytotoxic activity of the lipids (Christensen et al., 2007, 2011).

6.4.3 Liposomal vaccine clinical trials To date, two liposomal vaccines have been approved for clinical use. The first was the virosome Epaxal, which was approved in 1996 after demonstrating clinical efficacy against hepatitis A. Epaxal consists of inactivated hepatitis A virus adsorbed to virosomes formulated with influenza hemagglutinin and neuraminidase (Lim et al., 2014; Chappuis et al., 2017). Inflexal V is a virosomal influenza formulation that incorporates the hemagglutinin surface molecules of the influenza virus. It was approved for use against influenza after various trials demonstrating its efficacy in a wide range of patients (Herzog et al., 2009; Gasparini et al., 2013; Conne et al., 1997). In addition to Epaxal and Inflexal V, several vaccine formulations against other pathogens have been studied clinically. Various antimalarial vaccines have been tested using AS01, MPLA, and QS21 as adjuvants (see Table 2.1) (Alving et al., 2016). In 1999 the results of an intranasal liposomal vaccine against Streptococcus mutans showed an increased IgA response in subjects’ saliva following immunization with liposomal formulations and more recently, in 2012, a liposomal vaccine against Neisseria meningitides yielded promising results (Tandrup Schmidt et al., 2016). Liposomal tuberculosis vaccines containing CAF01 and AS01 have also shown promising results in early clinical trials (Montoya et al., 2013; van Dissel et al., 2014). CAF01 has also been tested as an adjuvant in an HIV-1 vaccine formulation (Roman et al., 2013; Childers et al., 1999). Liposomal formulations for cancer vaccines have been developed that target mucin 1, HPV viral proteins, and overexpressed tumor proteins, such as HER2/neu and telomerase (Tandrup Schmidt et al., 2016; Hamilton et al., 2012). Tecemotide, or Stimuvax, is a formulation containing MPLA tested as an adjuvant for the synthetic mucin 1 lipopeptide BLP25 in phase III clinical trials for non-small-cell lung cancer (Wu et al., 2011). BLP25 was also studied in prostate cancer, although it only reached phase II (North and Butts, 2005). DPX-0907 uses the liposomal vaccine platform DepoVaxt to target HLA-A2-restricted peptides as antigens in patients with advanced

Modulating the immune response with liposomal delivery

breast and prostate cancer (Karkada et al., 2014). Lipovacin-MM was recently evaluated for the treatment of malignant melanoma, although results have not yet been published (Grippin et al., 2017).

6.5 Conclusions and future directions The precedent established by the pioneers of liposomal drug delivery has positioned the field to have significant impact in a variety of therapeutic areas. The inventions that resulted in the discovery and approval of Doxil propelled the development of several other formulations that are currently approved or in clinical trials (Barenholz, 2012). Importantly, these discoveries have improved patient care and decreased disease burden by harnessing the characteristics of liposomes to enhance the solubility, pharmacokinetics, pharmacodynamics, and targeting of small-molecule therapeutics (Barenholz, 2012). Extension of these strategies into gene delivery and vaccine design has broadened the scope and clinical impact of LNPs. Importantly the advent of immune modulation provides an opportunity for synergistic discoveries to target and manipulate immune mechanisms associated with disease progression through the use of LNPs (Gregoriadis, 2016; Cullis and Hope, 2017; Alving et al., 2016). The utility of LNPs is highlighted by the number of ongoing clinical trials employing liposomal formulations presented throughout this chapter. Continued efforts to identify novel therapeutics and understand the ever-evolving pathogenic landscape positions LNPs to have continued impact in patients. These efforts require continued collaboration between industry, government, and academia to facilitate development and clinical translation of liposomal formulations by repurposing therapeutics for new indications and identifying beneficial formulations to enhance promising drugs with poor in vivo profiles. Strategically focused collaborative networks and partnerships will ensure that preclinical data are targeted and focused on clinical translation. Many investigational drug delivery strategies have not progressed past preclinical studies in mice, which may be indicative of a lack of appropriate clinical guidance (Venditto and Szoka, 2013). However, the established clinical utility of liposomal formulations has paved the way for continued investigation. As liposomes approach the half-century mark since their discovery, new avenues of research are possible that could not have been considered at the time of their inception. Most notably the shift toward personalized medicine has opened possibilities to tailor therapeutic interventions to patient’s individual needs based on gene sequencing, protein expression, and disease severity. The combinatorial approach to LNP development provides a diverse compositional space to meet a specific patient’s needs.

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Efficient strategies to exchange targeting moieties on preformulated LNPs will provide a mechanism to diagnose a patient, based on cell surface proteins and link an antibody or other targeting moiety for improved delivery and treatment. This strategy has clear benefits in cancer, due to the differential protein expression of tumors between patients. Utilizing this strategy in infectious disease provides a mechanism to attach relevant peptides to the surface of the LNP vaccine to induce epitope-specific immune responses. Importantly, altering the surface functionality of LNPs will impact the pharmacokinetic/pharmacodynamic (PK/PD) of the formulations. Algorithms capable of predicting the PK/PD of small-molecule therapeutics exist; however, these technologies have not translated to LNPs. Software capable of predicting LNP circulation and delivery, based on patient characteristics will have a profound impact on the future of the field, especially as rationally tailored formulations move closer to reality. Altogether, the current clinical landscape of liposomal therapeutics provides a foundation for continued exploration in a variety of disease states and suggests a bright future for continued success.

List of abbreviations APC ApoE ASO CTL DC-cholesterol DLinDMA DLinMC3DMA DODMA DOGS DOPE DOSPA DOTAP DOTMA DPPC DSPE eNOS HER2/neu HIV HLA HPV Ig IL

antigen-presenting cell apolipoprotein E antisense oligonucleotide cytotoxic T lymphocytes 3ß-[N-(N0 ,N0 -dimethylaminoethane)-carbamoyl]cholesterol hydrochloride 1,2-dilinoleyloxy-3-dimethylaminopropane dilinoleylmethyl-4-dimethylaminobutyrate 1,2-dioleyloxy-3-dimethylaminopropane dioctadecylamidoglycylspermine dioleoylphosphatidylethanolamine 2,3-dioleyloxy-N-[2(sperminecarboxamido)ethyl]-N,N-dimethyl1-propanaminium 1,2-dioleoyl-3-trimethylammonium-propane chloride 1,2-di-O-octadecenyl-3-trimethylammonium propane chloride dipalmitoylphosphatidylcholine distearoylphosphoethanolamine endothelial nitric oxide synthase receptor tyrosine-protein kinase erbB-2 human immunodeficiency virus human leukocyte antigen human papilloma virus immunoglobulin (e.g., IgM and IgD to IgA, IgE, and IgG) interleukin (e.g., IL-4, IL-6, IL-10)

Modulating the immune response with liposomal delivery

LLO listeriolysin O LNP lipid nanoparticle LPS lipopolysaccharide MHC major histocompatibility complex miRNA microRNA MPL monophosphorylated lipid A mRNA messenger RNA NF-κB nuclear factor κ B NOD-like receptor nucleotide-binding oligomerization domain-like receptor PAMP pathogen-associated molecular pattern PEG polyethylene glycol PRR pattern recognition receptor RISC RNA-induced silencing complex RNAi RNA interference S1P sphingosine-1-phosphate shRNA short hairpin RNA siRNA small-interfering RNA TH helper T cells TLRs toll-like receptors TNFα tumor necrosis factor α Treg regulatory T cell UGG unsaturated guanidinium glycoside

Acknowledgments This work was supported by University of Kentucky, College of Pharmacy. D.N. was supported by a Lyman T. Johnson fellowship from the University of Kentucky. D.H. was supported by a training grant through the National Center for Advancing Translational Sciences, NIH (TL1TR001997) and a Predoctoral Fellowship from the American Heart Association (19PRE34430120). J.S. was supported by a grant from the Kentucky Spinal Cord and Head Injury Research Trust (15-12A). V.V. was supported by a Scientist Development Grant from the American Heart Association (17SDG32670001).

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Recent advances in solid lipid nanoparticles formulation and clinical applications Helena Rouco1, Patricia Diaz-Rodriguez1, Carmen Remuñán-López2 and Mariana Landin1 1 R 1 D Pharma Group (GI-1645), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain 2 NanoBiofar Group (GI-1643), Department of Pharmacology, Pharmacy and Pharmaceutical Technology, Faculty of Pharmacy, University of Santiago de Compostela, Santiago de Compostela, Spain

7.1 Lipid nanoparticles Lipid nanoparticles are colloidal systems composed by solid lipids and stabilized by a surfactant layer. This lipid-based nanocarriers group includes solid lipid nanoparticles (SLN) and nanostructured lipid carriers (NLC) (Martins et al., 2007).

7.1.1 Solid lipid nanoparticles SLN are the first generation of lipid nanoparticles (Müller et al., 2011). SLN are constituted by a lipid core, solid at both room and body temperatures, surrounded by emulsifier molecules (Souto et al., 2007; Gordillo-Galeano and Mora-Huertas, 2018; Muchow et al., 2008). The main advantages of SLN include excellent biocompatibility, ability to include a wide variety of bioactive molecules, rapid and effective formulation through different techniques without the need to include organic solvents, and suitability for scaling up (Khosa et al., 2018). The main limitations of SLN are associated with a reduced drug loading (DL) ability due to the lipid’s crystalline nature and drug leakage during storage, related to gelation phenomena and lipids β-form that will described in Section 7.2.1.1. Gelation is the transformation of the colloidal suspension into a viscous gel, this process usually takes place after particle aggregation associated with lipid crystallization (Heurtault, 2003).

7.1.2 Nanostructured lipid carriers NLC were designed to overcome the drawbacks of SLN and constitute the second generation of lipid nanoparticles (Müller et al., 2011). As for SLN, NLC matrix is solid Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00007-2

© 2020 Elsevier Inc. All rights reserved.

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at both room and body temperatures (Muchow et al., 2008; Khosa et al., 2018). NLC show several advantages over the first generation of lipid nanoparticles, such as increased loading capacity, improved stability, and higher flexibility in drug release modulation (Khosa et al., 2018). Their improved characteristics are associated to the imperfect matrix structure created by mixing two spatially different molecules, a solid lipid and an oil (Muchow et al., 2008).

7.2 Formulation components SLN are composed of an oil phase and an aqueous phase in a ratio of 0.1% to 30% (w/w). Generally the aqueous phase includes between 0.5% and 5% of surfactants. The oil phase can be constituted by a solid lipid or a mixture of different solid lipids. NLC are characterized by the combination of liquid (oil) and solid lipids in a ratio between 70:30 and 99.9:0.1 (solid lipid:oil), and total solid content of formulations can reach 95% (Pardeike et al., 2009).

7.2.1 Lipids Lipids are a broad group of molecules that include triglycerides, partial glycerides, fatty acids, steroids, and waxes. Oils and fats are natural mixtures of mono, di, and triglycerides containing fatty acids of different chain length and unsaturation degree (Gaba et al., 2015). Most of them are approved as generally recognized as safe substances (Shah et al., 2015b). Biocompatibility, biodegradability, capability of forming small particles (in the nanorange), DL capacity, and stability in aqueous dispersion are the requirements of a lipid to be considered a suitable material to produce an optimal lipid nanoparticle formulation. Several aspects must be carefully considered in the selection of the appropriate lipid variety to encapsulate a specific therapeutic agent: polymorphic state, type, amount used in the formulation process, and liquid lipid combination possibility (Pathak et al., 2011). 7.2.1.1 Lipids polymorphic state Four polymorphic forms have been described for lipids: α, β, and βʹ as the main forms and βi form (transitional form between βʹ and β) can also be observed, but in a lesser extent. Polymorphs are chemically identical but show different solubility, X-ray diffraction (XRD) pattern, and melting point (Pathak et al., 2011). After nanoparticles formation, lipids tend to partially recrystallize, in a highly disordered α form. During storage, polymorphic transitions from α form to a more stable β form via βʹ can occur, which result in a reduction of the imperfections in the lipid matrix and the expulsion

Recent advances in solid lipid nanoparticles formulation and clinical applications

of the encapsulated drug. The extension of these phenomena depends on the storage conditions and the transition kinetics of the lipid(s) (Pathak et al., 2011). Recrystallization processes also affect particle size, size distribution, and shape of the resulting nanocarriers (Pathak et al., 2011). During storage, nanoparticles may undergo directional crystal growth, altering their morphology from the spherical shape, characteristic of the α form, to a needle or plated-shape morphology, characteristic of the β form. If these needle-shaped structures with increased surface area are not stabilized by surfactants, flocculation occurs destabilizing the suspensions (Pathak et al., 2011; Weiss et al., 2008). Flocculation phenomena may be responsible for the increase in particle size and size distribution reported in the literature for this type of systems (Suresh et al., 2007). On the other hand, extremely slow transitions from βʹ to the stable β form have been related to slight reductions in particle size during storage (Chen et al., 2006). Proper knowledge and control about polymorphic behavior of the lipids in the formulation are necessary to develop stable nanocarriers. 7.2.1.2 Types of lipids Different lipid types such as fatty acids, glycerides, waxes, hard fats, and steroids can be employed in lipid nanoparticle formulation (Pathak et al., 2011). Fig. 7.1 provides some examples of them. Fatty acids are hydrocarbon chains that end in carboxylic acid groups with different chain lengths and unsaturation degrees that modulate their properties. Fatty acids of biological systems usually contain between 14 and 24 carbon atoms (C), being 16- and 18-C fatty acids the most common ones (Berg et al., 2002).

Figure 7.1 Examples of chemical structures of different lipid types.

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Glyceride group includes triglycerides and partial glycerides (mono and diglycerides), which differ in their degree of esterification with fatty acids (Khatak and Dureja, 2018). Triglycerides are obtained by the esterification of glycerol with three fatty acids, which can be equal or not, usually showing a chain length of 16, 18, or 20 C. They can be classified according to their C chain length as short-chain ( , 5 C), mediumchain (6 12 C), and long-chain ( . 12 C), or according to their acyl substitution as mono-, di-, and tri-acid (Khatak and Dureja, 2018). Medium-chain triglycerides present several advantages over long-chain triglycerides, such as higher solvent capacity and better stability toward oxidation (Khatak and Dureja, 2018). Partial glycerides are esters of glycerol with fatty acids that contain unreacted hydroxyl groups. These substances named monoglycerides or diglycerides show interesting emulsifying properties, being considered as nonionic emulsifiers, since they do not produce ions in solution (Sharma et al., 2017). Mixtures of mono, di, and triglycerides (acylglycerol mixtures, e.g., Compritol 888 ATO) are widely used for lipid nanoparticle formulation (Aburahma and Badr-Eldin, 2014). Waxes are defined as esters of fatty acids with alcohols, other than glycerol. They may contain free fatty acid functions, as is the case of beeswax, or free hydroxyl groups inside the molecule (Jenning and Gohla, 2000). “Hard fat” includes (semi-) synthetic mixtures of mono, di, and triglycerides of fatty acids C10 C18. These compounds show a low melting point, usually between 33 C and 36 C (Bouwman-Boer, 2015). Steroids include any synthetic or natural compound with a 1,2-cyclopentanophenanthrene backbone along with a side chain of variable length and two or more oxygen functions, such as carboxyl, carbonyl, aldehyde, or alcohol (Goad and Akihisa, 1997). The type of lipid selected strongly affects formulation properties as drug loading. As a general rule, the lesser the lipid crystalline matrix order, the greater the DL capacity (Pathak et al., 2011). Therefore lipids that produce highly crystalline particles with perfect lattices (such as monoacid triglycerides) often lead to the expulsion of the drug, whereas lipids that are composed of mixtures of mono, di, and triglycerides (such as Precirol) or lipids containing fatty acids of different chain lengths, which form lessperfect crystals, offer more space for drug encapsulation (Vivek et al., 2007). Other factors such as lipids hydrophobicity also determine the entrapment efficiency (EE) of lipophilic drugs. The high-lipid hydrophobicity derived from the long-chain fatty acids attached to the triglycerides (e.g., tripalmitin) results in an increase in the accommodation of lipophilic drugs and, therefore in a higher drug EE (Kumar et al., 2007). In addition, emulsifying properties of mono and diglycerides also contribute to improve DL capacity (Müller et al., 2000).

Recent advances in solid lipid nanoparticles formulation and clinical applications

Lipid’s structure seems to play a role on nanoparticles size. Lipids with smaller acyl chains have been reported to produce small-sized particles (Triplett and Rathman, 2009). The composition of the lipid matrix influences nanoparticle size and size distribution. Lipid polarity determines the surface charge of the nanoparticles and the zeta potential, conditioning their degree of agglomeration. Therefore the use of lipids with a suitable polarity allows the production of nanoparticulate systems with a monodisperse particle size distribution and good stability (Pathak et al., 2011). Other lipid excipients properties, such as speed of lipid crystallization and self-emulsifying capacity, are also known to affect the final size of lipid nanoparticles (Vivek et al., 2007). Drug release profile can fluctuate between nanoparticles prepared using different lipids. In physiological conditions, drug release from lipid nanoparticles occurs by simultaneous diffusion and erosion mechanisms. The lipid matrix degradation is mainly promoted by enzymes, and in a less extent, by hydrolytic processes (Pathak et al., 2011). Prolonged release has been associated with slow drug diffusion from the lipid nanoparticles, which is highly related to some characteristics of the lipids such as hydrophobicity, hydrophilic lipophilic balance (HLB), or melting point (Pathak et al., 2011). Lipids of high hydrophobicity and low HLB are preferred to achieve controlled release drug delivery systems (Pathak et al., 2011). Nanoparticles stability during storage is an important factor that must be considered during lipid selection. Waxes, as cetyl palmitate and beeswax, exhibit superior stability in terms of particle size when compared to glycerides. However, its capacity for drug retention during storage depends, to a large extent, on the storage temperature and its tendency to form supercooled melts, as in the case for cetyl palmitate. Within glycerides, the presence of high amounts of partial glycerides, such as monoglycerides, produces physical destabilization. On the other hand, the emulsifying properties of mono and diglycerides improve the surfactant layer around the nanoparticles avoiding agglomeration (Jenning and Gohla, 2000). Acylglycerol mixtures such as glyceryl monostearate and glyceryl behenate are more efficient to prevent drug leakage and maintain DL during storage due to their less ordered lipid matrix (Jenning and Gohla, 2000). As noted earlier, highly ordered crystal packing of glycerides often leads to the drug expulsion in solidified state. Therefore pure triglycerides, as in the case of tripalmitate, show the same behavior as waxes (Jenning and Gohla, 2000). Finally it is necessary to emphasize that, since most commercial lipid excipients are composed of mixtures of different chemical compounds, the inter- and intramanufacturer variability is important and can have a great impact on the properties of the nanoparticles (Mehnert and Mader, 2001). Considering these aspects should help to avoid unexpected results after nanoparticle preparation.

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7.2.1.3 Lipid proportion The amount of lipids used during lipid nanoparticles formulation affects particle size and drug payload. Generally an increase in lipid amount leads to an increase in particle size due to less efficient homogenization and greater tendency of lipids to agglomerate (Pathak et al., 2011; Mehnert and Mader, 2001). Some authors have reported this phenomenon when using increasing amounts of lipids during the formulation of nanoparticles by microemulsion or solvent injection techniques (Tiyaboonchai et al., 2007; Shah and Pathak, 2010). Other authors have pointed out as a main factor the solvent:lipid ratio when using a modified solvent emulsification evaporation technique for the nanoparticle formulation (Vitorino et al., 2011). DL capacity is directly related to drug solubility in the lipid matrix. If drug solubility is high, a higher amount of lipid leads to an increase in drug payload. However, if drug solubility is low, the opposite effect occurs. In addition, the use of a high amount of lipid increases the risk of crystallization, which can drastically reduce encapsulation efficiency (Pathak et al., 2011). 7.2.1.4 Presence of a liquid lipid The inclusion of a liquid lipid (oil) in the nanoparticle matrix is the main difference between SLN and NLC (Severino et al., 2012). It affects formulation parameters such as DL, size, and drug release profile (Pathak et al., 2011). NLC are highly heterogeneous structures with a large number of imperfections, which improves their DL capacity (Pathak et al., 2011). In addition, the drugs solubility in liquid lipids is usually greater than in solid lipids (Müller et al., 2002). Several studies have described an increase in encapsulation efficiency and DL with the use of an increased amount of liquid lipid (Song et al., 2016; Rouco et al., 2018). However, the opposite effect has also been described due to a loss in the drug immobilization capacity of the matrix by insufficient proportion of solid lipids (Kim et al., 2010). Several authors have reported reductions in particle size as a consequence of the introduction of a liquid lipid into the nanoparticle matrix (Song et al., 2016; Rouco et al., 2018), which may be related to differences in viscosity between both solid and liquid lipids. The inclusion of a large amount of oil in the lipid matrix reduces the surface tension, which leads to smaller nanoparticles with narrower size distributions (Pathak et al., 2011). However, bigger particle sizes have also been described as a result of the increase in the amount of liquid lipid in the formulation (Pathak et al., 2011; Kim et al., 2010). If the amount of oil further increases, there is an abrupt particle size reduction related to the solid lipids crystallization inhibition (Pathak et al., 2011). Changes in particle size were also related to the reduced retention capacity of the lipid core described above, which leads to drug expulsion, producing changes in surface charge and leading to larger particle sizes and broader polydispersions (Kim et al., 2010).

Recent advances in solid lipid nanoparticles formulation and clinical applications

The amount of liquid lipid in the nanoparticles also determines drug release profile. A high amount of oil in the matrix leads to an increase in the drug release rate due to the reduction in crystallinity and the increase of drug mobility into the NLC lipid core. The amount of liquid lipid must be carefully selected to optimize drug release profile (Kim et al., 2010).

7.2.2 Surfactants or emulsifiers Surfactants are amphipathic molecules composed of a nonpolar hydrophobic portion, consisting of a branched or linear hydrocarbon or fluorocarbon chain containing 8 18 carbon atoms, attached to a polar hydrophilic moiety. They are classified as ionic, nonionic, and amphoteric (Som et al., 2012). Ionic (anionic or cationic) surfactants confer electrostatic stability, while nonionic emulsifiers provide stability by steric repulsion (Khatak and Dureja, 2018). Most of the nonionic emulsifiers are too small to confer steric stability, but they are able to increase stability through the Gibbs Marangoni effect (Shah et al., 2015b), which consists on the rapid migration of the emulsifier toward the particle surface. This movement drags water, restoring its presence between particles and reducing surface concentration gradients (Wilde et al., 2004). Nonionic surfactants are known to modulate the action of the lipase/colipase complex, which is responsible for “in vivo” lipid degradation. Pluronic and Tween are widely used nonionic emulsifiers (Shah et al., 2015b). Amphoteric surfactants have several functional groups showing characteristics of anionic or cationic emulsifiers at high and low pH, respectively (Khatak and Dureja, 2018). Phospholipids and phosphatidylcholines, such as Lipoid S 75 and Lipoid S100, are examples of amphoteric surfactants used in lipid nanoparticles formulation (Souto et al., 2007). Sodium cholate and sodium dodecyl sulfate, among others, are ionic surfactants, which are commonly used for the preparation of lipid nanoparticles (Shah et al., 2015b). Surfactants can also be classified according to their HLB, as hydrophilic (HLB . 11), intermediate (HLB 9 11), or lipophilic (HLB , 9). HLB values of emulsifiers are determined by the equilibrium between the strength and size of their hydrophilic and lipophilic groups (Severino et al., 2012). Emulsifiers play two essential roles in lipid nanosystems: (1) facilitate the dispersion of the molten lipid in the aqueous phase during formulation process and (2) confer stability to the lipid nanoparticles after cooling (Shah et al., 2015b). In addition, they modulate the crystallization process of lipid nanoparticles, since they interact with a large number of lipid molecules due to the small size of the nanoemulsions. The use of surfactants can improve the kinetic stability of the crystal structure generated even if it is thermodynamically less stable than other polymorphic forms (Weiss et al., 2008).

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Factors such as the administration route, the HLB of the emulsifier, the desired particle size, or the preferred “in vivo” degradation pattern for the lipid matrix strongly condition the selection of the optimal surfactant (Shah et al., 2015b). The best option for efficiently stabilizing lipid nanosystems and preventing particle agglomeration is using combinations of emulsifiers (Mehnert and Mader, 2001). A method of selecting emulsifiers based on HLB has been proposed. The HLB system consists of using a surfactant or combinations with an HLB as close as possible to the HLB of the lipids. This would ensure a good performance of the resulting nanosystem (Severino et al., 2012). The estimation of the required HLB can be carried out using the Grifinth equation or considering the values provided by manufacturers (Kovacevic et al., 2011).

7.2.3 Other components In addition to lipids and emulsifiers, other agents such as counterions, surface modifiers, cryoprotectants, antimicrobial agents, and other preservatives may also be part of the lipid nanoparticles composition (Gordillo-Galeano and Mora-Huertas, 2018; Shah et al., 2015b). Counterions such as anionic polymers or organic ions can be incorporated into lipid nanoparticle formulations to encapsulate water-soluble cationic drugs (Shah et al., 2015b). Cryoprotectants such as sorbitol, glucose, and fructose are used in lyophilized formulations. Parabens and thiomersal may also be included as antimicrobial agents and preservatives (Gordillo-Galeano and Mora-Huertas, 2018).

7.3 Preformulation studies 7.3.1 Solubility studies Drug solubility in the lipid or combination of lipids should be properly characterized to know drug affinity for the nanoparticle matrix and establish the optimal drug:lipid ratio. For solid lipids, solubility studies are often performed by heating them 10 C above their melting point and successively adding small amounts of drug until lipid saturation, which is considered to occur when the excess of solid drug remains for more than 8 hours (Gaspar et al., 2016). Alternatively this study can be performed using a known amount of drug and increasing amounts of solid lipids. In this case, the end point of the experiment will be the amount of molten solid lipids that allows the formation of a clear solution (Alam et al., 2018). For liquid lipids this analysis is usually accomplished by determining the highest amount of drug that can be dissolved in each candidate lipid. For this, an excess of drug is incorporated to a known amount of lipid in a vial, stirred at room temperature

Recent advances in solid lipid nanoparticles formulation and clinical applications

for 72 hours, centrifuged, and the amount of drug dissolved is quantified (Alam et al., 2018). A similar procedure can also be used for other formulation raw materials such as organic solvents and surfactants. Excess of drug is added to a known amount of the tested component(s). The blend is mixed, sonicated, and finally kept in an incubator shaker. After 8 12 hours, mixtures are centrifuged, and the drug is determined in the supernatant (Ranpise et al., 2014).

7.3.2 Partitioning analysis Partitioning nature assay constitutes a different approach to investigate the affinity of a drug for the lipid matrix selected and is designed to predict, in an approximately way, the drug encapsulation in a certain lipid matrix. In this assay, drug is dispersed in a melted lipid along with hot phosphate buffer pH 7.4. The mixture is shaken for 30 minutes 10 C above the melting point of the selected lipid. The aqueous phase is then separated from the lipid by centrifugation, and drug content is quantified in the supernatant. The partition coefficient is calculated by the subtraction of the amount of drug present in phosphate buffer to the total drug added and the subsequent division by the amount of total drug (Vivek et al., 2007).

7.3.3 Compatibility between solid lipids and liquid lipids In the formulation of NLC, it is advisable to evaluate the compatibility between the selected lipids. This study can be carried out easily by mixing the solid lipid and the oil in different proportions at a temperature 5 C above the melting point of the solid lipid. Mixtures are analyzed visually after 1 and 24 hours of matrix solidification. Those that show a single phase are considered suitable for nanoparticles formulation (Ranpise et al., 2014). An alternative method to evaluate the miscibility between solid and liquid lipids consists of mixing liquid and solid lipid in different proportions at 85 C and cooling until solidification. Then hydrophilic filter paper is smeared with these mixtures and observed. The absence of oil droplets on the paper is indicative of a good miscibility between lipids (Kasongo et al., 2011).

7.4 Formulation procedures Formulation procedures of SLN and NLC can be categorized into three groups. High-energy methods, low-energy methods, and approaches based on organic solvents (Gordillo-Galeano and Mora-Huertas, 2018). Moreover several formulation techniques to produce lipid nanoparticles in solid state have also been reported (Battaglia and Gallarate, 2012), although they are not widespread.

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7.4.1 High-energy methods High-energy methods are based on the use of equipment, which produces particle size reduction through the generation of high-shear forces or pressure distortions (Gordillo-Galeano and Mora-Huertas, 2018). 7.4.1.1 High-pressure homogenization High-pressure homogenization can be performed at high or low temperature (GordilloGaleano and Mora-Huertas, 2018; Khosa et al., 2018; Das and Chaudhury, 2011). First lipids are heated 5 C/10 C above their melting point and the drug is dispersed in the melt. If the high-temperature process is used, the aqueous solution is also heated and added to the oil phase. The preemulsion is obtained using a high-shear mixer and homogenized in a high-pressure homogenizer at elevated temperature until the desired particle size is achieved. Finally the nanoemulsion is cooled at room temperature (GordilloGaleano and Mora-Huertas, 2018; Das and Chaudhury, 2011). If the cold technique is performed, the drug lipid melt is cooled using liquid nitrogen or dry ice and ground by a ball mill or a mortar. The powder obtained is dispersed in a cold aqueous-surfactant solution and homogenized using a high-pressure homogenizer at room temperature or even lower (Gordillo-Galeano and Mora-Huertas, 2018; Das and Chaudhury, 2011). High-pressure homogenization method is simple and cost-effective. Nevertheless, in the hot method, high temperature during homogenization is mandatory (Khosa et al., 2018), being a limitation for thermolabile drugs. Cold homogenization helps to overcome this problem and constitutes a well-established advantageous procedure for large-scale manufacture avoiding organic solvents (Khosa et al., 2018). 7.4.1.2 Emulsification sonication technique This technique is carried out by melting the drug lipid mixture above the melting point of lipids and dispersing it in aqueous-surfactant solution kept at the same temperature by means of a high-shear mixer. The resulting emulsion is ultrasonicated to reduce particle size, and finally nanoparticles are obtained by cooling the nanoemulsion at room temperature. The main disadvantage is the possibility of dispersions contamination with metals during sonication (Gordillo-Galeano and Mora-Huertas, 2018; Das and Chaudhury, 2011). 7.4.1.3 Supercritical fluid technology A compound in the supercritical state, at temperature and pressure above the critical values behaves like gas and liquid simultaneously. The solubility of a substance in the supercritical fluid (SCF) can be regulated through small changes in pressure. CO2 is the most commonly used SCF because of its low critical point, nontoxicity, and low cost (Battaglia and Gallarate, 2012). Two main methods have been described for lipid nanoparticle production, gas-assisted melting atomization (GAMA) and supercritical fluid extraction of emulsions (SCFEE).

Recent advances in solid lipid nanoparticles formulation and clinical applications

In GAMA, lipids are melted and exposed to supercritical CO2 in a mixing chamber. The saturated lipid mixture is sprayed through a nozzle into a chamber. The rapid depressurization of the mixture generates a high supersaturation and precipitation of lipid microparticles. The microparticles are collected, dispersed in water, and using a vortex or ultrasound, fragmented to obtain nanoparticles (Battaglia and Gallarate, 2012). In the SCFEE process, drug-loaded O/W (oil-in-water) emulsions are introduced into the upper part of an extraction column. The introduction of supercritical CO2 countercurrent through the lower part leads to the extraction of solvents and the precipitation of drug lipid particles. The main advantage of this technique is the high efficiency of elimination of solvents (Battaglia and Gallarate, 2012). 7.4.1.4 Hot high-shear homogenization This technique is performed by heating separately an aqueous and an oil phase 10 C above lipids’ melting point. The aqueous phase is added onto the oil phase, and homogenization is performed by a high-shear laboratory mixer at constant temperature. The obtained nanodispersions are cooled down to allow lipid crystallization (Gaspar et al., 2016). This technique is extremely simple, but it has the limitation of the potential presence of microparticles in the final dispersion (Alvarez-Trabado et al., 2017).

7.4.2 Low-energy methods This category includes methods that do not require a large amount of energy and/or those in which the reduction of the particle size can take place spontaneously (Gordillo-Galeano and Mora-Huertas, 2018). 7.4.2.1 Microemulsion technique In this method, nanoparticles are obtained spontaneously due to the high surfactant lipid ratio used (Gordillo-Galeano and Mora-Huertas, 2018). To generate nanoparticles, a preheated surfactant-containing aqueous solution is added onto a melted drug lipid mixture and emulsified under mild agitation. Then the emulsion is dispersed in an excess of cold water maintaining the agitation to achieve a nanoemulsion, which results in the formation of nanoparticles after droplet crystallization (GordilloGaleano and Mora-Huertas, 2018; Das and Chaudhury, 2011). The main advantage of this technique is its suitability for large-scale production (Khosa et al., 2018; Das and Chaudhury, 2011). Among its disadvantages, the high amount of emulsifying agents and the great dilution of the nanoparticles can be noted (Gordillo-Galeano and MoraHuertas, 2018; Khosa et al., 2018).

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7.4.2.2 Double emulsion This method is used to produce lipid nanoparticles including hydrophilic drugs or peptides (Gordillo-Galeano and Mora-Huertas, 2018), which are usually known as lipospheres, because of their large particle size (Das and Chaudhury, 2011). A blend of melted lipids and an active ingredient aqueous solution is mixed to form a W/O (water-in-oil) microemulsion, which is dispersed in an aqueous solution containing a hydrophilic stabilizer to form a W/O/W (water-in-oil-in-water) double emulsion. Lipid nanoparticles are obtained after emulsion dilution in cold water (GordilloGaleano and Mora-Huertas, 2018). 7.4.2.3 Membrane contractor technique In this procedure, a drug lipid melt is pushed through a hydrophobic porous membrane toward an aqueous-surfactant solution which is circulating inside the membrane module dragging away the lipid droplets. Finally these lipids droplets form the nanoparticles when the aqueous phase is cooled at room temperature (Gordillo-Galeano and Mora-Huertas, 2018; Khosa et al., 2018). Its main advantages are simplicity (Khosa et al., 2018) and possibility to achieve a continuous production (GordilloGaleano and Mora-Huertas, 2018). 7.4.2.4 Phase inversion technique Phase inversion induced by heat is performed by preparing a mixture of water, emulsifier, lipid, and drug under stirring, which is subsequently subjected to three heating and cooling cycles (85 60 85 ) (Khosa et al., 2018). These cycles produce the inversion of emulsion phases, progressively reducing droplet size (Gordillo-Galeano and Mora-Huertas, 2018). Nanoparticles are finally obtained by dilution in cold water (Gordillo-Galeano and Mora-Huertas, 2018; Khosa et al., 2018). Its principal advantages are the avoidance of organic solvents, while its main disadvantage is the tedious formulation process (Khosa et al., 2018). 7.4.2.5 Coacervation This formulation method is based on the fact that the presence of surfactants produces the precipitation of free fatty acids from their micelles. A salt of a fatty acid is dispersed in an emulsifier solution. The mixture is heated until the Kraft point of the salt of the fatty acid under agitation. When a clear solution is achieved, an ethanolic solution containing the drug is added gently under constant stirring to obtain a single phase. The addition of an acidifying solution or a coacervation agent promotes the formation of the nanoparticle suspension (Gordillo-Galeano and Mora-Huertas, 2018).

Recent advances in solid lipid nanoparticles formulation and clinical applications

7.4.3 Organic solvent-based approaches In this category, the particle size reduction is achieved by the addition of organic solvents (Gordillo-Galeano and Mora-Huertas, 2018) which is precisely its main drawback (Das and Chaudhury, 2011). 7.4.3.1 Solvent emulsification evaporation method In this technique, lipids dissolved in water-immiscible organic solvents (chloroform, cyclohexane) are emulsified with an aqueous phase containing surfactants by stirring. The evaporation of the solvent during the emulsification leads to lipids precipitation and nanoparticles formation (Khosa et al., 2018; Das and Chaudhury, 2011). The main advantage of this method is its suitability for thermosensitive drugs. However, the high dilution of the dispersions obtained constitutes its major drawback (Khosa et al., 2018; Das and Chaudhury, 2011). 7.4.3.2 Solvent emulsification diffusion method Solvent emulsification diffusion method involves the use of partially water-miscible organic solvents (ethyl formate, benzyl alcohol), which are saturated with water, to dissolve the lipids (Das and Chaudhury, 2011). The oil-in-water emulsion is dispersed into water under continuous stirring, producing the diffusion of the organic solvent and the lipids precipitation with the subsequent nanoparticle formation (GordilloGaleano and Mora-Huertas, 2018; Khosa et al., 2018; Das and Chaudhury, 2011). The solvent can be removed by distillation or ultrafiltration (Gordillo-Galeano and Mora-Huertas, 2018). The main disadvantage of this formulation procedure is the high dilution of the dispersions obtained (Das and Chaudhury, 2011). 7.4.3.3 Solvent injection method Lipids and active ingredients are dissolved in a water-miscible organic solvent (methanol, acetone, isopropanol) or a water-miscible solvent mixture. After that the organic solution is injected through an injection needle in an aqueous solution of emulsifiers under stirring, producing the solvent migration and nanoparticle precipitation (Gordillo-Galeano and Mora-Huertas, 2018; Das and Chaudhury, 2011). Among the advantages of this technique are equipment simplicity, easy handling, and agile production process (Das and Chaudhury, 2011; Table 7.1).

7.5 Characterization techniques 7.5.1 Particle size and size distribution The nanoparticles particle size and size distribution are evaluated by laser diffraction (LD) and photon correlation spectroscopy, also known as dynamic light scattering

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Table 7.1 Methods of lipid nanoparticle production: advantages and disadvantages. Method Advantages Disadvantages Examples

References

High-energy methods

Hot high-pressure homogenization

Simple, cost-effective, no organic solvents, easy to scale up

High temperature, Progesterone-loaded not suitable for SLN and NLC, thermolabile drugs, dibucaine-loaded high-energy input SLN

Cold high-pressure homogenization

Simple, cost-effective, easy to scale up, no organic solvents, suitable for thermosensitive drugs in some cases, suitable for hydrophilic drugs

More cycles and Lysozime-loaded SLN, higher pressures vinorelbine bitartraterequired, highloaded SLN energy input, high temperature used in melting step

Emulsification sonication High-shear mixing, no technique organic solvents, simplicity, low-surfactant concentration

Supercritical fluid technology

Hot high-shear homogenization

Potential metallic Vorinostat-loaded SLN, contamination, raloxifene-loaded polydisperse SLN suspensions, not suitable for thermolabile drugs

High efficiency of solvent Use of organic Camptothecin-loaded elimination, nanoparticles solvents, use of SLN, ketoprofenin solid state, rapid, complex loaded SLN and suitability for labile equipment indomethacin-loaded compounds SLN Simplicity, no organic Microparticles in the Rifabutin-loaded NLC solvents, low-surfactant final product, not and SLN, resveratrolconcentration, low cost suitable for loaded SLN and thermolabile drugs NLC

Khosa et al. (2018), Iqbal et al. (2017), Naseri et al. (2015), Sala et al. (2018), Garces et al. (2018), and Nasrollahi et al. (2013) Khosa et al. (2018), Iqbal et al. (2017), zur Mühlen et al. (1998), Garces et al. (2018), Nasrollahi et al. (2013), Cortesi et al. (2011), and Lin et al. (2010) Gordillo-Galeano and MoraHuertas (2018), Khosa et al. (2018), Iqbal et al. (2017), Garces et al. (2018), Gainza et al. (2015), and Chen et al. (2013) Alvarez-Trabado et al. (2017), Nasrollahi et al. (2013), Wang et al. (2017), and Rigon et al. (2016) Som et al. (2012), Ranpise et al. (2014), zur Mühlen et al. (1998), Garces et al. (2018), Nasrollahi et al. (2013), and Joshi and Müller (2009) (continued)

Table 7.1 (Continued) Method

Advantages

Disadvantages

Easy scale up, no organic solvents, reduced particle size, and size distribution

High proportions of emulsifiers, high nanoparticle dilution

Examples

References

Low-energy methods

Microemulsion technique

Ofloxacin-loaded NLC, Gordillo-Galeano and Moratopotecan-loaded Huertas (2018), Khosa et al. (2018), Iqbal et al. SLN and NLC (2017), Garces et al. (2018), Fang et al. (2013), and Yang et al. (1999) Double emulsion No complex equipment, Large particle size, Vancomycin-loaded Gordillo-Galeano and Morasuitable for hydrophilic use of organic SLN, PuerarinHuertas (2018), zur Mühlen et al. (1998), drugs, suitable for solvents, high loaded SLN Garces et al. (2018), thermolabile drugs nanoparticle Nasrollahi et al. (2013), dilution Chen et al. (2001), and Liu et al. (2011) Membrane contractor Simplicity, possible Membrane Vitamin E-loaded SLN Gordillo-Galeano and Moratechnique continuous production obstruction Huertas (2018), Khosa et al. (2018), and Esposito et al. (2012) Phase inversion technique Easy to scale up, no organic Tedious process, Idebenone-loaded SLN Gordillo-Galeano and Morasolvent emulsion instability Huertas (2018), Khosa et al. (2018), Garces et al. (2018), Hsu et al. (2010), and Poonia et al. (2016) Coacervation Low cost, suitable for Possibility of Didodecylmethotrexate- Garces et al. (2018), Lin et al. thermolabile drugs, shape components loaded SLN, (2017), and Kumbhar and Pokharkar (2013) and size can be degradation due to temozolomide-loaded modulated by reaction acidic conditions SLN conditions (continued)

Table 7.1 (Continued) Method

Advantages

Disadvantages

Examples

References

Organic solvent-based approaches

Solvent emulsification evaporation

Suitable for thermosensitive Use of organic drugs, simple equipment, solvents, high small particle size and size nanoparticle distribution, dilution, emulsion reproducibility instability

Solvent Suitable for thermosensitive Use of organic emulsification diffusion drugs, simple equipment, solvents, high small particle size and size nanoparticle distribution, dilution, emulsion reproducibility instability

Solvent injection

Simplicity, fast production, suitable for thermosensitive drugs

NLC, Nanostructured lipid carriers; SLN, solid lipid nanoparticles.

Use of organic solvents

Diclofenac sodiumKhosa et al. (2018), Iqbal loaded SLN, et al. (2017), Garces et al. (2018), Nasrollahi et al. ritonavir-loaded SLN (2013), Poonia et al. (2016), Wang et al. (2013), and Date et al. (2011) Budesonide-loaded Gordillo-Galeano and MoraSLN, 5-fluorouracilHuertas (2018), Khosa et al. (2018), Iqbal et al. loaded (2017), Poonia et al. (2016), Weber et al. (2014), Ngan and Asmawi (2018), and Ezzati Nazhad Dolatabadi et al. (2015) Ciprofloxacin Khosa et al. (2018), Iqbal hydrochloride-loaded et al. (2017), Garces et al. (2018), Poonia et al. SLN, econazole (2016), Hu et al. (2010), nitrate-loaded NLC and Videira et al. (2012)

Recent advances in solid lipid nanoparticles formulation and clinical applications

(DLS) (Gordillo-Galeano and Mora-Huertas, 2018; Das and Chaudhury, 2011). The simultaneous use of both techniques is recommended because they complement each other in the range of particle sizes they determine (Müller et al., 2000). Particle size distribution is characterized by the polydispersity index, which ranges between 0 and 1. Values close to 0 correspond to monodisperse samples, while values close to 1 are indicative of highly polydisperse distributions (Gordillo-Galeano and Mora-Huertas, 2018). These techniques are based on projected surface light scattering effects of the particles, so not spherical shapes or broad particle size distributions can generate misunderstandings of the obtained values. It is recommended to confirm the results using imaging techniques (Gordillo-Galeano and Mora-Huertas, 2018; Müller et al., 2000).

7.5.2 Surface charge Zeta potential values are commonly used to express superficial charge magnitude in aqueous dispersions and can be estimated through the determination of the electrophoretic/electroacoustic mobility. Zeta potential can also be estimated by DLS and LD (Gordillo-Galeano and Mora-Huertas, 2018) being indicative of nanoparticles long-term physical stability. It has been reported that absolute values higher than 30 mV are required to assure stability in terms of pure electrostatic interactions. However, if surfactants providing steric stabilization are used, an absolute value of 20 mV is enough for nanoparticle stabilization (Gordillo-Galeano and Mora-Huertas, 2018; Khosa et al., 2018).

7.5.3 Morphology Nanoparticle morphology can be evaluated by scanning electron microscopy (SEM), transmission electron microscopy (TEM), atomic force microscopy (AFM), and their cryogenic variations (Gordillo-Galeano and Mora-Huertas, 2018). These techniques also provide information about particle size, surface topography, aggregation, and internal structure (Khosa et al., 2018). The most widely used technique is TEM, which usually requires negative staining, followed by SEM that involves drying the sample and coating it with a metallic layer (Gordillo-Galeano and Mora-Huertas, 2018), which can cause alterations of the nanoparticles (Müller et al., 2000). AFM allows sample observation in hydrated state, without coating or staining steps (Gordillo-Galeano and Mora-Huertas, 2018). Image techniques involving cryofixation, as CryoTEM or CryoFESEM (Cryo field emission SEM), can be suitable alternatives, since they allow particle observation in a frozen hydrated state (Khosa et al., 2018).

7.5.4 Degree of crystallinity and polymorphism Crystalline state and polymorphism should be carefully characterized, since they can affect lipid nanoparticle encapsulation efficiency and release profile. It should be

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considered that thermodynamic stability increases and loading capacity decreases in the following order: supercooled melt . α-modification . βʹ-form . β-modification (Khosa et al., 2018). Differential scanning calorimetry (DSC) and XRD are two commonly used techniques to evaluate crystal structure and polymorphic behavior of SLN and NLC (Gordillo-Galeano and Mora-Huertas, 2018; Das and Chaudhury, 2011). DSC technique allows the detection of melting point modifications and melting enthalpies of lipids (Müller et al., 2000; Das and Chaudhury, 2011), providing information about structure, physical state, phase transitions, and interactions between components (Gordillo-Galeano and Mora-Huertas, 2018). XRD allows the characterization and the identification of lipid and drug structures and the prediction of lipid molecules arrangement and phase behavior (Das and Chaudhury, 2011).

7.5.5 Coexistence of different colloidal structures When producing lipid nanoparticles, it is common to obtain dispersions of different coexisting colloidal structures (liposomes, nanoemulsions, micelles, mixed micelles, and supercooled melts) with dynamic phenomena between them, affecting particle stability, DL, and drug release behavior (Khosa et al., 2018). Electron spin resonance and nuclear magnetic resonance are useful noninvasive techniques to study these phenomena together with the existence of oily nanocompartments (multiple-type NLC) (Khosa et al., 2018; Müller et al., 2000; Das and Chaudhury, 2011).

7.5.6 Entrapment efficiency and drug loading DL capacity is defined as the percentage of drug incorporated in the particles regarding to the total nanoparticle weight. EE is defined as the percentage of drug incorporated into the particle with respect to the total amount of drug added (Gordillo-Galeano and Mora-Huertas, 2018; Muchow et al., 2008). It should be noticed that EE percentages can reach high values if a low amount of drug is added, so it is always advisable to pay attention to DL values (Muchow et al., 2008). To determine both parameters, it is necessary to separate the drug from the formulation, for which techniques such as ultrafiltration, centrifugation, or dialysis can be used (Gordillo-Galeano and MoraHuertas, 2018). The quantification of the drug can be determined indirectly, by measuring the free drug in the aqueous phase (Iqbal et al., 2017) or directly, disrupting the lipid matrix with organic solvents (Rouco et al., 2018).

Recent advances in solid lipid nanoparticles formulation and clinical applications

7.6 Drug incorporation models 7.6.1 Drug loading models of solid lipid nanoparticles Two different structural models have been proposed for SLN. The first model corresponds to the α-polymorph form of lipids which maintains a spherical shape after cooling. The second model has been associated to the β or βʹ forms of lipids that lead, after solidification, to a platelet-like layered structure with several creases and borders (Gordillo-Galeano and Mora-Huertas, 2018). The affinity of the drug for the different formulation components determines its distribution within the SLN (Gordillo-Galeano and Mora-Huertas, 2018). Three models have been proposed for SLN of the spherical-shaped structure: homogeneous matrix or homogeneous matrix of solid solution (Fig. 7.2A), drug-enriched shell (Fig. 7.2B), and drug-enriched core (Fig. 7.2C). 7.6.1.1 Homogeneous matrix The active ingredient is molecularly and homogeneously dispersed into the lipid nanoparticle matrix or forming amorphous drug clusters (Souto et al., 2007; Muchow et al., 2008). Drug release occurs by diffusion and degradation of the lipid matrix (Muchow et al., 2008) and a controlled drug release can be achieved (Souto et al., 2007). This type of drug incorporation can be obtained by avoiding heat during formulation process, optimizing the drug lipid ratio in hot techniques (Souto et al., 2007), without using surfactants or alternatively using emulsifiers that cannot solubilize the drug (Müller et al., 2000). Highly hydrophobic drugs with high-crystallization temperatures favors the generation of this model (Gordillo-Galeano and MoraHuertas, 2018).

Figure 7.2 SLN drug incorporation models. (A) Homogeneous matrix. (B) Drug-enriched shell. (C) Drug-enriched core. SLN, Solid lipid nanoparticles.

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7.6.1.2 Drug-enriched shell Nanoparticles present a core shell structure where drug is located in the shell (Muchow et al., 2008). If nanoparticle matrix started to solidify when the repartition process takes place, an accumulation of drug in the nanosystem shell occurs. Drug release profiles of this kind of nanoparticles are characterized by a burst release, due to the drug accumulation in the outer part of the particles (zur Mühlen et al., 1998). This type of drug incorporation is obtained when lipids of high-crystallization temperatures, high-surfactant concentration, low amount of drug, and hot techniques are selected (Souto et al., 2007; Gordillo-Galeano and Mora-Huertas, 2018; Muchow et al., 2008). 7.6.1.3 Drug-enriched core Nanoparticles present a core shell structure where drug is accumulated in the core (Muchow et al., 2008). Drug release profile is expected to follow a prolonged release pattern (Souto et al., 2007; Muchow et al., 2008). This structure is obtained when the recrystallization of the lipid occurs after the precipitation of the drug. This occurs when working at drug concentrations in the melted lipid near saturation so that the crystallization of the lipid occurs after the crystallization of the drug (Souto et al., 2007; Muchow et al., 2008; Müller et al., 2000). Besides these incorporation models, it is also necessary to consider that drug incorporation in the aqueous phase as well as in the nanoparticle interface is also possible, leading to a rapid release of the active ingredient. This phenomenon is associated with β polymorphic forms, which often lead to drug expulsion (Gordillo-Galeano and Mora-Huertas, 2018).

7.6.2 Nanostructured lipid carriers drug loading models Three different structural models of NLC, depending mainly on the type of lipid used, have been described (Souto et al., 2007): imperfect (Fig. 7.3A), amorphous (Fig. 7.3B), and multiple oil-in-solid fat-in-water (O/F/W) (Fig. 7.3C). 7.6.2.1 Imperfect type In this type of NLC, DL capacity is improved by increasing imperfections in the crystal order (Khosa et al., 2018). This can be performed by adding a sufficient amount of oil to the solid lipid to decrease the order of the lipid matrix due to the presence of mixtures of mono, di, and triacylglycerols and fatty acids of different chain lengths (Souto et al., 2007). In this structure, the drug molecules can be included in the gaps between the fatty acid chains of triglycerides in the crystal (Naseri et al., 2015).

Recent advances in solid lipid nanoparticles formulation and clinical applications

Figure 7.3 NLC structural types. (A) Imperfect type. (B) Amorphous type. (C) Multiple type. NLC, Nanostructured lipid carriers.

7.6.2.2 Amorphous type In amorphous NLC, drug expulsion during storage due to β-modification is avoided by the generation of an unstructured amorphous lipid matrix (Khosa et al., 2018). For this purpose, the solid lipid is mixed with special lipids such as isopropyl myristate, hydroxyoctacosanyl hydroxystearate, or dibutyl adipate (Souto et al., 2007; Khosa et al., 2018) that do not recrystallize during the cooling step, creating an amorphous structure (Souto et al., 2007). 7.6.2.3 Multiple oil-in-solid fat-in-water type Multiple-type NLC (Fig. 7.3C) show several tiny oil compartments embedded in the nanoparticles structure that increase DL. The structure reminds the W/O/W emulsions (Naseri et al., 2015). A prolonged release of drug controlled by the solid lipid matrix around the small compartments can be achieved with this model (Khosa et al., 2018). This type of NLC can be obtained using amounts of oil greater than its solubility in the solid lipid (Souto et al., 2007). During the cooling step, the oil droplets reach the miscibility region, precipitate, and remain fixed within the nanoparticles structure through the solidification of the solid lipid (Souto et al., 2007).

7.7 Administration routes Lipid nanoparticles offer a high potential as drug delivery systems through different administration routes such as topical, parenteral, oral, pulmonary, ocular, or intranasal (Khosa et al., 2018; Battaglia and Gallarate, 2012).

7.7.1 Topical administration Dermal and transdermal administration of drugs is an interesting approach for the management of dermatological pathologies, increasing the amount of drug in the site

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of action for a local effect or even looking for a systemic therapy (Sala et al., 2018). The main barrier of the skin is the stratum corneum, which can be crossed via the cell, paracellular, or follicular pathways. In addition to its barrier effect, the skin also has its own metabolism that hinders drug absorption by this route (Sala et al., 2018; Garces et al., 2018). Lipid nanoparticles have certain advantages for topical application: drug protection, modulation of drug release, improved bioavailability, increased skin penetration, and adhesion to the stratum corneum. It is known that after its topical application, SLN and NLC form a hydrophobic film with occlusive properties. As a consequence of the strengthening of the lipid barrier, the loss of transepidermal water is prevented and the skin hydration increases (Sala et al., 2018; Garces et al., 2018). Large lipid nanoparticles ( . 100 nm) can remain on skin surface or accumulate in hair follicles, generating a drug reservoir and releasing active ingredients toward epidermis and dermis, so having a local effect (Sala et al., 2018). Small lipid nanoparticles (,100 nm) can reach blood circulation, having a systemic effect. In addition, its composition gives them the ability to alter the structure of the stratum corneum. The lipids interact with those of the skin and the surfactants act as absorption promoters (Sala et al., 2018; Garces et al., 2018). During the last years, SLN and NLC have been tested in vitro, ex vivo, or in vivo animal models as drug delivery systems to topically administer different drugs to treat skin disorders, chronic wounds or even rheumatoid arthritis (Sala et al., 2018; Garces et al., 2018). Some selected drug used examples are tretinoin (Nasrollahi et al., 2013), acyclovir (Cortesi et al., 2011), methotrexate (Lin et al., 2010), recombinant human epidermal growth factor (Gainza et al., 2015), astragaloside IV (Chen et al., 2013), minoxidil (Wang et al., 2017), or resveratrol (Rigon et al., 2016). The use of lipid nanoparticle formulations has been linked to various therapeutic benefits such as improved efficacy, transdermal drug permeation, and penetration of the stratum corneum, or an increase in the local concentration of the drug, together with a reduction in systemic absorption and the irritation of the skin (Garces et al., 2018). Lipid nanoparticle suspensions can be transformed in semisolid formulations to facilitate topical application, for example, by dispersing the nanosystems in freshly prepared hydrogels or by adding gelling agents to the aqueous phase (nanoemulgels). Moreover SLN and NLC have been widely studied for cosmetic purposes, as is the case of formulations with antiage, antioxidant, ultraviolet blocker, and hydrating properties (Garces et al., 2018).

7.7.2 Parenteral administration Among the advantages of parenteral lipid nanoparticles are their particulate nature, the protection of the drug, the suitability for the encapsulation of lipophilic and

Recent advances in solid lipid nanoparticles formulation and clinical applications

hydrophilic drugs, the potential controlled release of the drug, the cost of the raw materials, and the ease scaling up process (Joshi and Müller, 2009). In addition, the biodegradability of lipids also makes this kind of nanoparticles a good alternative for this purpose(Fang et al., 2013). The main limitation associated with the intravenous administration of these systems is their rapid clearance by the reticuloendothelial system, which can be avoided by modifying them superficially with compounds such as polyethylene glycol or Pluronic F68 (stealth nanoparticles) (Fang et al., 2013). SLN and NLC have been widely studied to treat cardiovascular diseases, rheumatoid arthritis, parasites, inflammatory processes, liver diseases, pain, or cancer, via parenteral route (Joshi and Müller, 2009; Fang et al., 2013). Several formulations have been developed and tested both in vitro and in vivo, as is the case of camptothecin- and paclitaxelloaded SLN (Yang et al., 1999; Chen et al., 2001) or docetaxel-loaded NLC (Liu et al., 2011), among others. Formulations show an improved efficacy together with an increase retention time after parenteral administration. SLN and NLC functionalization with several ligands allows targeting to specific organs, as brain, being useful for treating brain diseases by taking advantage of their ability to overcome the blood brain barrier (Joshi and Müller, 2009; Fang et al., 2013). Bromocriptine- and apomorphine-loaded NLC or clozapine-loaded SLN are some examples of lipid nanoparticles designed for brain targeting using parenteral route (Esposito et al., 2012; Hsu et al., 2010).

7.7.3 Oral administration Oral route is regarded as the preferred administration route due to its many advantages, such as patient compliance, ease of self-administration, painlessness, costeffectiveness, and feasibility for outpatients. The main challenges for drugs oral administration are related to the characteristics of the active molecules (low-aqueous solubility, stability, or permeability) and to the physiological barriers of the gastrointestinal tract: chemical and enzymatic environment, gastrointestinal epithelium (GI), mucosal barrier, presence of multidrug efflux proteins such as glycoprotein P (P-gp) in the membrane of epithelial cells, first-pass metabolism, and variability due to the presence of food (Poonia et al., 2016; Lin et al., 2017). The formulation of drugs in lipid nanoparticles for oral administration confers various benefits such as increased drug solubilization capacity in the GI tract, the protection of labile drugs, the potential controlled release properties, the increase in the residence time, and the potential selective drug delivery (Poonia et al., 2016; Lin et al., 2017). In addition, nanoparticles undergo lymphatic absorption that increases the bioavailability and half-life of drugs that undergo hepatic first-pass effect and are also useful in the case of drugs with toxic metabolites. Drugs that interact with P-gp can also be efficiently delivered through the oral route when administered as lipid nanoparticles (Poonia et al., 2016).

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Lipid nanoparticles can be administered orally as aqueous dispersions or solid dosage forms (powders, capsules, pellets, or tablets) (Battaglia and Gallarate, 2012). They can undergo a digestive process by enzymes, releasing the drug, which crosses the enterocytes and reaches the systemic circulation or alternatively the drugs released together with the products obtained from the digestion of lipids, can form micelles with the bile salts and reach the enterocytes, avoiding the first-pass metabolism by the formation of chylomicrons (Poonia et al., 2016). If nanoparticles avoid digestion process, they could reach portal circulation via paracellular route or lymphatic system through M cells (Poonia et al., 2016), although there is still no consensus about this possibility (Poonia et al., 2016; Lin et al., 2017). Recently the usefulness of SLN has been explored for the administration of different drugs and natural products through the oral route for the treatment of several pathologies, such as cancers, central nervous system diseases, cardiovascular diseases, infections, diabetes, and osteoporosis. They have shown promising results both in vitro and in animal models (Lin et al., 2017). NLC have also been employed to administer orally a wide variety of active molecules such as anticancer (Kumbhar and Pokharkar, 2013), antiinflammatory (Wang et al., 2013), antihypertensive (Ranpise et al., 2014), or antidiabetic drugs (Date et al., 2011), among others. Strategies such as surface modification to improve adhesion and diffusion across the mucus as well as peptide ligands have also been evaluated. Those lipid nanosystems have helped to overcome low solubility, stability, permeability problems or drugs’ high toxicity, and to improve pharmacokinetic parameters of NLC-loaded drugs. Various oils from plant and animal sources have been included in NLC (as liquid lipid) and drug delivery in the oral cavity using lipid nanoparticles included in hydrogels to treat infections has also been studied (Poonia et al., 2016).

7.7.4 Pulmonary administration Drug delivery through the inhalation offers a huge potential for both local and systemic therapy (Weber et al., 2014; Ngan and Asmawi, 2018). Lipid nanoparticles advantages for pulmonary drug delivery include good biocompatibility and biodegradability, deep lung deposition, prolonged mucoadhesion, lung retention and controlled release, improving therapeutic efficacy of drugs, reducing required dose, and increasing dose interval (Weber et al., 2014). SLN and NLC for pulmonary administration should be biocompatible and fulfill aerodynamic requirements that vary depending on whether the therapy is systemic or local. For systemic treatments, lipid particles with 1 3 μm aerodynamic diameters are required to achieve deep lung deposition and drug absorption. For local delivery, the requirements depend on the lung area to be treated (Weber et al., 2014).

Recent advances in solid lipid nanoparticles formulation and clinical applications

Lipid nanoparticles can be administered as suspensions using nebulizers or as dry powders and their size can be modulated as a function of its composition and formulation method (Weber et al., 2014; Ngan and Asmawi, 2018). Systemic treatments via inhalation can be interesting alternatives for peptide and protein delivery, which can be either incorporated into the lipid matrix or adsorbed onto the surface (Weber et al., 2014). In addition, lipid nanoparticles have also shown their potential to deliver other drugs, such as alendronate, to achieve a systemic effect, in this case, for osteoporosis treatment (Ezzati Nazhad Dolatabadi et al., 2015). Regarding local pulmonary therapy, lipid nanoparticle formulations have been studied for the treatment of several diseases such as chronic obstructive pulmonary disease, asthma, lung cancer, cystic fibrosis, or mycosis (Weber et al., 2014; Ngan and Asmawi, 2018). Some examples are epirubicin- and paclitaxel-loaded SLN (Hu et al., 2010; Videira et al., 2012), or beclomethasone dipropionate- (Jaafar-Maalej et al., 2011), tobramicin- (Moreno-Sastre et al., 2016), and itraconazole-loaded NLC (Pardeike et al., 2011). Pulmonary tuberculosis treatment is another goal for lipid nanoparticles (Weber et al., 2014). It has been shown that lipid nanoparticles of size around 1 μm provide an efficient alveolar macrophage targeting (Mu and Holm, 2018). As an example, SLN loaded with rifampicin, isoniazid, and pyrazinamide have shown promising results (Weber et al., 2014). In addition, pulmonary administration of gene vectors, as plasmid-DNA and siRNA, using lipid nanoparticles has also been explored (Weber et al., 2014; Ngan and Asmawi, 2018).

7.7.5 Ocular administration Topical ocular delivery is the most suitable route to treat superficial diseases and pathologies related to the anterior segment of the eye, due to its ease of administration, noninvasive nature, patient compliance, reduced systemic absorption, and avoidance of hepatic first-pass effect. However, the presence of barriers, metabolic (esterases and cytochrome P-450), static (corneal epithelium and stroma, and blood aqueous), or dynamic (lymphatic flow, conjunctival flow, and drainage of tears), hinders drugs administration. Therefore treating the posterior segment of the eye by topical route still constitutes a challenge (Sanchez-Lopez et al., 2017a). SLN and NLC offer certain advantages for the administration of drugs via the eye; they confer protection to drugs against lacrimal enzymes and binding proteins; they are able to overcome the eye’s blood barriers; they allow controlled and prolonged release of drugs; they are biodegradable and biocompatible and they allow to decrease the dosage and therefore reduce side effects (Sanchez-Lopez et al., 2017a,b). During the last years, several SLN and NLC in eye drops or in semisolid hydrogels have been developed for the treatment of pathologies of the anterior and posterior

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segment of the eye as infections, ocular inflammation, glaucoma, cataracts or treatments of age related macular degeneration, which have produced higher bioavailability of the drugs (Sanchez-Lopez et al., 2017b). They have also been shown to be useful for administering genetic material, supplying it to the nucleus and favoring the transfection of the genes into the cells of the retina (Apaolaza et al., 2014). Research shows that lipid nanoparticles are ideal for efficiently administering therapeutic agents in the posterior segment of the eye (Sanchez-Lopez et al., 2017b).

7.7.6 Intranasal delivery Intranasal delivery has been proposed as an alternative to parenteral route due to several advantages such as high surface area, rich blood supply in nasal mucosa, noninvasiveness, and avoidance of hepatic and gastrointestinal metabolism, allowing a quick absorption and side effects and doses reduction (Khosa et al., 2018). However, this route of administration is hampered by the presence of several enzymes, the adenosine triphosphate binding cassette of the nasal cavity and mucociliary clearance. In addition, it only allows the administration of high potency drugs in small volume formulations (Battaglia et al., 2018). The administration of lipid nanoparticles by this route helps to overcome some of those problems. Surface charge (cationic) modified SLN and NLC offer bioadhesion to nasal mucosa, biocompatibility and drug protection from degradation and P-gp efflux proteins activity. Moreover the surfactants included in their composition are able to disrupt tight junctions in mucosal epithelium and improve absorption (Battaglia et al., 2018). SLN have been successfully used as a release system for diclofenac, carvedilol, and flavonoids through intranasal route (Battaglia et al., 2018). In addition, the rich blood supply of respiratory epithelium makes this route adequate for systemic drug delivery, avoiding first-pass effect (Battaglia et al., 2018). During the last years intranasal route has been explored for targeting the central nervous system as it provides direct access from the nose to the brain and to the cerebrospinal fluid through the intra- and extraneuronal pathways, overcoming blood brain barrier (Battaglia et al., 2018; Samaridou and Alonso, 2018). This administration route has a great potential for the treatment of several neurodegenerative diseases such as Alzheimer, Parkinson, Amyotrophic Lateral Sclerosis, and Huntington’s disease (Battaglia et al., 2018). Some examples of drug delivery systems based on lipid nanoparticles studied for neurodegenerative diseases management are rivastigmine-loaded SLN (Shah et al., 2015a), rosmarinic acid-loaded SLN (Bhatt et al., 2015), astaxanthin-loaded SLN (Bhatt et al., 2016), tarenflurbilloaded SLN (Muntimadugu et al., 2016), or chitosan-coated SLN complexed with siRNA and cell penetrating peptides (Rassu et al., 2017). The administration of these drugs by means of lipid nanoparticles through the intranasal route provided different benefits such as increased bioavailability, improved brain biodistribution profiles, higher mucoadhesiveness, and enhanced intracellular transport (Battaglia et al., 2018).

Recent advances in solid lipid nanoparticles formulation and clinical applications

7.8 Solid lipid nanoparticles and nanostructured lipid carriers case studies in humans for medical applications The main commercial application of lipid nanoparticles is nowadays focused on the cosmetics, food/nutrition, and nutraceuticals fields (Danaei et al., 2018). The use of lipid nanoparticles is well stablished for cosmetic uses with more than 25 formulations already marketed (Pardeike et al., 2009). Despite the huge potential of SLN and NLC as drug delivery platforms, no current treatments based on this strategy are in the clinic (Cipolla et al., 2014). However, several preclinical and clinical studies have been performed using these nanoparticulated systems for medical applications. While the preclinical evaluation of these formulations were already described in the previous section, the clinical studies performed based on SLN and NLC are described in this section, categorized by the administration route selected.

7.8.1 Topical administration (skin and mucosa) SLN have been widely included in cosmetic and pharmaceutical industrial products for the skin care without the need of clinical trials for their commercialization approval (Lakshminarayanan et al., 2018; Kaul et al., 2018). Their ability to modulate skin penetration make them really useful for this purpose (Souto and Müller, 2008). Moreover the treatment of numerous acute or chronic dermal and mucosal pathologies with lipid nanoparticles has also been explored. This administration route usually requires their combination with other pharmaceutical formulation forms such as gels or creams in where the nanoparticulate systems are incorporated to facilitate their administration. The efficacy and safety of fluconazole-loaded SLNs incorporated on a gel for the local therapy of skin fungal infections was assessed. Patients diagnosed with Pityriasis versicolor were topically treated twice daily for 4 weeks with one of the following groups: (1) 1% Carbopol 934 gel containing SLNs of 10% Compritol 888 ATO 1 0.5% Cremophor RH40 1 1% Fluconazole; (2) 1% Carbopol 934 gel containing SLNs of 10% Precirol ATO5 1 0.5% Poloxamer 407 1 1% Fluconazole, or (3) commercially available product Candistan (Clotrimazole 1%). The final drug concentration used was 8 9 mg/1 per gram of gel. Treatments were applied twice a day during 4 weeks. The percentage of cure was significantly improved by the treatment with the gels containing SLN (groups 1 and 2) when compared to the commercially available product. Moreover no adverse reactions were reported in any of the groups neither during treatment nor afterward (El-Housiny et al., 2017). SLN have also been used for the topical treatment of recurrent condyloma acuminatum. For this purpose, SLN were loaded with podophyllotoxin and embedded on a gel. Patients with recurrent condyloma acuminatum were treated either with standard

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podophyllotoxin gel or with podophyllotoxin-SLN gel. The treatment with the nanoparticulated formulation reached similar wart clearance than the standard treatment but decreased recurrence rates and adverse effects (Xie et al., 2007). Another approach used for topical treatment has been focused on the development of mucoadhesive dosage forms containing curcumin-loaded SLN for the treatment of precancerous lesions in buccal mucosa. In this regard, SLN have shown an excellent loading capacity and a more controlled release when compared to the drug dispersed on the mucoadhesive gel. Moreover curcumin-SLN also showed higher retention and penetration deep into the mucosal layers in an “in vivo” model. Patients diagnosed with erythroplakia were treated with either a conventional mucoadhesive gel or with the developed mucoadhesive formulation containing curcumin-loaded SLN at a dose of 6 mg curcumin per day for 6 weeks. Clinical evaluation consisted on pain index and lesion size. The incorporation of curmumin-loaded SLN into the formulation was able to significantly reduce the pain and size of the lesion. As a result, SLN were able to successfully deliver curcumin in a stable form enhancing its activity. The incorporation of curcumin-SLN into mucoadhesive gel provided an efficient approach for oral mucosal targeting (Hazzah et al., 2016). Even though SLN are usually included into gels to facilitate topical administration, their nature also allows them to be incorporated into similar pharmaceutical formulation forms as creams (Lakshminarayanan et al., 2018). On this sense, clobetasol propionate was loaded in SLN and they were dispersed (amount equivalent of 0.05 g of drug) on a topical cream. The loading capacity of the SLN reached 35% even though the formulation selected for the clinical assessment was the one incorporating 6% of drug based on the size of the SLN. Chronic eczema patients were treated with either clobetasol standard cream or clobetasol-SLN cream for 6 weeks. The therapeutic response was significantly improved on those patients treated with clobetasol-SLN cream in terms of degree of inflammation and itching when compared to the control group (Kalariya et al., 2005). Similar approaches for local drugs administration have been also developed and clinically tested for the second generation of lipid nanoparticles. NLC have been used for the treatment of acne vulgaris by loading spironolactone on them. In this study, NLC were obtained by probe-ultrasonication and composed by stearic acid (solid lipid), oleic acid (liquid lipid), Span 80 (lipophilic surfactant), and spironolactone (SP). NLC were dispersed on a 1% Carbopol gel containing methyl paraben as preservative. This formulation was clinically tested for acne treatment in comparison to an alcoholic gel containing free spironolactone. With over 8 weeks of treatment, patients used 60 g of either the gel containing 10 mg SP-NLC per gram of gel or alcoholic gel 50 mg SP per gram of gel. The treatment with SP-NLC promoted a reduction in the acne total lesion score together with a decrease in noninflammatory lesions when compared to the base line, similarly to the alcoholic gel after 8 weeks of treatment. Therefore the use of SP-NLC has shown a good therapeutic effect on mild to moderate acne vulgaris and it was well tolerated (Kelidari et al., 2016). On a similar study, acitretin-loaded nanostructured lipid carriers (Act-NLC) were obtained by the combination of the following components: oleic acid (liquid lipid),

Recent advances in solid lipid nanoparticles formulation and clinical applications

Precirol ATO 5 (solid lipid), and Tween 80 (surfactant) containing 5% of drug. Lyophilized Act-NLC at an equivalent drug amount of 150 mg were incorporated to a Carbopol 934P based gel. Psoriasis patients were then treated with either the standard acitretin gel twice daily or Act-NLC gel once a day for 4 weeks. Patients treated with the NLC-based formulation presented a reduction in erythema and a significant reduction in scaling, indicating moderate to excellent improvement in the disease symptoms when compared to the standard treatment (Agrawal et al., 2010).

7.8.2 Oral administration Although several SLN and NLC formulations have been designed and tested in vitro or in preclinical studies for systemic administration, clinical studies for these systems are still limited. The high costs of the experiments together with their unexplored systemic side effects could be the reason of the lack of clinical trials on this route (Lin et al., 2017). To clarify this point, the safety and tolerability of oral administered curcumin-loaded SLN have been assessed on healthy volunteers and cancer patients. Nanoparticles containing curcumin in the range of 20% 30% of total formulation were administered orally as a single dose in a capsule of 2000 mg (containing 400 600 mg of curcumin), 3000, or 4000 mg to late-stage osteosarcoma patients. The pharmacokinetic analysis showed high plasma curcumin concentrations and dose-related AUCs reaching the plasma concentration peak after 3.5 hours of administration. No adverse events were reported in either healthy volunteers or osteosarcoma patients (Gota et al., 2010). Despite the promising clinical results of the above studies, it should be stated that all the clinical trials included presented a relatively small number of patients and these are treated for a short period of time, 8 weeks is the longest time point selected. Therefore there is a long way to go for the establishment of SLN and NLC as a realistic clinical treatment alternative to conventional therapies. However, taking into account the improvement in their formulation technology over the last few years and the increased number of patented NLC-based formulations, these systems represent a promise for the pharmaceutical market which is expected to promote an increase in the number of clinical trials performed in the near future (Beloqui et al., 2016).

Self-assessment questions 1. What is the main difference between solid lipid nanoparticles (SLN) and nanostructured lipid carriers (NLC)? 2. What are the main factors related to lipids which should be considered during excipient selection?

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3. Which polymorphic transitions can occur in lipid nanoparticles during storage? How do they affect nanoparticle properties? How can polymorphic behavior be modulated? 4. What are the functions performed by emulsifiers in the case of lipid nanoparticles? 5. Which preformulation studies would you perform for the development of a SLN formulation? And in the case of NLC? 6. Which are the possible drug incorporation models in SLN? Which of them would result in a burst-type drug release? 7. Which are the main advantages of NLC over SLN? 8. How could an amorphous type NLC be obtained? And a multiple type NLC? Which release profile could be expected from this last type? 9. Which are the main formulation techniques of SLN and NLC? 10. Which type of organic solvent should be used in solvent emulsification evaporation method? And in the case of solvent injection? 11. Which formulation techniques can be used in the case of thermosensitive drugs? 12. Which technique(s) should be used for particle size characterization of a lipid nanoparticle formulation? And for crystalline state characterization? 13. Which techniques should be used to evaluate particle morphology in the case of samples sensitive to staining or desiccation? 14. Which are the main advantages derived from the use of lipid nanoparticles for drug delivery through the topical route? 15. How could lipid nanoparticle suspensions be transformed into a semisolid formulation for topical application? 16. Which strategy could be employed to overcome clearance by reticuloendothelial system in lipid nanoparticles administered through the parenteral route? 17. Which are the main advantages derived from the use of lipid nanoparticles for drug delivery through the oral route? 18. Which requirement should a lipid nanoparticle formulation fulfill to achieve a systemic effect through the pulmonary route? 19. Is there any SLN- or NLC-based formulation in the market at the present time? 20. What are the administration routes in which SLN or NLC have reached clinical trials? 21. Which administration route could show a shorter time to market for SLN and NLC with therapeutic purposes?

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Khosa, A., Reddi, S., Saha, R.N., 2018. Nanostructured lipid carriers for site-specific drug delivery. Biomed. Pharmacother. 103, 598 613. Kim, J.K., Park, J.S., Kim, C.K., 2010. Development of a binary lipid nanoparticles formulation of itraconazole for parenteral administration and controlled release. Int. J. Pharm. 383, 209 215. Kovacevic, A., Savic, S., Vuleta, G., Müller, R.H., Keck, C.M., 2011. Polyhydroxy surfactants for the formulation of lipid nanoparticles (SLN and NLC): effects on size, physical stability and particle matrix structure. Int. J. Pharm. 406, 163 172. Kumar, V.V., Chandrasekar, D., Ramakrishna, S., Kishan, V., Rao, Y.M., Diwan, P.V., 2007. Development and evaluation of nitrendipine loaded solid lipid nanoparticles: influence of wax and glyceride lipids on plasma pharmacokinetics. Int. J. Pharm. 335, 167 175. Kumbhar, D.D., Pokharkar, V.B., 2013. Engineering of a nanostructured lipid carrier for the poorly water-soluble drug, bicalutamide: physicochemical investigations. Colloids Surf. A Physicochem. Eng. Asp. 416, 32 42. Lakshminarayanan, R., Ye, E., Young, D.J., Li, Z., Loh, X.J., 2018. Recent advances in the development of antimicrobial nanoparticles for combating resistant pathogens. Adv. Healthcare Mater. 7, e1701400. Lin, Y.K., Huang, Z.R., Zhuo, R.Z., Fang, J.Y., 2010. Combination of calcipotriol and methotrexate in nanostructured lipid carriers for topical delivery. Int. J. Nanomed. 5, 117 128. Liu, D., Liu, Z., Wang, L., Zhang, C., Zhang, N., 2011. Nanostructured lipid carriers as novel carrier for parenteral delivery of docetaxel. Colloids Surf. B Biointerf. 85, 262 269. Lin, C.H., Chen, C.H., Lin, Z.C., Fang, J.Y., 2017. Recent advances in oral delivery of drugs and bioactive natural products using solid lipid nanoparticles as the carriers. J. Food Drug. Anal. 25, 219 234. Martins, S., Sarmento, B., Ferreira, D.C., Souto, E.B., 2007. Lipid-based colloidal carriers for peptide and protein delivery liposomes versus lipid nanoparticles. Int. J. Nanomed. 2, 595 607. Mehnert, W., Mader, K., 2001. Solid lipid nanoparticles: production, characterization and applications. Adv. Drug. Deliv. Rev. 47, 165 196. Moreno-Sastre, M., Pastor, M., Esquisabel, A., Sans, E., Vinas, M., Fleischer, A., et al., 2016. Pulmonary delivery of tobramycin-loaded nanostructured lipid carriers for Pseudomonas aeruginosa infections associated with cystic fibrosis. Int. J. Pharm. 498, 263 273. Mu, H., Holm, R., 2018. Solid lipid nanocarriers in drug delivery: characterization and design. Expert. Opin. Drug. Deliv. 15, 771 785. Muchow, M., Maincent, P., Muller, R.H., 2008. Lipid nanoparticles with a solid matrix (SLN, NLC, LDC) for oral drug delivery. Drug. Dev. Ind. Pharm. 34, 1394 1405. zur Mühlen, A., Schwarz, C., Mehnert, W., 1998. Solid lipid nanoparticles (SLN) for controlled drug delivery drug release and release mechanism. Eur. J. Pharm. Biopharm. 45, 149 155. Müller, R.H., Mader, K., Gohla, S., 2000. Solid lipid nanoparticles (SLN) for controlled drug delivery a review of the state of the art. Eur. J. Pharm. Biopharm. 50, 161 177. Müller, R.H., Radtke, M., Wissing, S.A., 2002. Nanostructured lipid matrices for improved microencapsulation of drugs. Int. J. Pharm. 242, 121 128. Müller, R.H., Shegokar, R., Keck, C.M., 2011. 20 years of lipid nanoparticles (SLN and NLC): present state of development and industrial applications. Curr. Drug. Discov. Technol. 8, 207 227. Muntimadugu, E., Dhommati, R., Jain, A., Challa, V.G., Shaheen, M., Khan, W., 2016. Intranasal delivery of nanoparticle encapsulated tarenflurbil: a potential brain targeting strategy for Alzheimer’s disease. Eur. J. Pharm. Sci. 92, 224 234. Naseri, N., Valizadeh, H., Zakeri-Milani, P., 2015. Solid lipid nanoparticles and nanostructured lipid carriers: structure, preparation and application. Adv. Pharm. Bull. 5, 305 313. Nasrollahi, S.A., Abbasian, A.R., Farboud, E.S., 2013. In vitro comparison of simple tretinoin-cream and cream loaded with tretinoin-SLN. J. Pharm. Technol. Drug. Res. 2, 13. Ngan, C.L., Asmawi, A.A., 2018. Lipid-based pulmonary delivery system: a review and future considerations of formulation strategies and limitations. Drug. Deliv. Transl. Res. 8, 1527 1544. Pardeike, J., Hommoss, A., Müller, R.H., 2009. Lipid nanoparticles (SLN, NLC) in cosmetic and pharmaceutical dermal products. Int. J. Pharm. 366, 170 184.

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Triplett, Ii. M.D., Rathman, J.F., 2009. Optimization of β-carotene loaded solid lipid nanoparticles preparation using a high shear homogenization technique. J. Nanopart. Res. 11, 601 614. Videira, M., Almeida, A.J., Fabra, A., 2012. Preclinical evaluation of a pulmonary delivered paclitaxelloaded lipid nanocarrier antitumor effect. Nanomed. Nanotechnol. Biol. Med. 8, 1208 1215. Vitorino, C., Carvalho, F.A., Almeida, A.J., Sousa, J.J., Pais, A.A., 2011. The size of solid lipid nanoparticles: an interpretation from experimental design. Colloids Surf. B Biointerf. 84, 117 130. Vivek, K., Reddy, H., Murthy, R.S.R., 2007. Investigations of the effect of the lipid matrix on drug entrapment, in vitro release, and physical stability of olanzapine-loaded solid lipid nanoparticles. AAPS Pharm. Sci. Tech. 8. Wang, Q., Cheng, H., Zhou, K., Wang, L., Dong, S., Wang, D., et al., 2013. Nanostructured lipid carriers as a delivery system of biochanin A. Drug. Deliv. 20, 331 337. Wang, W., Chen, L., Huang, X., Shao, A., 2017. Preparation and characterization of minoxidil loaded nanostructured lipid carriers. AAPS Pharm. Sci. Tech. 18, 509 516. Weber, S., Zimmer, A., Pardeike, J., 2014. Solid lipid nanoparticles (SLN) and nanostructured lipid carriers (NLC) for pulmonary application: a review of the state of the art. Eur. J. Pharm. Biopharm. 86, 7 22. Weiss, J., Decker, E.A., McClements, D.J., Kristbergsson, K., Helgason, T., Awad, T., 2008. Solid lipid nanoparticles as delivery systems for bioactive food components. Food Biophys. 3, 146 154. Wilde, P., Mackie, A., Husband, F., Gunning, P., Morris, V., 2004. Proteins and emulsifiers at liquid interfaces. Adv. Colloid Interf. Sci. 108-109, 63 71. Xie, F.M., Zeng, K., Chen, Z.L., Li, G.F., Lin, Z.F., Zhu, X.L., et al., 2007. Treatment of recurrent condyloma acuminatum with solid lipid nanoparticle gel containing podophyllotoxin: a randomized double-blinded, controlled clinical trial. Nan fang. yi ke da xue xue bao (J. South. Med. Univ.) 27, 657 659. Yang, S.C., Lu, L.F., Cai, Y., Zhu, J.B., Liang, B.W., Yang, C.Z., 1999. Body distribution in mice of intravenously injected camptothecin solid lipid nanoparticles and targeting effect on brain. J. Control. Release 59, 299 307.

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CHAPTER EIGHT

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants Marilena Vlachou and Angeliki Siamidi Section of Pharmaceutical Technology, Department of Pharmacy, School of Health Sciences, National and Kapodistrian University of Athens, Athens, Greece

8.1 Introduction The oral route is the most commonly used route for facile drug administration not only due to patient acceptance, but also due to the gastrointestinal physiology. The development of modified release dosage forms over the past decades offered significant advances in the area of drug delivery. As a general term, modified release systems are all drug-delivery systems that are deviating from the conventional release pattern of the immediate release systems. They are developed to alter the drug’s absorption and/or site of release to attain possible therapeutic benefits, like improved efficacy and reduced adverse events, increased patient’s convenience and compliance, optimized clinical performance, greater selectivity of activity, and so forth (Qiu and Lee, 2017). Moreover the development of a modified release product may also offer commercial advantages, such as market expansion and increased cost effectiveness. To achieve the predetermined plasma profile that will not only reduce the dose frequency, but will also improve the efficacy-to-safety ratio, scientists alter the drug release rates by adjusting the physical, chemical, and/or biological components of the delivery systems. It is well known that many bodily functions (hormones fluctuations, gastric secretion, etc.) and diseases (bronchial asthma, hypertension, rheumatitis, etc.) follow a circadian rhythm. Chronotherapeutics aim at timing the administration of a programmed release system so that the therapeutic plasma concentration can be obtained at the optimal time with reference to the condition (e.g. asthmatic attacks during early morning hours) (Qiu and Lee, 2017). Another aim of modified release systems is to maintain the concentration of the therapeutic agent in the blood or in target tissues at a desired value. Therefore an initial (a fraction of the dose) rapid release (burst release) may be followed by a maintenance supply of effective drug

Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00008-4

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concentration level. In all situations it is necessary to have full knowledge of the drug release transport mechanisms and be able to predict the drug release kinetics.

8.1.1 Terminology used to describe modified release systems Modified release systems are formulations altered to provide any other kind of release other than immediate. They can be categorized into the following (Collett and Moreton, 2001): • Delayed release systems indicate that the drug is not being released immediately following administration, but at a later time (e.g., enteric-coated tablets, pulsatilerelease capsules) • Repeat action systems indicate that an individual dose is released fairly soon after administration, and the second or third doses are subsequently released at later stages. • Prolonged action/release systems indicate that the drug is provided for absorption over a longer period of time than for a conventional dosage form. However, there is an implication that onset is delayed because of an overall slower release rate from the dosage form. • Sustained action/release systems indicate that an initial dose of the drug sufficient to provide a therapeutic dose is released soon after administration, and then a gradual release over an extended period can be observed. • Extended release systems indicate that the drug is released slowly, so that plasma concentrations are maintained at a therapeutic level for a prolonged period of time, usually between 8 and 12 hours. • Controlled release systems indicate that the drug is released at a constant rate and provide plasma concentrations that remain invariant with time.

8.1.2 Advantages and limitations of modified release systems The widespread proliferation of modified release systems is due to their significant advantages over immediate release systems (Tiwari and Rajabi-Siahboomi, 2008). Specifically: • Modified release systems contribute to maintaining the concentration of the active substance in the blood at therapeutic levels for a longer period of time compared to the levels achieved by single or repeated administration of immediate release formulations. • Fluctuation of levels of the active substance in the blood is reduced, thus side effects and toxicity, which are mainly due to the sharp increases in concentrations can be avoided. • By modified release formulations administration, the patient receives a lower number of doses, which results not only in a reduction in toxicity but also has a wider economic health benefit. In contrast, the underlying limitations of modified release systems can be summarized to the following (Tiwari and Rajabi-Siahboomi, 2008):

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

• Possible risk of accumulation of the active substance. • Delay in the therapeutic effect onset, due to the longer time required to achieve therapeutic blood concentrations. • A decrease in the bioavailability of certain substances may be observed. • There is a restriction on substances, which can be formulated in modified release formulations. Some of the limiting factors are the pharmacokinetic properties (the short half-life requirement), the required dose (which must be low), and the aqueous solubility of the substance (extremely water-soluble substances are difficult to be released at the desired rate).

8.1.3 Historical review of modified release systems The evolution of the modified release methods can be divided into three time phases. The first period, from 1950 to 1970, is related to formulations based on hydrophobic polymers and waxes to prolong the pharmacological action and reduce the dosage. During the second period, from 1970 to 1990, the emphasis was placed on the nonclassical kinetics and the drug targeting in specific sites of the body. In the 1980s and early 1990s, the rapid advances in biotechnology and molecular biology have promoted research into the modified release of drugs by highlighting the optimization of pharmaceutical formulations (Ding et al., 2005). The progress in peroral modified release technology is largely due to the significant improvements in industrial equipments, particularly in coating machines, which have not only led to efficient and controlled production processes, but also to the development of improved biocompatible and biodegradable polymeric materials to modify the drug release rate (Charman and Charman, 2003). Significant efforts have been made over the past decades to develop safe nanomaterial carriers suitable for the development and production of pharmaceutical drug delivery systems.

8.1.4 Formulation of modified release dosage forms When formulating a modified released drug-delivery system for peroral drug delivery, it is essential to consider some factors that may influence the design strategy. These factors include the physiology of the gastrointestinal tract (resident time, pH, bacteria with enzymatic action, etc.), the physicochemical properties of the drug (aqueous solubility and stability, pka, partition coefficient, etc.), the design of the appropriate dosage form (single/multiple unit dosage forms, insoluble/hydrophilic matrices, technical feasibility etc.), the drug release mechanism (dissolution or diffusion), the particular disease factors (presence/absence of pharmacological rationale), and the biological properties of the drug (large dose size, permeability values, half-life, poor colonic permeability, etc.) (Qiu and Lee, 2017; Collett and Moreton, 2001).

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8.1.5 Classification of modified release systems according to the drug release rate mechanism Modified release systems can be classified according to the drug release rate mechanism to the following categories and subcategories: • Diffusion-controlled systems, where the two main types of systems through which the release of encapsulated substances is determined by diffusion mechanisms are: • Reservoir systems and • Monolithic systems, that can be considered as two different groups: • Hydrophilic matrix systems and • Hydrophobic matrix systems. • Solvent-controlled systems, which depending on the penetration mechanism, can be divided into two groups: • Osmotically controlled systems and • Swelling-controlled systems. • Chemically controlled systems, where the chemical reaction that affects the rate of release of the active substance may be either hydrolysis or an enzymatic reaction. These systems can be divided into two categories: • Systems where the drug release is controlled by polymer erosion and • Systems where the drug is bonded to the polymer, that can be considered as two different groups: • Systems where the drug is chemically bonded to the polymer. • Systems where copolymerization of the active substance or its products is achieved.

8.2 Mathematical models for drug release The use of mathematical equations that describe the drug release in function of time is very important when designing a pharmaceutical formulation or evaluating the drug release process either in vitro or in vivo. The mathematical models aim to predict the amount and type of active drug, excipients, size, and shape of the delivery system that will achieve the desired drug release profile. The evaluation of drug release kinetics with mathematical functions ensures the optimal design of the pharmaceutical formulation and provides understanding upon the release mechanism through experimental verification. The main model-dependent methods that describe the release kinetic are the zero-order, first-order, Higuchi, Hixson
Crowell, Korsmeyer
Peppas, and Weibull equations and are depicted in the following section.

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

8.2.1 Zero-order kinetics The dissolution of an active pharmaceutical ingredient contained in a nondisintegrating dosage form that releases the drug in a very slow rate (considering that there are no changes in the equilibrium conditions), can be represented by the following equation: W0 2 Wi 5 Kt where W0 is the initial mass of active pharmaceutical ingredient in the dosage form, Wi is the remaining mass in the dosage form, on time t, and K is the constant of proportionality. This can be simplified by dividing with W0, to the following equation: fi 5 K0 t where fi 5 1 2 (Wi/W0) is the fraction of the active pharmaceutical ingredient dissolved during the time t and K0 is the constant of apparent velocity of dissolution. In this case, the graphical representation of this equation results in a straight line, in which the angular coefficient corresponds to K0. Therefore for zero-order kinetics, the drug release is simply a function of time and the process takes place at a constant rate independent of the active pharmaceutical ingredient concentration. In terms of concentration the previous equation can be described as follows: Ct 5 C0 1 K0 t where Ct is the amount of the active pharmaceutical ingredient released during the time t, C0 is the initial concentration of the active pharmaceutical ingredient released (generally C0 5 0), and K0 is the zero-order constant. This equation can be used to describe various drug-delivery systems (e.g., transdermal, slow release matrix, coated, and osmotic systems). Zero-order kinetics can be achieved in peroral drug delivery by tablet coating or osmotic systems. When the tablet coat is permeable to both the active pharmaceutical ingredient and the gastrointestinal fluids, after swallowing, the core tablet becomes hydrated and the active pharmaceutical ingredient dissolves until it reaches its saturation concentration or solubility. The active pharmaceutical ingredient is released through the membrane to the gastrointestinal fluids with a constant release rate due to its concentration saturation in the tablet core. When this concentration falls below saturation, the drug release rate will decay to zero. Consequently osmotic systems (tablets or capsules consisting of a core of an active pharmaceutical ingredient surrounded by a membrane, which is permeable to gastrointestinal fluids but not to the active pharmaceutical ingredient) can also provide zero-order release kinetics. In this case, upon administration, gastrointestinal fluids pass into the tablet core through the semipermeable membrane and push the active pharmaceutical

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ingredient at a constant rate out of the system through a small hole that is drilled onto the membrane. When drug concentration falls below its solubility, the rate of this osmotic pumping will decay. These drug-delivery systems that display zero-order release kinetics are suitable for prolonged drug release (Costa and Lobo, 2001).

8.2.2 First-order kinetics The phenomenon of dissolution of a solid particle in a liquid medium implies a surface action and can be described by the Noyes and Whitney equation: dc 5 K ðCs 2 C Þ dt where C is the concentration of solute at the time t, Cs is the equilibrium solubility at the temperature of the process, and K is the first-order proportionality constant. This equation can be modified to include the surface variable S: dc 5 K1 S ðCs 2 C Þ dt where K1 is a new constant of proportionality. Using Fick’s first law, it is possible to obtain a relationship for the constant K1: K1 5

D Vh

where D is the diffusion coefficient in the release medium, V is the volume of the liquid release medium, and h the thickness of the diffusion layer. Hixson and Crowell proposed a modification of the Noyes and Whitney equation: dW 5 KS ðCs 2 C Þ dt where W is the amount of drug soluble on time t, dW/dt is the ratio of the solute passing to the solution on time t, S is the surface of dissolution, and K is the constant of proportionality. Multiplying both terms of the equation by V, considering that K 5 K1V and k 5 D/h, the equation can be rewritten as follows: dW KS 5 ðVCs 2 W Þ 5 K ðVCs 2 W Þ dt V where k 5 k1S. When a therapeutic system maintains a constant area under sink conditions, it is possible to use the last equation and, after integration, it results as follows: W 5 VC s ð1 2 e2kt Þ

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

This equation can be transformed by applying decimal logarithms in both terms: log ðVC s 2 W Þ 5 log VC s 2

kt 2:303

However, the term VCs is equal to the mass Q, and the last equation can be rewritten as: log Q1 5 log Q0 1

k1 t 2:303

where Q1 is the amount of active agent released on time t, Q0 is the initial amount of drug dissolved, and K1 is the first-order constant. This equation corresponds to a linear function, and the graph of Napierian or decimal logarithm of the mass released by the drug will result in a straight line, with angular coefficient K1/2.303 and linear coefficient equal to log Q0. In conclusion, when a drug-delivery system follows first-order release kinetics, it implies that the change in drug concentration in relation to time is dependent only on the active pharmaceutical ingredient concentration (Costa and Lobo, 2001).

8.2.3 Higuchi model To understand drug release, Higuchi developed mathematical relationships for drug particles dispersed in homogeneous matrix that act as a diffusing medium (Higuchi, 1961, 1962, 1963). The dissolution of a lipophilic drug in a homogeneous, planar matrix can be described from the following equation: pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi f1 5 Q 5 Dð2C 2 Cs ÞCs t where Q is the amount of drug released on time t by area unit, C is the initial amount of drug contained in dosage form, Cs is the drug solubility in the matrix medium, and D is the diffusion constant in the matrix medium. This equation is valid over the time of dissolution, except when the drug release levels tend to saturate into the liquid medium contained in the matrix. Despite the conditions adopted (lipophilic, homogeneity, and planar form), this model is suitable for other types of dosage forms. To study the dissolution profile from heterogeneous matrix systems, where the drug concentration in the matrix is lower than its solubility and the drug is released through the matrix pores, Higuchi transformed the equation to the following: rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Dε f1 5 Q 5 ð2C 2 εCs ÞCs t τ where Q is the amount of drug released on time t by area unit, ε is the porosity of matrix, τ is the capillary tortuosity factor, C is the initial amount of drug contained in

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the dosage form, Cs is the solubility of the active pharmaceutical ingredient in the matrix medium, and D is the diffusion constant in the matrix medium. This model assumes no important alteration of matrix structure during its contact with liquid media. In 1962 Higuchi proposed the following equation that can be applied when the drug is dissolved from a saturated solution to a matrix system (Higuchi, 1962): rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Dt f1 5 Q 5 2C0 ε τπ where C0 is the solution concentration. This equation can be rewritten as follows:

pffi f1 5 Q 5 KH t

where KH is the Higuchi release constant. This equation is the simplified Higuchi model, which relates the concentration of the drug to the square root of time, representing a linear function. When Higuchi model is used, the following assumptions have to be considered (Siepmann and Peppas, 2012): • The matrix system contains the initial drug concentration which is much higher than the solubility of the drug. • The diffusion is unidirectional, because the edge effects are negligible. • The thickness of the dosage form is greater than the size of the drug molecules. • The swelling or dissolution of the matrix is negligible. • The drug diffusion is constant. • The perfect sink conditions are attained in the release environment.

8.2.4 Hixson
Crowell model In 1931 Hixson and Crowell realized that a group of particles’ regular area is proportional to the cubic root of its volume (Hixson and Crowell, 1931). Using this relationship, they proposed the following equation: ffiffiffiffiffi p p ffiffiffiffiffi 3 w0 2 3 w t 5 K s t where W0 is the initial amount of the drug in the dosage form, Wt is the amount of the drug remaining in the system at time t, and Ks is the constant of incorporation, relating the surface and the volume. This equation applies to dosage forms such as tablets, in which dissolution occurs in plane areas parallel to the surface area of the dosage form, allowing the surface to decrease proportionally over time but the geometrical characteristics of the form to

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

maintain constant. If the drug concentration and density as well as the spherical particles are considered, the equation can also be described as the following: ffiffiffiffiffi 0p 3 N DC s t K ffiffiffiffiffi p p ffiffiffiffi ffi 3 w0 2 3 w t 5 δ where N is the number of particles, K0 is a constant that relates surface, form, and particle density, D is the diffusion coefficient, Cs is the solubility at the equilibrium (saturation concentration) at the temperature of process, and δ is the thickness of the diffusion layer. The surface factors of cubic or spherical particles should be kept constant if the dissolution is constant over the entire system and the dosage form dissolves in an equal manner by all sides. In case of the particles being irregular, this behavior is different and the equation cannot be applied. Dividing the initial equation by the cubic root of W0, it is possible to simplify: p ffiffiffiffiffiffiffiffiffiffiffi 3 1 2 ft 5 1 2 K β t where ft 5 1 2 (Wt/W0) and represents the fraction of drug dissolved at time t and Kβ is a release constant. This equation represents a linear function when the cubic root of the unreleased fraction of the drug is related with time in a graph. This can occur only when the equilibrium conditions are not modified and the surface of the particle decreases proportionally over the period of time. When this model is used, it must be assumed that the drug release is limited by dissolution velocity and not by diffusion, which can occur through the polymeric matrix (Costa and Lobo, 2001).

8.2.5 Korsmeyer
Peppas model Korsmeyer et al. (1983) and Ritger and Peppas (1987a,b) developed a semiempirical model, establishing the exponential relationship between the release and the time: f1 5

Mi 5 Kt n MN

where f1 is the amount of drug released, Mi is the amount of drug released over time t, MN is the amount of drug released as time approaches infinity (  amount of drug contained in the dosage form), K is the release constant (incorporating structural modifications and geometrical characteristics of the system), and n is the release exponent (related to the drug release mechanism) in function of time t. NB: To determine the value of the exponent n, it is suggested to use the first 60% of the release curve. Also it must be assumed that the drug release occurs in just one

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direction and that the relationship between the diameter and thickness of the system is at a maximum of 1:10 (Ritger and Peppas, 1987a,b; Peppas and Sahlin, 1989). This model, also called the power law model, was developed mainly for the drug released from a polymeric matrix, such as a hydrogel, and it is very useful when the release mechanism is not known or when more than one type of drug releases are involved. As mentioned earlier, the value of n is related to the drug release mechanism and can be classified to a Fickian model (Case I) or non-Fickian models (Case II, Anomalous Case and Super Case II) (Table 8.1). In the Fickian model (Case I), the drug release is governed by diffusion as the solvent penetration rate is much greater than the polymeric chain relaxation. This kind of Fickian diffusion occurs in polymers with a glass transition temperature (Tg) lower than the temperature of the environment. Non-Fickian models are normally found in glassy polymers, when the environmental temperature is less than the Tg and include Case II, Anomalous Case, and Super Case II transport mechanisms. In anomalous transport, the mechanism of drug release is governed by both diffusion and swelling. The polymeric chains are rearranged slowly while the diffusion process takes place very quickly leading to the timedependent anomalous effects. Transport Case II models are characterized by a rapid penetration of the solvent to the center of the matrix. The drug release rate corresponds to zero-order release kinetics and the drug release mechanism is governed by the swelling and relaxation of the polymeric chains. At the end of Case II t806-ransport, a fast increase of solvent diffusion may sometimes be observed and the transport mechanism changes to Super Case II transport, as the forces increase by the swollen gel in the glassy nucleus. Table 8.1 Release exponent n representing the drug release mechanism from polymeric matrices with different geometries. Release mechanism model Geometry of the system Release exponent n

Fickian diffusion Case I Anomalous transport

Case II transport

Super Case II transport

Thin films Cylinders Spheres Thin films Cylinders Spheres Thin films Cylinders Spheres Thin films Cylinders Spheres

0.50 0.45 0.43 0.50 , n , 1.0 0.45 , n , 0.89 0.43 , n , 0.85 1.0 0.89 0.85 n.1 n . 0.89 n . 0.85

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

Super Case II model is an extreme transport mechanism where the polymer chains break. The outer gel layer meets the glassy nucleus, producing tension of compression on the nucleus up to a breaking point where the nucleus breaks. The main difference between the three mechanisms of non-Fickian diffusion (Case II, Anomalous Case, and Super Case II transport) is the rate of solvent diffusion into the matrix system. In Case II, the rate solvent diffusion is less than the polymeric relaxation process, in Anomalous transport, they have similar extent while, in Super Case II, the rate of solvent diffusion is much higher, causing an increase in solvent penetration (Klech and Simonelli, 1989). Usually in a drug-delivery system, more than one transport mechanism may occur at the same time, especially when swelling polymers are used. In these cases, Fickian diffusion may take place first, following a Case II transport mechanism and last, Super Case II might occur.

8.2.6 Weibull model This model is based on a function proposed originally by Weilbull (1951) and adapted later by Langenbucher (1972) to describe the drug release process. The equation that can be easily applied to almost all dissolution profiles is expressed in terms of the drug fraction accumulated (m) in the solution on time t:   2 ðt2Ti Þb m 5 1 2 exp a or log ½ 2lnð1 2 mÞ 5 b log ðt 2 Ti Þ 2 log a where a is the scale parameter that defines the timescale of the process, Ti is the localization parameter that represents the latency time of the release process (many times is zero), b is the form parameter that characterizes the type of curve and can be either b 5 1 (exponential), b . 1 (sigmoid, with ascendant curvature delimited by an inflection point), or b , 1 (parabolic, displaying high initial slope, and a consistent exponential character). The Weibull equation is an empirical model, and therefore can only describe and not characterize the dissolution kinetics of the drug. There is no parameter regarding the intrinsic dissolution factor of a drug and the equation does not allow any in vitro/ in vivo correlation (Costa and Lobo, 2001).

8.2.7 Other release parameters Other parameters used to characterize the drug release are the time taken for a given proportion of the active drug to be released into solution (i.e., t20%, t50%, and t90%),

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the fraction of drug in solution after a given time (sampling time), the dissolution efficiency (DE), and the mean dissolution time (MDT). DE is defined as the area under the dissolution curve up to a certain time, t, expressed as a percentage of the area of the rectangle described by 100% dissolution, in the same time according to the following equation: Ðt DE 5

ydt

0

y100 t

100%

where DE sums up the drug release profile into a single number and therefore allows comparison between a large number of dissolution profiles (Khan, 1975). MDT is defined as the mean time needed for the active ingredient to be dissolved and is described as the following equation: MDT 5

ABC WN

where ABC is the area between the drug dissolution curve and its asymptote and WN is the maximal amount of the drug substance that is dissolved (Rinaki et al., 2003).

8.3 Release profiles comparison Statistical methods are often used to study the drug release kinetics from modified release formulations. The comparison of the dissolution profiles in both graphical and numerical form is a useful but complex form and this is why researchers frequent use the aforementioned empirical equations. Moore and Flanner (1996) proposed an independent mathematical model for comparing the solubility profiles using the difference factor ( f1) and the similarity factor ( f2). 9 8 n P > > > jRt 2 Tt j > < = t51 n  f1 5 3 100 P > > > > : ; Rt (" f2 5 50 3 log

t51

#20:5 )  X 1 n 2 11 ðRt 2Tt Þ 3 100 n n21

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

where n is the number of observations, Rt is average percentage drug dissolved from reference formulation, and Tt is average percentage drug dissolved from test formulation time t. Factor f1 is proportional to the mean difference between the two profiles, whereas factor f2 is inversely proportional to the mean of squares of the difference between the two profiles with emphasis on the points with the largest deviation. When the two profiles are the same f2 5 100. When all time points have a mean difference of 10% f2 5 50. The FDA has established a range of f2 with 50 , f2 , 100 which indicates similarity between the two dissolution profiles.

8.4 Biopolymers in modified peroral drug delivery The development of pharmaceutical formulations involves the use of various excipients in addition to the active ingredient. Over the past decades, excipient development has become a very interesting research area in drug delivery as it plays an important role in the drug release process especially in modified release formulations. Moreover, the use matrix tablet systems as modified release formulations has advanced the industrial production of pharmaceutical dosage forms as there is no need for complex processes like coating and pelletization. In these cases, the drug release rate from the dosage form is controlled mainly by the type and proportion of the excipients used in these formulations. Among various excipients, biopolymers are an attractive choice due to their low toxicity and immunogenicity, stability, biocompatibility, and biodegradability (Mano et al., 2007). Natural polymers are polysaccharides, proteins, and polyesters from plant and animal species. Some of the most common biopolymers that are utilized as excipients in pharmaceutics are polysaccharides derived from plants (i.e., starch, cellulose, pectin, and some gums like guar and arabic gum), algae (i.e., alginates, galactans, and carrageenans), animals (i.e., chitin, chitosan, glycosaminoglycans, and hyaluronic acid), and microorganisms (i.e., dextran, bacterial cellulose, and some gums like gellan and xanthan gum). Polysaccharides consist of multiple monosaccharide units linked together by O-glycosidic bonds. The differences in the monosaccharide composition, the linkage types/patterns, the chain shape/length, and the molecular weight state the polysaccharides’ physical properties, such as solubility, viscosity, gelling potential, and/ or surface and interfacial properties (d’Ayala et al., 2008; Klouda and Mikos, 2008). Nanotechnology ensures the fabrication of these biopolymers, such as silk fibroins, collagen, gelatin, albumin, starch, cellulose, and chitosan to nanometer scale for various medical applications including delivery vehicles for both macro- and minidrug molecules (Jacob et al., 2018).

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Cellulose is a highly crystalline linear homopolymer found in most plants. Despite being a homopolymer of glucose, cellulose is not water soluble because of extensive intra- and intermolecular hydrogen bonding. However, the chemical modification or the use of suitable solvents can render cellulose soluble (Singh et al., 2018). Large-scale commercial cellulose ethers include carboxymethyl cellulose, methyl cellulose, hydroxyethyl cellulose, hydroxypropylmethylcellulose, hydroxypropylcellulose, ethyl hydroxyethyl cellulose, and methylhydroxyethylcellulose. Hemicelluloses are found in most plants and being mostly heteropolysaccharides; they are classified according to the sugar residues present, namely, xylans, mannans, arabinans, and galactans. They can be either linear or branched polymers (Singh et al., 2018). Starch is a heterogeneous polymer of α-D-glucose units derived from plants. The anhydrous glucose units are mainly linked by α-(1,4)-bonds and to some extent by α-(1,6)-linkages. The biopolymer consists of not only two distinguished structural forms, amylose and amylopectin, but also contains very small amounts of proteins, lipids and inorganic compounds (Singh et al., 2018). In pharmaceutical technology, starch can be used as a tablet diluent, disintegrant, or binder, and it can be compressed directly (Rowe et al., 2006). Studies have shown that it can be used in slow release and bioadhesive tablets (Hartesi et al., 2016). Pectins are complex polysaccharides present in plants. Researchers have studied pectins as hydrophilic excipient in controlled release tablets or colon-specific drugdelivery systems due to their gelling properties when used as a film coating or as ingredient in the matrix tablets (Singh et al., 2018; Vlachou et al., 2017). Gums and mucilages are also derived from plants and are used in pharmaceutics as sustained release excipients, binders, disintegrants, and so forth. Alginates are linear polysaccharides found in brown seaweed and marine algae. They have high molecular weights of 20
600 kDa and are used in pharmaceutics as stabilizers in emulsions, suspending agents, binders, and disintegrants in tablet manufacturing (Singh et al., 2018). Chitin is a naturally occurring polysaccharide, found widely in nature in fungi and yeast. Most forms of chitin, chitosan, and their derivatives are used in pharmaceutics as excipients in tablet manufacturing (diluent, binder, lubricant, disintegrating agent). Moreover, chitosan can be an ideal excipient for local delivery of drugs in the oral cavity due to its mucoadhesive properties (Singh et al., 2018). Dextran is synthesized by a large number of bacteria and is composed of the simple sugar glucose monomers. Dextran polymers have a number of medical/pharmaceutical applications particularly these derived from Leuconostoc mesenteroids. These are characterized by their content of 95% α-1,6-glucopyranosidic linkages and 5% 1,3-linkages (Dhaneshwar et al., 2006). Dextrans have been used for wound coverings, in surgical sutures, as blood volume expanders, to improve blood flow in capillaries in the treatment of vascular occlusion, and in the treatment of iron deficiency anemia in both

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

humans and animals (Singh et al., 2018). In pharmaceutical technology, dextran has been studied for its modified release characteristics due to the thick gelled layer that is produced when in contact with the biological media (Vlachou et al., 2017; Casettari et al., 2015; Efentakis and Siamidi, 2014; Vlachou et al., 2016).

8.5 Nanofibers in modified peroral drug delivery During the past years, drug-loaded nanofiber matrices have gained significant attention due to their unique characteristics that allow their use in a wide range of applications (Kikionis et al., 2015, 2017a,b; Akhgari et al., 2017; Kumbar et al., 2006; Al-Enizi et al., 2018; Cheng et al., 2018; Kenry and Lim, 2017; Chang et al., 2016; Gallego-Perez et al., 2016). Nanofibers are solid nanomaterials, with a diameter ranging between 1 and 1000 nm, that have high porosity and permeability, high surfacearea-to-volume ratio, mechanical strength, and flexibility (Garg et al., 2015; Sharifi et al., 2016). Nanofibers can be produced by an inexpensive technique called electrospinning from charged polymer solutions under a high electric field (Hamori et al., 2014; Frenot and Chronakis, 2003; Thakkar and Misra, 2017; Bhardwaj and Kundu, 2010; Greiner and Wendorff, 2007; Teo and Ramakrishna, 2006). The use of various biocompatible and biodegradable natural or synthetic polymers in the polymer solution and the electrospinning parameters tuning can tailor the nanofibers’ mechanical and biological properties (Ignatious et al., 2010; Toskas et al., 2011, 2012). Nanofibers may find applications in the biomedical sector as drug-delivery systems, in wound healing and in tissue engineering (Agarwal et al., 2008; Chakraborty et al., 2009; Ramakrishna et al., 2006). In oral drug delivery, electrospun nanofibers have been successfully used for the modified release of various incorporated drugs, achieving desirable release profiles (Akhgari et al., 2017). Researchers have prepared and characterized melatonin-loaded nanofibrous systems based on cellulose acetate, polyvinylpyrrolidone, and hydroxypropylmethylcellulose. The electrospun nanofiber mats that were inserted in hard gelatin capsules exhibited variable release profiles in the gastric-like fluids, ranging from 30 to 120 minutes. The electrospun nanofiber mats that were inserted in DRcaps capsules released melatonin at a slower pace (6 hours) (Vlachou et al., 2019a). In another study, nanofibers of cellulose acetate and polyvinylpyrrolidone loaded with melatonin were prepared and compressed at various pressures into monolayered tablets. In addition, because multilayered tablets consisting of a core containing the active ingredient and 1 or 2 barrier layers applied on the core during tableting have brought a new set of challenges in the design of oral-modified drug-delivery systems

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(Gavate et al., 2013; Efentakis et al., 2010; Efentakis and Peponaki, 2008), the nanofiber mats were incorporated in three-layered tablets, containing in the upper and lower layers, the combinations of lactose monohydrate and hydroxypropylmethylcellulose as modifying accessories. The in vitro dissolution profiles have showed promising results in modified peroral drug delivery (Vlachou et al., 2019b). Researchers have also prepared nanofibers loaded with mebeverine hydrochloride and the in vitro dissolution tests of the Eudragit fibers show pH-dependent drug release profiles. The results indicated very limited release at pH 2.0, but sustained release over approximately 8 hours at pH 6.8, and therefore Eudragit nanofibers have the potential to be developed as oral drug-delivery systems targeted the drug release in the intestinal tract (Illangakoon et al., 2014). In another study, researchers synthesized gelatin nanofibers by electrospinning, and tested them as a potential carrier for oral drug delivery for a model hydrophobic drug, piperine. The results illustrated good compatibility of the hydrophobic drug in gelatin nanofibers with promising controlled drug release patterns by varying crosslinking time and pH of release medium (Laha et al., 2016). Nanofiber-based capsules including as model drugs uranine (as a water-soluble drug) and nifedipine (as a water-insoluble drug) were prepared for controlled release delivery systems using Eudragit S100 as a polymer, and evaluated their in vitro drug dissolution profiles and in vivo pharmacokinetics in rats. The in vitro release of both drugs from the nanofiber-packed capsules and milled-powder of nanofiber-packed capsules showed controlled release as compared with capsules of a physical mixture of the polymer and each drug. The in vivo pharmacokinetic study in rats demonstrated that application of nanofibrotic technique as a drug-delivery system offers drastic changes in pharmacokinetic profiles for both water-soluble and water-insoluble drugs (Hamori et al., 2014). Nanofiber-based tablets with acetaminophen were developed for controlled release delivery systems and evaluated in vitro and in vivo. The in vitro dissolution rate profiles showed controlled release aspects, whereas the in vivo pharmacokinetic studies in rats after intraduodenal administration showed sustained release properties. The nanofibers showed promising results as oral delivery systems for sustained release regulation as these new tablet formulations did not disintegrate in the intestine in the lower pH region, and could regulate the release of the model drug in a pH-dependent manner (Hamori et al., 2016).

8.6 Examples of liposomal-modified release formulations in clinical use Liposomal drug-delivery systems have been used widely in the recent years for various diseases ranging from cancer treatment to pain management. The main

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

advantage of liposomes when used in pharmaceutical products is the control over the pharmacokinetic and pharmacodynamic properties of the drug that leads to the improvement of the product’s therapeutic effect (Bulbake et al., 2017). The following are examples of liposomal-modified release formulations in clinical use that are currently in the market. Doxil (an i.v. injection of DOX hydrochloride) developed by Sequus Pharmaceuticals, is the first FDA-approved nanodrug-delivery system based on PEGylated liposome technology. It is used for the management of ovarian cancer, multiple myeloma, and HIVassociated Kaposi’s sarcoma. In this formulation, the drug is within the liposome core, allowing higher retention with less drug efflux in circulation, whereas at the same time, providing acceptable drug rates in tissues (Gabizon, 1995; Gabizon et al., 1994). Marqibo (a vincristine sulfate liposomal injection) has been developed by Talon Therapeutics, Inc. for the treatment of adult patients with Philadelphia chromosomenegative acute lymphoblastic leukemia with second or greater relapse or whose disease has advanced after two or more antileukemia therapies. In this formulation, the drug is encapsulated in the aqueous interior core of a sphingomyelin/cholesterol liposome (optisome). These optisomes were specifically developed not only to promote the higher loading and holding of the drug but also to increase its circulation time and its slow release into the tumor vasculature (Webb et al., 1995; Johnston et al., 2006; Krishna et al., 2001). DepoDur is an epidural morphine sulfate sustained release liposome injection developed by SkyePharma. In this formulation, the drug is encapsulated in a lipid foam by DepoFoam Technology, permitting its sustained release into the epidural space for a longer time. The multivesicular structure of the foam encapsulates the active drug and then ruptures in a time-dependent way to release it for a 48-hour period (Alam and Hartrick, 2005; Hartrick and Manvelian, 2004). Exparel (a bupivacaine liposome injection) has been developed by Pacira Pharmaceuticals, Inc. for the management of postsurgical analgesia. This formulation is based on Depofoam Technology, containing a novel phospholipid excipient, dierucoylphosphatidylcholin and is intended for extended release of the drug up to 72 hours (Angst and Drover, 2006; Richard et al., 2011).

8.7 Conclusion The development of tools that facilitate product development by reducing the biostudies needed is very important especially in modified release systems. The use of mathematical equations as well as other parameters used to characterize the drug release (the time taken for a given proportion of the active drug to be released

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into solution, the fraction of drug in solution after a specified time, the dissolution efficiency and the mean dissolution time) are essential keys when designing a pharmaceutical formulation or evaluating the drug release process either in vitro or in vivo. The comparison of the dissolution profiles can be achieved by statistical methods in either graphical and numerical form or the use of the difference ( f1) and similarity factor ( f2). The careful selection of excipients, with the appropriate physicochemical properties and quantity, plays an important role in the drug release process especially in modified release formulations. Among various excipients, biopolymers and nanomaterials are attractive choices due to their unique characteristics. Electrospun nanofibers are recently studied for the modified release of various incorporated drugs with promising results. Liposomal drug delivery systems have been used widely for control over the pharmacokinetic and pharmacodynamic properties of the drug that leads to the improvement of the product’s therapeutic effect.

Self-assessment questions 1. Which one of the following is NOT true? a. Modified release formulations are most useful for drugs with a long half-life b. Modified release formulations can often reduce side effects. c. Modified release formulations can improve patient compliance d. Modified release formulation can be used for local drug delivery 2. Which one of the following is NOT true? a. Drug release from reservoir systems is controlled by diffusion b. Drug release from matrix systems is controlled by diffusion c. Drug release from reservoir systems normally follow zero-order kinetics d. Drug release from matrix systems normally follow zero-order kinetics 3. Which one of the following is TRUE? Biopolymers that are plant and animal species can be: a. Polysaccharides b. Proteins c. Polyesters d. All of the above 4. Which one of the following is TRUE? The benefit(s) of electrospinning when compared with other nanofabrication techniques is: a. High surface-to-volume ratio b. Inexpensive

Biopolymers, liposomes, and nanofibers as modified peroral drug release formulants

c. Easy d. All of the above 5. Which one of the following is TRUE? Liposomal-modified release formulations is in clinical use for various diseases like: a. Pain management b. Cancer management c. Postsurgical analgesia d. All of the above

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5621. Akhgari, A., Shakib, Z., Sanati, S., 2017. A review on electrospun nanofibers for oral drug delivery. Nanomed. J. 4, 197
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Efentakis, M., Peponaki, C., 2008. Formulation study and evaluation of matrix and three-layer tablet sustained drug delivery systems based on carbopols with isosorbite mononitrate. AAPS PharmSciTech. 9, 917
923. Efentakis, M., Siamidi, A., 2014. Design and evaluation of a multi layer tablet system based on Dextran. J. Pharm. Drug. Deliv. Res. 3, 2. Efentakis, M., Naseef, H., Vlachou, M., 2010. Two- and three-layer tablet drug delivery systems for oral sustained release of soluble and poorly soluble drugs. Drug. Dev. Ind. Pharm. 36, 903
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Kikionis, S., Ioannou, E., Toskas, G., Roussis, V., 2015. Electrospun biocomposite nanofibers of ulvan/ PCL and ulvan/PEO. J. Appl. Polym. Sci. 132, 42153. Kikionis, S., Ioannou, E., Konstantopoulou, M., Roussis, V., 2017a. Electrospun micro/nanofibers as controlled release systems for pheromones of Bactrocera oleae and Prays oleae. J. Chem. Ecol. 43, 254
262. Kikionis, S., Ioannou, E., Andrén, O.C.J., Chronakis, I., Fahmi, A., Malkoch, M., et al., 2017b. Nanofibrous nonwovens based on dendritic-linear-dendritic poly (ethylene glycol) hybrids. J. Appl. Polym. Sci. 135, 45949. Klech, C.M., Simonelli, A.P., 1989. Examination of the moving boundaries associated with non-Fickian water swelling of glassy gelatin beads: effect of solution pH. J. Membr. Sci. 43, 87
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297.

CHAPTER NINE

Grafted polymethacrylate nanocarriers in drug delivery Dorota Neugebauer Silesian University of Technology, Faculty of Chemistry, Department of Physical Chemistry and Technology of Polymers, Gliwice, Poland

9.1 Graft poly(meth)acrylates, including molecular brushes The macromolecules, which contain side chains grafted to backbone, are known as the one-dimensional (1D) structures representing extraordinary properties including selective solubility, wormlike conformation, compact molecular dimension, and chain end functionalities. Because of that, they are of great interest in the field of polymer and material science finding broad applications in catalysis, nanolithography, optoelectronics, biomineralization, medical diagnosis, or drug delivery. In comparison to the corresponding linear analogs with similar molecular composition, the graft copolymers are characterized with better solubility, smaller hydrodynamic radius, lower melt and solution viscosity, lower glass transition temperature, and lower crystallinity in case of a crystalline polymers. The unusual architectures of graft copolymers are classified onto loosely and densely grafted (brush) structures, homo- and heterografted (two or more types of polymeric side chains) structures, where the core
shell (block structure of side chains), Janus-type (phase separation of statistically distributed side chains), blocktype (block distribution of side chains among the backbone) cylinders can be distinguished (Fig. 9.1). The well-defined graft polymers are efficiently synthesized by using the living/controlled methods, such as anionic and cationic polymerizations, ringopening polymerization (ROP) including ring-opening metathesis polymerization and reversible-deactivation radical polymerizations [atom transfer radical polymerization (ATRP), ATRP by initiators for continuous activator regeneration, ATRP via activators generated by electron transfer, addition-fragmentation chain transfer (RAFT)]. The building blocks with different chemical nature can be obtained by the combination of various polymerization methods to design the graft copolymers, which are able to construct hierarchical nanostructures with required properties for the specific applications. There are three main techniques: (1) “grafting onto” applying multi- and monotelechelic polymers, (2) “grafting from” with the use of multifunctional macroinitiator, and (3) “grafting through” in via polymerization of macromonomer. Nanomaterials for Clinical Applications. DOI: https://doi.org/10.1016/B978-0-12-816705-2.00009-6

© 2020 Elsevier Inc. All rights reserved.

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Figure 9.1 Graft copolymer structures.

The development of synthetic strategies is advantageous for the macromolecular engineering of precisely controlled composition, structure, and functionality (Bhattacharya and Misra, 2004; Hadjichristidis et al., 2006; Sheiko et al., 2008; Ito et al., 2014; Neugebauer, 2015; Verduzco et al., 2015; Neugebauer, 2016). Chemical heterogeneity of backbone and side chains provides predispositions for stimuli-sensitive behavior by changing configuration/conformation via a supramolecular intra- and interactions. The hydrophobic/hydrophilic ratio is responsible for the formation of micelles/nanoparticles with various morphologies and intermicellar aggregation in an aqueous solution with possible reversible effects in another selective solvent (Abetz and Simon, 2005; Rodriguez-Hernandez et al., 2005; Riess, 2003). The self-assembling behavior is convenient for encapsulation of drug and then the controlled release due to the macromolecular carrier prolongs blood circulation time and reduces nonspecific adsorption of proteins (Kadajji and Betageri, 2011; Felice et al., 2014). The drug release can be modulated by additional external stimuli, such as temperature, pH, and reducing agent. It is well known that the tumor tissues show acidic pH, high salt, and glutathione concentrations, which are competent to stimulate the activity of the smart polymer carriers with the release of the anticancer drugs. The polymers are also applied for the other types of therapies, such as antiinflammatory or antibacterial, to support controlled pharmacokinetics and pharmacodynamics. Because of drug delivery applications, the micelles should exhibit excellent biocompatibility and cellular uptake property. The efficient gene therapy for numerous hard curable diseases, such as cancer, genetic and infectious diseases, is supported by the nonviral polycations as the attractive carriers due to their good biocompatibility, high condensation ability, as well as easily adjustable and controlled structures (Xu and Yang, 2011). Cationic polymers with negatively charged nucleic acids form positively charged polyplexes, which can

Grafted polymethacrylate nanocarriers in drug delivery

be translocated across negatively charged cell membranes and then dissociated with release nucleic acid into the nucleus for gene expression. The self-assembling polymers are also functionalized with antibodies, proteins, peptides (especially with sequence of arginyl
glycyl
aspartic acid), carbohydrates (mostly glucose, mannose, galactose, and lactose), transferrin (polypeptide glycoprotein), aptamers, vitamins, folates, and nucleic acids to design the targeting drug delivery systems (DDSs) (Bertrand et al., 2014; Wang et al., 2016). The attached ligands are recognized by receptors of the injured cells, where drug-carrying polymer systems is transported to start therapeutic process. In this chapter, the amphiphilic graft copolymers, which contain poly(meth)acrylate backbones and polyether, polyester, and/or poly(meth)arylate side chains, including molecular brushes, linear-graft block, and disulfide linked structures, are discussed in the aspect of the self-assemblies, which are able to encapsulate small biologically active molecules.

9.2 Carriers based on poly(ethylene glycol) poly(meth)acrylate brushes The unique properties of poly(ethylene glycol) (PEG), such as water solubility, biocompatibility, nontoxicity, make this polymer as the fascinating building block in the drug carriers. The studies on micellization, drug loading, and release profiles confirmed that the grafted topology of nanocarriers is advantageous in drug delivery due to the extra parameters, such as length of PEG side chains and grafting degree, which are responsible for specific adjustment of hydrophilic/hydrophobic balance in the selfassembled nanoparticles (Neugebauer, 2007a). Amphiphilicity of brush polymers based on PEG macromonomers is caused by the combination of hydrophobic polymethacrylate backbone and hydrophilic PEG side chains, whereas the interactions between densely grafted chains (one side chain per each repeating unit of backbone, Fig. 9.1A) result in soft elastomeric properties. The methoxy-functionalized PEG methacrylates (mPEGMA, Fig. 9.2A) are the most

Figure 9.2 Terminal end groups of PEG side chains in graft copolymers based on (A) mPEGMA (poly(ethylene glycol) methyl ether methacrylate/methoxy-functionalized poly(ethylene glycol) monomethacrylate)), (B) poly(ethylene glycol) bismethacrylate, (C) PEGMA (hydroxy-functionalized poly(ethylene glycol) monomethacrylate), (D) dimethylsilyl polystyrene resin-functionalized poly (ethylene glycol) monomethacrylate), (E) ePEGMA (poly(ethylene glycol) ethyl ether methacrylate), (F) phPEGA (poly(ethylene glycol) phenyl ether acrylate).

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studied to prepare the brush/graft polymers, for which amorphous morphology was observed in case of shorter PEG chains, whereas introduction of macromonomer with longer PEG chains has generated the crystalline properties. The spontaneous polymer networks were prepared by conventional free radical polymerizations, whereas the ATRP-supported high-macromonomer conversions and the well-defined soluble brush polymers with no gel fraction (Neugebauer et al., 2003b; Ying et al., 2004). The cross-linking effect can be promoted by the incorporation of bismethacrylate PEG macromonomer (Fig. 9.2B) (Dubrovskii et al., 2003; Kozlov et al., 2006). In the case of hydrogels, the relationship between the elastic modulus and the swelling behavior in association with the length of side chains was evident providing classical soft gels at long side chains and rigid gels at short side chains. Similarly to mPEGMA, the well-defined brushes of hydroxy-functionalized PEGMA (Fig. 9.2C) have been synthesized, although the risk of condensation of hydroxyl groups is highly probable. The homopolymer and copolymers with methyl methacrylate (MMA) containing more than 24 mol.% (57 wt.%) of PEGMA (n 5 5, where n is the number of repeating units in the PEG) were water soluble and exhibited sharp lower critical solution temperatures (LCSTs), which was increasing with PEGMA content (Ali and Stover, 2004). The chlorinated dimethylsilyl polystyrene resin can be used during the polymerization of PEGMA to exclude side reactions by isolation of hydroxyl end group (Fig. 9.2D), which were deprotected in the obtained polymer by the anion exchange with chloride acid (Taniguchi et al., 2006). The solubility and thermoresponsive properties corresponding to the phase transitions, due to a reversible coil-globule transformation, can be regulated by the length of mPEG chains increasing hydrophilicity of PEG brushes. The extension of grafts from 2 to 7
8 ethylenoxy units provided the shifting of LCST in a wide range of temperatures (26 C
90 C with respect to the polymer concentration and molecular weight) (Lutz, 2008). The LCST has been also precisely tuned by changing the ratios of two mPEGMA comacromonomers (two vs nine repeating ether units) to induce insolubility in 31 C
67 C (Lutz et al., 2007b). These phase transition results are in agreement with those to their analogous hydrogels, which were obtained by controlled radical polymerization with the addition of crosslinker (ethylene glycol dimethacrylate), exhibiting swelling/deswelling behavior in water at  40 C and  50 C, for copolymers containing 10 and 20 mol.% of mPEGMA (n 5 9) per chain, respectively (Lutz et al., 2007a). Variations in PEG content and PEG chain length have caused the temperature changes in the phase transition effect, but the LCST can be also shifted through the addition of salt as it was reported for copolymers of mPEGMA and ethoxy-functionalized PEG methacrylate (ePEGMA, Fig. 9.2E) with nine versus three repeating ether units, respectively (Magnusson et al., 2008). The statistical copolymers, in which the part was replaced with a more hydrophilic mPEG graft block demonstrated the clouding points at slightly higher temperature than LCST of the fully statistical analogs. However, this temperature was significantly

Grafted polymethacrylate nanocarriers in drug delivery

reduced in the presence of a kosmotropic salt yielding stable micelle-like assemblies with burst release of an encapsulated carboxyfluorescein. The crystalline-amorphous morphology has been displayed by dual PEG brushes based on mPEGMA and phenoxy-functionalized PEG acrylate (phPEGA, Fig. 9.2F) varying the length of side chains (23 vs 4 repeating ether units, respectively) (Neugebauer et al., 2006a). In addition, the variety in relative reactivity of methacrylate versus acrylate macromonomers provided the spontaneous gradient structure with the instantaneous composition, which is continuously altering along the main chain (predominated methacrylate gradually changing through the heterosequences to acrylate domination). The differences in comonomer sequences, that is, block versus statistical versus gradient, have influenced on thermal and mechanical properties due to various mobilities of side chains with consequences for the self-assembling behavior in aqueous solution, but also degradation rates and toxicity of copolymers (Jakubowski et al., 2008).

9.3 Carriers based on poly(ethylene glycol) grafted poly(meth) acrylates The hydrophilic
hydrophobic properties of PEG brush copolymers, which are crucial for the drug delivery, are constantly modified by adding of the extra units or segments. These complex structures, mostly loosely PEG-grafted copolymers, semigrafted linear-brush block copolymers, and densely heterografted copolymers are self-organized in aqueous media into the micellar nanoparticles with ability to carry hydrophobic drug in the core covered by the hydrophilic PEG shell (Fig. 9.3A).

Figure 9.3 Typical micellar superstructures from graft copolymers with loaded drug.

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The randomly distributed hydrophilic PEG grafts among the hydrophobic polymethacrylate backbone (side chain per more than one repeating unit of backbone related to grafting degree lower than 100%, Fig. 9.1B) have been resulted by copolymerization of PEG methacrylate macromonomer with the low-molecular weight comonomers, such as MMA (Table 9.1 1) (Maksym-Be˛benek et al., 2014; MaksymBe˛benek and Neugebauer, 2015), lilial-functionalized methacrylate (LILMA, Table 9.1 2) (Morinaga et al., 2009), and prodrug monomers: (Son et al., 2016) methacrylate functionalized with camptothecin (Table 9.1 3a) or kinase inhibitor dasatinib (Table 9.1 3b). In case of the latter ones, that is, the PEG-grafted polymer
drug conjugates (prodrug polymer carrier systems with the chemotherapeutic agent), drug release by statistical copolymers was faster than for diblock analogs yielding micelles with the drug block separated from the aqueous phase. The introduction of hydrophilic comonomer, for example, N,N-dimethylacrylamide (DMAAm, Fig. 9.4A), which is able to be quaternized to ammonium units, increased solubility of the PEG graft copolymers (Neugebauer, 2008), but in some cases, it may limit the self-assembling due to too high hydrophilicity. Because of that, the positively charged water-soluble random copolymers of mPEGMA and (3-(methacryloylamino)propyl) trimethylammonium chloride (Fig. 9.4C) were completed with oppositely charged poorly water-soluble biosurfactants, such as of cholate, deoxycholate, oleate, and laurate sodium salts (Nisha et al., 2004; Kizhakkedathu et al., 2005). The water solubility of the polyion complexes dependent on anion type and PEG content gives opportunity for solubilization of hydrophobic drug in the hydrophobic core, but the presence of ionic units leads to the core
shell
corona micelles with the middle layer of the ionized fraction (Fig. 9.3B). The block copolymers of PEG polymethacrylate brush have been prepared by the sequential polymerization to include the linear segment (Fig. 9.1C), for example, based on the UV-sensitive 4,5-dimethoxy-2-nitrobenzyl methacrylate (DNBMA, Table 9.1 4) (Xiang et al., 2018) or smectic mesogen cholesteryl-functionalized methacrylate (ChEMA, Table 9.1 5a) (Li et al., 2018). The hydrophilicity of diblock copolymer can be increased by the stimulus-triggered cleavage of sensitive groups, what not only caused the formation of extra pendant hydrophilic groups at the backbone (carboxylic acid, hydroxyl), but also the release of hydrophobic drug from disassembled PEG-based micellar nanocarrier. In the case of DNBMA o-nitrobenzyl groups were cleaved by the photochemical reaction upon the near-infrared light, whereas an acid-labile acetal moieties in LILMA and ChEMA units were stable in the aggregated micelles, but the temperature responsive copolymer can be hydrolyzed providing micelle dissociation with controlled release of the lilial or cholesteryl as the model hydrophobic biocompounds. The cholesterol molecules have been also conjugated with the use of “postpolymerization” modification strategy, which is convenient to introduce the acetal linker by alcoholysis via hydroxyl groups of 2-hydroxyethyl methylacrylate (HEMA, Fig. 9.4H)

Table 9.1 Drug carriers based on poly(ethylene glycol) graft polymethacrylates. R No. O O O R Drug Characteristics , where R2 is O R1 1

2

n

1

H n 5 9, 6

2

CH3 n 5 8.5

3a

PEGMA n 5 19

3b

4 5a


CH3

O

O

Lilial

O

O

O O

O

HN

O

O

O

O

5b CH3 n58

O

N

N

N

O

O N

O

N

O

O

CH3 n54

CPT

N

O

O

O

O

CH3

IMC

N

N

H N

S O

KID

Cl

DOX

O O

O O

H

Cholesteryl

H H

O

O O

H H H

DOX

Water-soluble polymers (min. 17 mol.% of PEGMA), temperature-sensitive, LCST 5 39 C
70 C, DLC 5 22%
88%, faster drug release at pH 7.4 than at pH 5.0 Acetal linking lilial-conjugate, aggregationdissociation dependent on temperature (40 C
50 C) in NaCl presence Drug
polymer conjugates, live animal imaging in human prostate cancer cell line (PC-3), fluorescently labeled copolymers trafficked to the tumor 24 hours post injection, ex vivo analysis polymer accumulated in the tumor with kidney excretion, in vitro cytotoxicity on K562-S and K562-R cells Near-infrared light-responsive nanocarrier, photocontrolled drug release Liquid-crystalline block copolymer, conjugate, tunable morphologies (short cylindrical micelles, nanofibers, fringed platelets, and ellipsoidal vesicles) Cytotoxicity for HepG2 cells, pH-sensitive micelles, anticancer drug carriers, in vitro release rate higher in pH 6.0 than in pH 7.4, DLC 5 15%, DEE 5 50%

Ref.

Maksym-Be˛benek et al. (2014), MaksymBe˛benek and Neugebauer (2015) Morinaga et al. (2009)

Son et al. (2016)

Xiang et al. (2018) Li et al. (2018)

Zhang et al. (2014)

(continued)

Table 9.1 (Continued) R No. O O O R , where R2 is O R1 1

5c

O

7

CH3 n55

O

8

CH3 n 5 45, 114

O

H

O

H H

O

N

N

O

, P

H+ N

O

Characteristics

Ref.

H

5

CH3 n 5 45 CH3 H

6

Drug

2

n

p

O

P

O N H

O- Na+ O- Na

+

O-Na+ O- Na+

H

DOX

Effectively internalized by HeLa cells, physical drug entrapment, DLC 5 21%, DLE 5 85% Gal Efficient transfection of polyplex into HepG2 cells, cell imaging system to specific ligand
receptor interactions Gal/ASGPRs on the HepG2 cell surface DOX Inner-layer crosslinked micelles, high inhibition for HepG2 cell proliferation, DLC 5 15%, DLE 5 60%
85% DOX, Manganese nanocarriers, DLE 5 70%
100%, cisplatin, fast release in acidic pH, anticancer activity carboplatin against MCF-7 cells, nontoxic against mouse hepatocytes, contrast agents, simultaneous MRI diagnosis, and cancer treatment

Nguyen et al. (2014) Hao et al. (2014b)

Zhang et al. (2018)

Pothayee et al. (2014)

ASGPRs, Asialoglycoprotein receptors; CPT, camptothecin; DLC, drug loading content; DLE, drug loading efficiency; DOX, doxorubicin; IMC, indomethacin; KID, kinase inhibitor dasatinib; LCST, lower critical solution temperature; MRI, magnetic resonance imaging; PEG, poly(ethylene glycol); PEGMA, PEG methacrylates.

Grafted polymethacrylate nanocarriers in drug delivery

Figure 9.4 Stimuli-sensitive comonomers, (A) DMAAm (N,N-dimethylacrylamide), (B) NIPAAm (N-isopropylacrylamide), (C) (3-(methacryloylamino)propyl) trimethylammonium chloride, (D) DMAEMA/DMAEA (2-(N,N-dimethylamino)ethyl (meth)acrylate), (E) DEAEMA (2-(N,N-diethylamino)ethyl methacrylate), (F) EOxMA (2-ethyl-2-oxazoline methacrylate), (G) MAA (methacrylic acid), and (H) HEMA (2-hydroxyethyl methylacrylate).

units statistically distributed with pH-sensitive 2-(N,N-diethylamino)ethyl methacrylate (DEAEMA, Fig. 9.4E) in the diblock terpolymer P(HEMA-co-DEAEMA)-bPmPEGMA (Table 9.1 5b) (Zhang et al., 2014). Similarly the pendant hydroxyl groups of PEGMA units were modified with galactosamine (Gal), which is easily recognized by asialoglycoprotein receptors (ASGPRs) over-expressing hepatoma cells (Hao et al., 2014b). This graft copolymer was combined with segment of 2-(N,N-dimethylamino)ethyl methacrylate (DMAEMA, Fig. 9.4D), which at acidic and neutral conditions is protonated gaining positive charges as the advantageous to bind DNA via electrostatic interaction with the formation of polycation/DNA polyplex (Fig. 9.3E). The ligand-conjugated PEG brush-cation copolymer P (mPEGMA-co-PEGMA)-b-PDMAEMA (Table 9.1 6), which was internalized via a receptor-mediated endocytosis process, has been claimed as a promising platform for hepatoma-targeting delivery of genes. The hydrophilic ABC-type triblock copolymers composed of a charged middle block (PDMAAm quaternized with CH3I), and thermosensitive outer blocks with different LCSTs, where segment A corresponds to ePEGMA polymer, whereas segment C is a random copolymer of mPEGMA and fluorescent rhodamine B-functionalized methacrylate (LCSTA , LCSTC), have been recognized as doubly thermosensitive (Hu et al., 2016). The micelles with A block core were formed in water at temperature above the LCSTA but below the LCSTC, whereas above the LCSTC, the C-block corona was collapsed yielding a two-compartment 3D network micellar hydrogel. The gelation effect with addition of hairy silica nanoparticles, which acted as “seeds” adsorbing the collapsed C block, was enhanced by the formation of bridging chains between micellar cores and nanoparticles. The triple-stimuli-responsive

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behavior activated by temperature-sensitive N-isopropylacrylamide (Fig. 9.4B), pHsensitive DMAEMA and methacrylic acid (MAA, Fig. 9.4G) is represented by tetrablock copolymer PmPEGMA-b-PDMAEMA-b-PNIPAAm-b-PMAA (Table 9.1 7). In acidic aqueous solution, it was self-assembled into four-layer micelles (Fig. 9.3C), whereas the inner-layer crosslinked micelles, including core-crosslinked, shellcrosslinked, and shell
core dilayer-crosslinked micelles can be obtained via thiolresponsive disulfide bonding (Fig. 9.3F) (Zhang et al., 2018). The release rate of encapsulated drug from the crosslinked micelles was accelerated by combined stimulations [pH 5, 37 C, 10 mM dithiothreitol (DTT)]. The dual delivery systems encapsulating drug and metal nanoparticles are represented by the amphiphilic block copolymer of PEG brush with the liquid-crystalline cholesterol pendant polymethacrylate (Table 9.1 5c) end-capped by thiol group, which was able to anchor gold nanoparticles via the coordination bonding (Nguyen et al., 2014). The encapsulation of metal nanoparticles (Fig. 9.3D) has been also reported for the pHresponsive hydrogel of statistical terpolymer of mPEGMA (n 5 21), HEMA, and MAA, which was prepared by in situ reducing silver ions anchored by the deprotonated carboxyl acid groups (Xiang and Chen, 2007). The silver nanoparticles dispersed in hydrogel have supported much higher swelling ratio and faster deswelling rate as well as higher pH switchable electrical properties then the pure hydrogel. Similarly the ultrasmall superparamagnetic iron oxide nanoparticles was stabilized by block copolymer with a PMAA segment containing pendant carboxylates as multidentate anchoring groups and a hydrophilic PEG-grafted polymethacrylate segment (Chan et al., 2014). Their colloidal stability and effectiveness of T-1-weighted contrast for magnetic resonance imaging (MRI) can be enhanced by the ratios of both blocks. In another case, the manganese ions were complexed by units of ammonium bisphosphonate methacrylate with propyl or hexyl spacer (Table 9.1 8) in the statistical copolymer of mPEGMA (Fig. 9.5A)

Figure 9.5 Self-assembling ionic graft polymers carrying DOX; DOX, Doxorubicin.

Grafted polymethacrylate nanocarriers in drug delivery

(Pothayee et al., 2014). These macromolecular MRI-positive contrast agents for chemotherapeutic delivery have demonstrated a few times higher relaxivities at physiological pH in comparison to a commercial controls.

9.4 Poly(ethylene glycol) and biodegradable polyester nonlinear amphiphilics The delivery polymeric systems are developed by the combination of the biocompatible thermoresponsive PEG graft polymers with (bio)degradable hydrophobic polyesters, such as poly(ε-caprolactone) (PCL), poly(lactic acid) (PLA). Polyesters are also promising biomaterials in tissue engineering as surgical sutures, implants, bioresorbable stents, or a fibrous coatings inside the vascular prostheses. In all these biomedical applications, degradation time of polyesters should be arranged with respect to the efficiency time of the healing process (Washington et al., 2017). The interactions between the hydrophilic PEG and hydrophobic polyester chains are self-assembled into nanoparticles with various morphologies, but in the aqueous environment, like blood stream, the polyester segments are hydrolyzed supporting decomposition of micelles or disaggregation. After that the remaining short chains of polyesters are biodegraded by metabolic pathway(s) to the Krebs cycle products, which are easily excreted from organism. The core
shell micellar systems solubilizing model drugs in the degradable core (Fig. 9.3A) have been reported for the self-assemblies of amphiphilic linear-PEG brush block copolymers (Fig. 9.1C), including AB-type with linear PCL (Table 9.2 1) (Knop et al., 2013), or cholesteryl-functionalized PLA (Table 9.2 2) (Bagheri et al., 2013) and ABA triblock copolymers containing the outer PEG brush segments in combination with the inner linear polyester, that is, PCL (Table 9.2 3) (Luzon et al., 2010) or PLA (Table 9.2 4) (Bakkour et al., 2013; Hu et al., 2015). Usually the encapsulation of hydrophobic drug caused reduction of micelle size and LCST in comparison to the blank micelles. The enhanced interactions between hydrophobic polyester with drug and polymethacrylate chains slowed down the burst-like release. The drug and gene codelivery with the fluorescent detection has been postulated for the multiresponsive nanoparticles prepared by the amphiphilic diblock copolymer containing an acid-cleavable acetal linkage between fluorescent coumarinfunctionalized PCL and hydrophilic loosely grafted copolymer of thermosensitive mPEGMA with ionizable DMAEMA (Table 9.2 5) (Hao et al., 2014a). The micelles with simultaneously encapsulated anticancer drug and DNA have formed a micelleplex with hydrophilic PEG chains on the surface (Fig. 9.3E), whereas the cleavage of the acetal linkage under intracellular acidic conditions caused the drug release.

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Table 9.2 Drug carriers based on poly(ethylene glycol) and biodegradable polyester based nonlinear amphiphilics. No. Copolymer

nPEG

Drug

Characteristics

Ref.

1 2

PCL-b-PmPEGMA PLA-b-PmPEGMA

9 114

DOX Naproxen

3

PmPEGMA-b-PCL-b-PmPEGMA

8
9

Cell lines: L929, HEK293, MCF-7, MDA-MB231 Cholesteryl-functionalized PLA, DLE 5 54%
60% and DLC 5 11%
12% LCST at 75 C
85 C

4a

PmPEGMA-b-PLA-b-PmPEGMA

5

Knop et al. (2013) Bagheri et al. (2013) Luzon et al. (2010) Bakkour et al. (2013) Hu et al. (2015)

Clofazimine

2 vs Curcumin 4b P(mPEGMA1-co-mPEGMA2)-boligo PLA-b-P(mPEGMA1-comPEGMA2) 5 PCL-b-P(mPEGMA-co-DMAEMA) 5 DOX 6

PmPEGMA-b-PCLMA

7

PmPEGMA-b-P(PLAMA-co-MAA) 8
9

Paclitaxel, curcumin Ibuprofen

8

P(PLAMA-co-mPEGMA)

114

Naproxen

9

P(CLA-co-mPEGA)

23

Quercetin

45

CPT

114

DOX

10 P(PLAMA-co-PDEAEMA-comPEGMA) 11 P(HEMA-graft-PLA-b-PEG)

19

LCST 45 C (blank micelles) and 41 C
38 C at 6%
11% drug loading, temperature-dependent release (20% higher above the LCST than that at 37 C) Polymer
DNA polyplex, synergistic gene and drug therapy, acetal Hao et al. (2014a) linking coumarin, fluorescence detection Subsequent release, DLEB90% Colombo et al. (2016) Oral drug administration, pH-dependent release, in gastric fluid Yang et al. (pH 1.2) micelle shrinking and aggregation, drug protected (2012a,b) from the harsh environment in stomach, in intestinal fluid (pH 7.4) ionized carboxyl groups, micelle swelling, accelerated release Cholesteryl-functionalized PLA, phase of slow drug release, Bagheri and DLE 5 85%
89% and DLC 5 17%
18% Bigdeli (2013) Oral drug delivery system, DLE and DLCB5%, biphasic drug Nikfarjam et al. release profile with an initial burst effect (2014) Multifunctioning nanoparticles efficient carrier for cancer cytosolic Shen et al. (2008) drug delivery, DLE 5 60%, DLC 5 95% Multifunctional theranostic micelles, targeted by TRC105, PET Guo et al. (2014) imaging, studies on 4T1 murine breast tumor-bearing mice, HUVEC (CD105-positive), MCF-7 cells (CD105-negative)

CLA, Caprolactone acrylate; CPT, camptothecin; DLC, drug loading content; DLE, drug loading efficiency; DOX, doxorubicin; HEK293, human embryonic kidney 293 cells; HUVEC, human umbilical vein endothelial cells; HEMA, 2-hydroxyethyl methylacrylate; LCST, lower critical solution temperature; MAA, methacrylic acid; MCF-7, Michigan Cancer Foundation-7, breast cancer cells; mPEGMA, methoxy-functionalized PEG methacrylates; nPEG, number of repeating units in PEG; PCL, poly(ε-caprolactone); PDEAEMA, poly(2-(diethylamino)ethyl methacrylate); PEG, poly (ethylene glycol); PEGA, PEG acrylate; PET, positron emission tomography; PLA, poly(lactic acid); PLAMA, poly(lactic acid) methacrylate.

Grafted polymethacrylate nanocarriers in drug delivery

There is also variety of the heterografted brushes composed of both PEG and polyester side chains attached to the poly(meth)acrylate backbone. They are represented by the block sequence of poly(methacrylate macromonomer)s of PEG and PCL (Table 9.2 6, Fig. 9.1E) (Colombo et al., 2016), where the hydrophobic segment can be randomly grafted by introduction of pH-sensitive units, for example, P (PLAMA-co-MAA) (Table 9.2 7) (Yang et al., 2012a,b), to accelerate drug release. Another strategy corresponds to Janus-type cylindrical brushes obtained by random copolymerization of the methacrylate macromonomers of mPEG and cholesterylfunctionalized PLA (Table 9.2 8, Fig. 9.1D) (Bagheri and Bigdeli, 2013). The polymers synthesized by controlled ATRP were self-assembled into smaller micelles at lower critical micelle concentrations than that by conventional radical polymerization. Similarly the acrylate macromonomers have been used to design the structures with softer backbone (in comparison to the polymethacrylate ones) and mixed side chains of mPEG and PCL functionalized with the pH-sensitive DL-malic acid providing the core
shell
corona micelles with carboxyl groups in the shell (Table 9.2 9, Fig. 9.3B) (Nikfarjam et al., 2014). Three-layered nanoparticles have also been obtained by terpolymerization of the PCL and mPEG methacrylates with methacrylamide macromonomer, that is, pH-responsive PDEAEMA (Table 9.2 10, Fig. 9.5B) (Shen et al., 2008). The middle layer formed by the ionizable grafts is insoluble at pH above 7, but it becomes positively charged and soluble via protonation at pH lower than 6.5, what promoted endocytosis for fast cellular internalization in the acidic interstitium of solid tumors and damage of the lysosomal membrane to transport the nanoparticles into the cytosol. The concept of multifunctional theranostic system has been investigated for the amphiphilic core
shell-type brushes with block structure of side chains PLA-b-PEG, which at the end were conjugated with a monoclonal antibody against CD105 (TRC105) and a macrocyclic isotope chelator for 64Cu-labeling, that is, 1,4,7-triazacyclononane-N,N0 ,N-triacetic acid (Table 9.2 11, Fig. 9.1F) (Guo et al., 2014). Their extraordinary activities were confirmed by targeting efficiency, capability of noninvasive tomography imaging, and pH-responsive cancer drug release. The advantageous cellular uptake demonstrated by the targeted micelles in CD105-positive cells, because of CD105-mediated endocytosis, was in contrast to CD105-negative cells.

9.5 Other thermoresponsive graft polymethacrylate nanocarriers As noted earlier, the thermoresponsive behavior of amphiphilic copolymers is generated by PEG/PEGMA segments, but it is also provided by other hydrophilic polymers,

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like poly(2-oxazoline)s (Hoogenboom and Schlaad, 2017) or polyacrylamides. In the first group, the most investigated polymer based on 2-ethyl-2-oxazoline (Fig. 9.4F) displays LSCT between 60 C and 70 C, which can be tuned to much broader range of 0 C
100 C. Its macromonomer, that is, oligo(2-ethyl-2-oxazoline methacrylate) (OEOxMA) has been used in synthesis of the amphiphilic linear-graft block copolymers PCL-b-POEOxMA, which were self-assembled into the micelles with encapsulated doxorubicin (DOX) to show the reduced cytotoxic effect in comparison to the free drug (Knop et al., 2013). The studies have confirmed that the molecular weight of the particular hydrophilic polymer are more influential for the efficient stabilization of hydrophilic shell than the number of repeating units. Another thermoresponsive graft copolymer consisting of polymethacrylate backbone and hydrophilic PNIPAAm (Fig. 9.4B) side chains, that is, P(MMA-co-(HEMA-g-PNIPAAm)), has indicated ability to the significant drug release at 33.5 C (Song et al., 2016).

9.6 Heterografted Janus-type carriers The chemical heterogeneity of brush copolymers is also beneficial to change solubility characteristics, thermal properties, micelle stability, or in the case of polymeric films the environment-dependent wettability. Copolymerization of two macromonomers, including (meth)acrylates, known as the “grafting through” technique is commonly used for synthesis of heterografted copolymers with the well-defined side chains, whereas the backbone length has to be controlled. The properties of copolymer can be adjusted by various nature of side chains, which are combined in one macromolecule, mostly hydrophilic versus hydrophobic, amorphous versus crystalline, soft versus hard (in relation to thermal properties), and nondegradable versus degradable. Because of variable composition, they provide phase separation of heterografts in response to selective environment/stimulus yielding the Janus-type cylindrical structures (Fig. 9.1D). This extraordinary behavior has been adopted to design drug carriers due to amphiphilic properties, which were observed for statistical copolymers of crystalline mPEGMA (n 5 23) with methacrylate of soft polydimethylsiloxane (Neugebauer et al., 2005), crystalline poly(octadecyl (meth)acrylate) (Neugebauer et al., 2006b), and degradable isotactic or atactic poly(3-hydroxybutyrate) (Neugebauer et al., 2007). However, using two macromonomers with different relative reactivities, the spontaneous gradient of both types of side chains at the backbone can be attained, for example, by copolymerization of the more reactive mPEGMA with the less reactive poly(propylene glycol)-4-nonylphenyl ether acrylate (phPPGA) (Neugebauer, 2007b).

Grafted polymethacrylate nanocarriers in drug delivery

The well-dispersed individual nanoparticles with the regularly spherical core
shell structure have been prepared for the polyacrylate brushes based on poly(t-butyl methacrylate) and PNIPAAm macromonomers. The thermoreversible solubility at LCST  29 C introduced by PNIPAAm side chains was used to activate caffeine delivery with controlled release (Luo et al., 2012a,b). In addition, tert-butyl groups can be removed by hydrolysis to deprotect carboxylic groups with formation of pHsensitive PMAA side chains. The modified brushes were characterized with slightly higher LCST ranged from 38 C to 47 C, whereas the release behavior of prednisone encapsulated in the micelles was dependent on pH and temperature (Luo et al., 2013). In another case the amphiphilic polymethacrylates containing degradable PCL and pH-sensitive PDMAEMA side chains have been self-assembled into the micelles stabilized by “core-surface cross-linking” (Xu et al., 2004). The improved nanoparticles exhibited about 100 times higher stability than the micelles obtained from corresponding amphiphilic block copolymers. Alternative option is the combination of two grafting techniques to form in the first step, the graft copolymers, by copolymerization of macromonomer (“grafting through”) and then it can be used as the multifunctional initiator due to the initiating sites located at the backbone (“grafting from”). In this way a binary and ternary graft copolymers have been designed by synthesis of loosely grafted copolymer of mPEGMA and grafting of poly(n-butyl acrylate) (Fig. 9.6A) (Neugebauer et al., 2003a) or block copolymer based on alternating graft segment of styryl-functionalized mPEG and maleimide-functionalized PCL, and grafting of polymethacrylate or polyacrylamide chains from linear segment of the modified PDMAEMA (Fig. 9.6B) (Tang et al., 2014). In the latter system comprising PNIPAAm grafts, the DOX release was significantly activated by temperature. The ternary graft copolymers (Fig. 9.6C) can be also obtained by grafting “onto” using the common “click” chemistry to join multitelechelic polymer, for example, poly(3-azido-2-hydroxypropyl methacrylate) as the main chain, with monotelechelic polymers, such as alkyne end-functionalized mPEG, poly(t-butyl acrylate), and lightsensitive poly(2-cinnamoyloxyethyl methacrylate) as the side chains (Mo et al., 2017). Both binary and ternary graft copolymers are able to change micellar morphology in the selective solvents (nonpolar vs polar) resulting a pearl-necklace, wormlike, or globular structure, where the intramolecular folding of hydrophobic grafts yields the compacted core surrounded by the hydrophilic mPEG (Neugebauer et al., 2003a; Mo et al., 2017). Polyacrylates with densely grafted polymethacrylate side chains and additionally containing carboxylic group on each backbone unit (Fig. 9.6D) were obtained from poly(methoxymethyl acrylate) modified with bromoester groups for “grafting from” by ATRP and hydrolysis to remove methoxymethyl groups (Peng et al., 2006; Gu et al., 2007). The synthesis of analogical polymer brushes has been developed by the

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Figure 9.6 Combination of various grafting strategies to complex polymethacrylate brushes. ATRP, Atom transfer radical polymerization; ROP, ring-opening polymerization.

use of “trifunctional monomer” strategy (Fig. 9.6E), which has also been convenient for the Janus-type brushes (Fig. 9.6F) (Feng and Huang, 2018). For example, “trifunctional monomers” containing a polymerizable double bonds, with all possible combinations of a hydroxyl group for ROP, an halogenoester-initiating group for ATRP, and a potential carboxyl moiety have found approach to yield biocompatible polymer brushes with polyacrylic backbone functionalized by acidic or hydroxyl groups and biodegradable PCL/PLA and/or poly(meth)acrylic side chains (Song et al., 2013; Qian et al., 2016; Jiang et al., 2014; Sun et al., 2016). A versatile polymer platform formed in the first step by polymerization of “trifunctional monomer” can be also efficient

Grafted polymethacrylate nanocarriers in drug delivery

both for “grafting from” by polymerization method dependent on the type of functional group and “grafting onto” by “click” reaction based on Cu-catalyzed alkyneazide cyclization to control grafting density. In this way, the polymer brushes with polyacrylate backbone and PDMAEA/PEG, poly(pentafluorophenyl methacrylate)/ PEG, or polystyrene/PEG have been synthesized in one shot (Xu et al., 2017).

9.7 Core
shell graft copolymers The unconventional versions of amphiphilic pluronics [polypropylene glycol (PPG)/PEG linear block copolymers], which are well known as the polymeric surfactants, have been designed by combination of a linear mPEG and polymethacrylic segments with loosely distributed PPG grafts, that is, mPEG-b-P(PPGMA-co-MMA) (Maksym and Neugebauer, 2016). The spherical micelles with encapsulated indomethacin (IMC, 15%
90%) demonstrated pH-dependent in vitro release (lower rates at pH 5 5.0 than at pH 5 7.4). The end hydroxyl groups in the side chains are convenient for further modification to balance the hydrophilic
hydrophobic ratio by extension/attachment of the hydrophilic chains (Fig. 9.7A). This strategy has been used for the hydroxy-functionalized loosely grafted copolymers of P(PPGMA-co-MMA) (Maksym-Be˛benek et al., 2015) and P(PEGMA-co-MMA) containing less than 17 mol.% of PEGMA units (Maksym-Be˛benek et al., 2014), which were insoluble in water. The extension of polyether side chains with weak polyacid PMAA segments introduced/increased the hydrophilic content, and efficiently improved the polymer

Figure 9.7 Strategies to core-shell graft copolymers with block structure of side chains.

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Figure 9.8 Self-assembling core-shell graft copolymers as drug nanocarriers. PEG, Poly(ethylene glycol); PPG, polypropylene glycol.

solubility in polar solvents, including water. Both types of amphiphilic graft copolymers containing polyether-b-polymethacrylate side chains were proposed as carriers for IMC (Maksym and Neugebauer, 2017). Depending on the polyether nature (PEG vs PPG) two different types of micelles were prepared (Fig. 9.8), but additionally the grafting degree and the content of acidic fraction have influenced on the stability, drug encapsulation efficiency, and controlled drug release. Generally slower drug release rates were monitored in acidic conditions for copolymers with long side chains and large content of acidic units because of enhanced particle stabilization and drug
polymer interactions. Similarly the amphiphilicity has been generated by grafting PMAA side chains from the copolymers of hydroxyfunctionalized caprolactone 2-(methacryloyloxy)ethyl ester (Neugebauer et al., 2013; Bury and Neugebauer, 2014). Their stable nanoparticles are promising systems in longer-term release of IMC and quercetin, for which the loading efficiency and release profiles can be adjusted by the composition of the self-assembling polymers and drug nature. The core/shell wormlike polymer brushes can be also synthesized via “grafting onto” by “click” reaction of an alkynyl-terminated block copolymer with a welldefined azido-functionalized polymethacrylate (Fig. 9.7B), for example, grafting PCL-b-mPEG onto the modified poly(glycidyl methacrylate) (Zhao et al., 2012). In comparison to the spherical micelles from linear PCL-b-mPEG, the cylindrical molecular brushes demonstrated a lower efficiency of DOX loading into the PCL core, but a faster release rate and easier entrance into HeLa and HepG2 cells within 1 hour. The supramolecular chemistry is alternative approach to link hydrophilic segment as the second block in the side chains through complementary multiple hydrogen bonding between nucleobases, which are molecular recognition moieties in drug delivery. The interaction of adenine units at the ends of PCL graft copolymer with the uracilfunctionalized mPEG was applied to achieve the amphiphilic brush copolymers, PHEMA-graft-(PCL-adenine: uracil-mPEG) (Fig. 9.7C) for the advanced anticancer drug delivery (Wang et al., 2012). Their supramolecular micelles with loaded DOX indicated low cytotoxicity and high anticancer efficacy against HeLa cancer cells, excellent biodegradability, adequate drug loading capacity (up to 70%), and an salt/pH dual-responsivity with accelerated drug release in response to the intracellular level of pH (5.0) and a high salt concentration.

Grafted polymethacrylate nanocarriers in drug delivery

9.8 Graft polymers containing disulfide linkers In response to reductive reactions, disulfides in thiol-responsive micelles are cleaved causing dissociation of micelles to smaller-size assembled structures in water. Usually the disulfide bridge(s) are introduced into the polymer by a disulfide-labeled difunctional initiator or disulfide-functionalized RAFT agent in the chain-growth processes, or disulfide-labeled diols and diacids through carbodiimide coupling or high temperature processes. Reductive (thiol-responsive) degradation of ss-based micelles activates the release of encapsulated drugs. Disulfide linkages may connect the amphiphilic block copolymers to yield symmetric triblock copolymers with monocleavable moiety in the middle block (BAss-AB, Fig. 9.9A). The cleavage of reductive disulfide linkages in the presence of reducing agents (DTT) results in the degradation of hydrophobic ss-segment in the micelle core. These brush-like triblock copolymers consisted of a hydrophobic polyacrylate block having pendant phPPG and a hydrophilic polymethacrylate block having pendant mPEG (Sourkohi et al., 2011).

Figure 9.9 Graft polymers with variety of disulfide localization. DTT, Dithiothreitol.

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The dual-sensitive brush copolymers containing centrally located disulfide linkage in the polymethacrylate backbone with statistically distributed PCL and two lengths of PEG side chains (17 vs 9) have been designed to form the self-assemblies for in vitro encapsulation and the release of DOX. The longer mPEG was linked via acidcleavable acetal moieties to HEMA units yielding three types of side chains, that is, -ss-P(mPEGMA-co-(HEMA-graft-PCL, mPEG)) (Fig. 9.9B) (Shao et al., 2013), or to hydroxyl-end capped PCL resulting in the amphiphilic block structure of side chains, which coexisted with hydrophilic homopolymer ones, that is, -ss-P(mPEGMA-co(HEMA-graft-PCL-b-mPEG)) (Fig. 9.9C) (Miao et al., 2014). The aggregates of both brushes were able to dissociate and reaggregate in the presence of acidic PBS solution or/and DTT with faster release of the encapsulated DOX in comparison to that with a lack of cleavable linkages. In the case of analogous monoresponsive copolymers the higher release rate was demonstrated by pH-sensitive than thiol-responsive systems. Response to the reductive conditions has also been confirmed for the asymmetric linear-brush block copolymer PLA-ss-PmPEGMA (Fig. 9.9D), in which the thiolresponsive disulfide linking both blocks can be cleaved (Ko et al., 2014). This concept is advantageous to tune the surface properties, including biodegradable PLA fibers, by attachment/detachment of hydrophilic PEG brush segment (hydrophobic to hydrophilic, and again to hydrophobic). The degradation rate can be significantly enhanced with the increasing amounts of disulfide linkages in the hydrophobic chains. This strategy has been used for amphiphilic block copolymers of PmPEGMA brush with linear polyesters, where the reductively degradable linkers were located repeatedly at regular intervals (Fig. 9.9E) (Nelson-Mendez et al., 2011). Larger amount of disulfide linkages can be also introduced as the pendant groups, which are labeled on the main chain (Fig. 9.9F). The polymethacrylate diblock brushes composed of the neutral graft segment of PmPEGMA and cationic PDMAEMA side chains grafted through disulfide bonds onto the second block were designed as a tumor-targeted redox-responsive degradable nonviral gene delivery vector (Han et al., 2016). This brush polymer in acidic solution was self-assembled into the micelles with mPEG/PDMAEMA mixed corona, which can be transformed to hydrophilic nanocapsules by cross-linking of the PDMAEMA chains and cleavage of the disulfides. The formation of thiol groups on the inner walls of the nanocapsules was used in the efficient functionalization with fluorescein Omethacrylate. In another system, block containing ss-linked PDMAEMA side chains was joined to the loosely PEG-grafted segment with methacrylate units functionalized by folic acid, which is the well-known cancer cell-targeting ligand (Li et al., 2014). The electrostatic interactions of cationic grafts with plasmid DNA were used to form polyplex micelles, which displayed relatively high gene transfection efficiency. In reductive conditions the disulfide linkages were effectively cleaved releasing pDNAbinding cationic PDMAEMA.

Grafted polymethacrylate nanocarriers in drug delivery

9.9 Summary A variety of polymers are intensively investigated in biomedical fields because they can be designed to achieve the well-defined molecular weight, cross-linking, (bio)degradation, self-assembling, stimuli-sensitivity, bioactive functionalization, biocompatibility, and toxicity. The state-of-the-art polymer carriers in drug delivery are responsible for effective biodistribution and controlled release without the threshold of toxicology level to minimize side effects, but remaining drug concentration on the desired therapeutic level. Generally, the sophisticated polymer architectures, including the highlighted graft polymers, give more opportunities to “tailored” nanocarrier formulations for the advanced DDSs with “drug cocktail,” multisensitivity and/ or targeting activity. The successful polymeric DDSs are already commercially available products approved by FDA, but the concepts of multifunctional DDSs are currently emerging area of the multidisciplinary studies to discover the medical products with even more unusual therapeutic possibilities than that offered on the trade market.

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295

Index Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A Acidic pH, 101, 104 105 Acitretin-loaded nanostructured lipid carriers (Act-NLC), 19, 240 241 Acne, 15 Act-NLC. See Acitretin-loaded nanostructured lipid carriers (Act-NLC) Active targeting mechanism, 143 145 Acyclovir with solid lipid nanoparticles (ACVSLNs), 13 Acylglycerol mixtures, 217 Adapalene SLNs (Ada SLNs), 19 Adaptive immunity, 163 aDDnSs. See Advanced drug delivery nanosystems (aDDnSs) Adhesiveness, 10 Adjuvants, 189 192 Advanced drug delivery nanosystems (aDDnSs), 103 104 Aerodiol, 43 Aerosolized liposomes inhalation, 172 173 AFM. See Atomic force microscopy (AFM) Albumin, 129 Alginates, 128, 262 Aliphatic polyesters, 131 134 PCL, 133 134 PLA, 132 133 PLGA, 133 polyglycolic acid, 132 All-trans retinol (AR), 20 Allopregnanolone, 38 α cyclodextrins (αCD), 29 31 α-L guluronic acid (G), 128 Alphaxalone, 39 Ammonium bisphosphonate methacrylate, 280 281 Amphiphilic polymers, 137 Amphoteric surfactants, 219 Amphotericin B loaded SLNs, 14 Analgesia, 36 37, 75 76 Animal models, 234 Anti-TNFα siRNA, 181 182

Anticancer drugs, 149 150 Antigen, 188 189 Antigen-presenting cells (APCs), 163, 188 Antimicrobials, delivery of, 13 14 Antisense oligonucleotides (ASOs), 177 AP. See Ascorbyl palmitate (AP) APCs. See Antigen-presenting cells (APCs) ApoE. See Apolipoprotein E (ApoE) Apolipoprotein E (ApoE), 180 181 Appendagael route, 73 74 Appendageal pathway. See Intrafollicular pathway Aquasomes, 2 AR. See All-trans retinol (AR) Ascorbyl palmitate (AP), 20 Asialoglycoprotein receptors (ASGPRs), 276 279 ASOs. See Antisense oligonucleotides (ASOs) Astragaloside IV, 15 Atom transfer radical polymerization (ATRP), 271 272 Atomic force microscopy (AFM), 229 ATRP. See Atom transfer radical polymerization (ATRP) Auraptene-SLNs, 16 Autologous stem cell transplantation, 39 Azelastine hydrochloride, 44 Azido-functionalized polymethacrylate, 288

B Bacillus megaterium, 130 Basal lamina, 73 β cyclodextrins (βCD), 29 31, 30f, 33 34 β-D mannuronic acid (M), 128 Bioactive lipids, 160 163 Biodegradable nanomaterials clinical applications of, 145 151 approved and investigational drugs with biodegradable polymeric nanoparticles, 148 151 properties of BNPs, 147 148 regulatory aspects, 147 future perspectives, 151 152 natural polymers, 125 130

297

298

Index

Biodegradable nanomaterials (Continued) polymeric nanoparticles, 136 145 synthetic polymers, 130 136 Biodegradable polymers (BPs), 123, 128, 130, 132, 147 inorganic, 135 136 medicinal properties, 124f Biological/physiological stimuli, 100 101 Biolubrification of damaged joints, 90 91 Biomineralization, 271 272 Biopolymers, 125, 261 γ-PGA, 130 in modified peroral drug delivery, 261 263 PHA, 130 BLP25 (synthetic mucin 1 lipopeptide), 192 193 Body mass index (BMI), 56 BPs. See Biodegradable polymers (BPs) Breathing mode, 116 Bricks, 73 Budesonide, 44 Buserelin, 43

C C12H25-poly(acrylic acid) (C12H25-PAA), 107 108 C12H25-poly(N-isopropylacrylamide)-COOH (C12H25-PNIPAM-COOH), 109 110 Caffeine, 85 86 Camellia sinensis.. See Green tea (Camellia sinensis) Camptothecin (CRLX101), 39 40 Cancer liposomal formulations for cancer vaccines, 192 193 Carbamazepine, 38 Carbopol, 13 Carbopol 934, 15 Carbopol 940 polymer, 52 53 Carbopol Ultrez 10, 19 gel, 240 Carboxymethyl-βCD (CMβCD), 50 51 Cardiolipin, 160 Catalysis, 271 272 Cationic deformable liposomes, 81 1,2-dioleoyl-3-trimethylammonium-propanebased liposomes, 77 lipids, 181 nanoparticles, 184 particles, 160 163 CDs. See Cyclodextrins (CDs)

Cellulose, 262 and cellulose derivatives, 128 ethers, 262 Central nervous system (CNS), 173 Ceramides, 76 Ceteth-20, 16 ChEMA. See Cholesteryl-functionalized methacrylate (ChEMA) Chemical penetration enhancer, 72 Chemical stability, 11, 33, 45 47, 59 60 Chemical/biochemical stimuli, 100 101 Chemically controlled systems, 252 Chewable tablets, 48 Chimeric nanosystems, 107 108 Chimeric stimuli-responsive liposomes with incorporated stimuli-responsive polymers, 103 106 pH-responsive liposomes, 104 105 thermoresponsive liposomes, 105 106 Chitin, 262 Chitosan, 127, 150, 262 Cholesterol, 56 57, 76 Cholesteryl sulfate, 76 Cholesteryl-functionalized methacrylate (ChEMA), 276 Chronotherapeutics, 249 Clodronate, 173 Clotrimazole-loaded SLN and NLC, 14 CLSM. See Confocal laser scanning microscopy (CLSM) CMβCD. See Carboxymethyl-βCD (CMβCD) CNS. See Central nervous system (CNS) Coacervation, 140 141, 224 Coenzyme Q10, 11, 21 Coil-to-globule transition, 108 109 Cold homogenization, 222 Cold shock, 103 Collagen, 72 73, 129 Confocal laser scanning microscopy (CLSM), 75 76, 85 86 Continuous venovenous hemofiltration, 34 35 Controlled release drug delivery systems, 217, 250 Conventional liposomes lipid vesicles for breaching skin barrier, 75 77 liposomal formulation in clinics, 88 90 Copolymers, 123 Coprecipitation, 141 Core shell graft copolymers, 287 288, 287f, 288f Corneocytes, 12, 73

299

Index

Cosmetic and topical applications of SLNs, 5 11, 20 22 Counterions, 220 CpG oligonucleotides, 191 192 Cremophor EL, 39 Cryo field emission SEM (CryoFESEM), 229 Cryogenic-transmission electron microscopy (cryo-TEM), 113 114, 229 Cryoprotectants, 220 Cryotherapy, 103 Crystalline poly(octadecyl (meth)acrylate), 284 Crystalline-amorphous morphology, 275 CTLs. See Cytotoxic T lymphocytes (CTLs) Cubosomes, pH-responsive, 114 115 Curcumin, 18, 173 174 curcumin-SLN, 240 Cyclic dinucleotides, 191 192 Cyclin D1, 182 Cyclodextrins (CDs), 29 CD-based sublingual formulation, 49 50 complexation, 33, 37 38 in control of obesity and hyperlipidemia, 57 60 cyclodextrin-based products in clinical practice, 33 60 in dermal formulations, 51 53 inclusion complexes, 31 33, 32f, 82 as multifunctional excipients in dosage form design, 33 53 in nasal formulations, 43 45 as novel therapeutically active pharmaceutical ingredients, 53 56 NP-C, 56 60 Sugammadex, 53 56, 54f in ocular formulations, 40 43 in oral formulations, 45 48 in oromucosal formulations, 48 51 in parenteral formulations, 34 40 structure, physiochemical properties, and toxicological profile, 29 31, 30t Cyclosporin-A, 17 Cytokines, 163 164 Cytotoxic agents, 173 Cytotoxic T lymphocytes (CTLs), 163 164

D Dacarbazine-laden nanocream (DZNC), 17 Dacarbazine-laden nanoparticle (DZNP), 17 DDS. See Drug delivery systems (DDS)

DE. See Dissolution efficiency (DE) DEAEMA. See 2-(N,N-Diethylamino)ethyl methacrylate (DEAEMA) Deformable liposomes, 78 79, 81 Degradation, 124 125 Degree of crystallinity, 229 230 Delayed release systems, 250 Dendritic cells, 160 163 DepoDur, 265 Depth of penetration, 79 80 Dermaceuticals, 1 case studies of topical delivery with lipidic nanoparticles, 13 14 cosmetic and topical applications of SLNs, 5 11 delivery of agents for other skin diseases, 15 22 drug penetration mechanism with SLNs, 12 13 lipidic nanoparticles from SLNs to NLCs, 3 5 semisolid vehicle, 13 skin penetration with SLNs, 11 12, 12f SLNs, 2 3 Dermal formulations, CDs in, 51 53 Dermis, 72 73 Desmosomes, 73 Dexamethasone, 41 43, 181 Dextran, 262 263 Dialysis, 141 Dichloromethylene bisphosphonate, 171 172 Diclofenac, 36 37 Diclofenac sodium, 40 41 2-(N,N-Diethylamino)ethyl methacrylate (DEAEMA), 276 279 Differential scanning calorimetry (DSC), 82 84, 106, 230 Diffusion, 71 72 diffusion-controlled systems, 252 Diflucortolone valerate, 17 Dihydroergotamine, 43 Dimethyl-β-cyclodextrin, 43 2-(N,N-Dimethylamino)ethyl methacrylate (DMAEMA), 276 279 DiMiL, 82 Dioleoylphosphatidylethanolamine (DOPE), 102, 180 1,2-Dipalmitoyl-sn-glycero-3-phosphocholine (DPPC), 108, 181 Dissolution efficiency (DE), 260 Distearoylphosphoethanolamine (DSPE), 181 Dithiothreitol (DTT), 279 280, 289f

300

Index

DL. See Drug loading (DL) DLS. See Dynamic light scattering (DLS) DMAAm. See N,N-dimethylacrylamide (DMAAm) DMAEMA. See 2-(N,N-Dimethylamino)ethyl methacrylate (DMAEMA) DNA vaccines, 189 DOPE. See Dioleoylphosphatidylethanolamine (DOPE) Dorzolamide, 41 43 DOTAP-based LNPs, 181 182 Double emulsion technique, 223 DOX. See Doxorubicin (DOX) Doxil, 265 liposomal pharmacology based on, 165 171 Doxorubicin (DOX), 165 171, 173, 283 284 DPPC. See 1,2-Dipalmitoyl-sn-glycero-3phosphocholine (DPPC) DPX-0907, 192 193 Drospirenone, 47 Drug delivery, 143, 236, 271 272 based on lipid nanoparticles, 238 through inhalation, 236 SLN and NLC on, 239 drug-enriched core structure, 232 drug-enriched shell structure, 232 drug-free nanovesicles, 90 91 drug-loaded nanofiber matrices, 263 drug/CD complex aggregates, 41 43 flux, 77 incorporation models, 231 233, 231f drug loading models of SLN, 231 232 nanostructured lipid carriers DL models, 232 233 penetration mechanism with SLNs, 12 13 solubility, 220 Drug delivery systems (DDS), 123, 126t, 127 128, 132, 134, 136, 144, 151, 253 254, 273 Drug loading (DL), 213, 230 Drug release classification of modified release systems according to drug release rate mechanism, 252 mathematical models for, 252 260 first-order kinetics, 254 255 Higuchi model, 255 256 Hixson Crowell model, 256 257 Korsmeyer Peppas model, 257 259 release parameters, 259 260

Weibull model, 259 zero-order kinetics, 253 254 mechanism, 251 profile, 217 DSC. See Differential scanning calorimetry (DSC) DSPE. See Distearoylphosphoethanolamine (DSPE) DTT. See Dithiothreitol (DTT) Dynamic light scattering (DLS), 110 111 DZNC. See Dacarbazine-laden nanocream (DZNC) DZNP. See Dacarbazine-laden nanoparticle (DZNP)

E Ebola virus, 189 ECN. See Econazole nitrate (ECN) Econazole, 75 76 Econazole nitrate (ECN), 14 EE. See Entrapment efficiency (EE) ELA-MAX, 88 89 Elastic liposome formulation, 183 vesicles, 80 Elastin fibers, 72 73 Elastosomes, 2 Electron spin resonance, 230 Electroporation, 72 Electrospun nanofiber mats, 263 EMLA cream, 89 90 Emulsification sonication technique, 222 Emulsifiers, 219 220 Emulsion solvent evaporation, 139 140 Encapsulation, 177 179, 188 189 Endocytosis, 179 180 Endosomal effect, 102 Endothelial nitric oxide synthase (eNOS), 182 183 Enhanced permeability and retention (EPR), 143, 145, 148 eNOS. See Endothelial nitric oxide synthase (eNOS) Entrapment efficiency (EE), 230 Epaxal, 192 Epidermis, 72 73 EPR. See Enhanced permeability and retention (EPR) Equimolar ketotifen HPβCD inclusion complex, 50 51 Estradiol, 43

301

Index

Ethanol-based lipid vesicles, 82 88 ethosomes, 82 85, 83f transethosomes, 85 86 Ethinylestradiol, 81 inclusion complex of, 47 Ethosomal carbomer gel, 84 85 Ethosomes ethanol-based lipid vesicles, 82 85, 83f liposomal formulation in clinics, 92 EU. See European Union (EU) Eudragit S100, 264 European Pharmacopeia, 31 European Union (EU), 88 Evomela, 39 Ex vivo studies, 15 Exparel, 265 Extended release systems, 250 Extracellular stimuli, 100 101 Extrusion process, 80 81

F Fatty acids, 215 FDA. See Food and Drug Administration (FDA) Fenebrutinib, 173 174 Fick’s first law, 142, 254 Fick’s second law, 142 Fickian model, 258 Film formation, 10, 11f First-order kinetics, 254 255 Flocculation, 215 Fluconazole-loaded SLNs (FLZ-loaded SLNs), 14, 239 5-Fluorouracil (5-FU), 18 FLZ-loaded SLNs. See Fluconazole-loaded SLNs (FLZ-loaded SLNs) Food and Drug Administration (FDA), 31, 147, 174 177 Fourier transform infrared spectroscopy, 85 86 Free cyclosporine A, 172 173 Free fatty acids, 76 Freeze fracture electron microscopy, 78 79 5-FU. See 5-Fluorouracil (5-FU) Functional genes, 177 Fusogenic lipid DOPE, 76

G GAMA. See Gas-assisted melting atomization (GAMA)

γ cyclodextrins (γCD), 29 32, 41 43 Gamma irradiation, 3 Gas-assisted melting atomization (GAMA), 222 Gastrointestinal tract, 251 Gelatin, 129 Gelation, 213 ionic, 141 Gene delivery, 184 interference, 177 therapy, 174 177 Gene therapy, nanoparticles for, 151 GF-SLNs. See Griseofulvin-loaded SLNs (GFSLNs) Gibbs Marangoni effect, 219 Glucagon, 44 45 Glucocorticoids, 173 D-Glucuronic acid, 127 Glyceride group, 216 Glyceryl monooleate (GMO), 112 Glycolysis, 101 102 Glycolytic metabolism, 101 102 GMO. See Glyceryl monooleate (GMO) Graft copolymers core shell, 287 288, 287f, 288f disulfide linkers, 289 290, 289f ternary, 285 Grafted polymethacrylate nanocarriers core shell graft copolymers, 287 288 graft polymers containing disulfide linkers, 289 290 heterografted Janus-type carriers, 284 287 molecular brushes, 271 273, 272f poly(ethylene glycol) and biodegradable polyester nonlinear amphiphilics, 281 283 on poly(ethylene glycol) grafted poly(meth) acrylates, 275 281 on poly(ethylene glycol) poly(meth)acrylate brushes, 273 275 thermoresponsive graft polymethacrylate nanocarriers, 283 284 “Grafting from” technique, 271 272, 285 287 “Grafting onto” technique, 271 272, 285 287 “Grafting through” technique, 271 272, 284 285 Green tea (Camellia sinensis), 18 Grifinth equation, 220 Griseofulvin, permeation of, 84 85

302

Index

Griseofulvin-loaded SLNs (GF-SLNs), 14 Gums, 262

H

“Hard fat”, 216 HBsAg. See Hepatitis B surface antigen (HBsAg) Helper T cells (TH cells), 163 164 Hemicelluloses, 262 Hepatitis B surface antigen (HBsAg), 81 Heterografted Janus-type carriers, 284 287, 286f Hexagonal HII phase structures, 179 180 High-energy methods, 222 223. See also Lowenergy methods emulsification sonication technique, 222 high-pressure homogenization, 222 hot high-shear homogenization, 223 SCF technology, 222 223 High-melting point lipid(s), 4 High-performance liquid chromatography (HPLC), 107 108 High-pressure homogenization, 222 Higuchi model, 255 256 Hixson Crowell model, 256 257 HLA. See Human leukocyte antigen (HLA) HLB. See Hydrophilic lipophilic balance (HLB) Homeostasis, 163 164 Homogeneous matrix, 231 Homopolymers, 123 Hot high-shear homogenization, 223 HPLC. See High-performance liquid chromatography (HPLC) HPβCD. See 2-Hydroxypropyl-β-cyclodextrin (HPβCD) HPγCD. See 2-Hydro-xypropyl-γ-cyclodextrin (HPγCD) HSPC bilayers, 106 107 Human leukocyte antigen (HLA), 163 164 Human skin barrier to xenobiotics, 72 74 Hyaluronan, 85 86 Hyaluronic acid, 127 2-Hydro-xypropyl-γ-cyclodextrin (HPγCD), 30 31, 40 41 Hydrophilic CD derivatives, 30 31 Hydrophilic penetration enhancers, 87 Hydrophilic lipophilic balance (HLB), 16, 217, 220 Hydrophobic drugs, 82 penetration enhancers, 87 Hydroquinone-loaded SLNs, 15

Hydroxyl groups, 29 2-Hydroxypropyl-β-cyclodextrin (HPβCD), 30 31, 34 37, 56 57 Hyperlipidemia, CDs in control of, 57 60 Hypotension, 38 39 Hypromellose, 41 43

I Ibrutinib, 173 174 Ichthyosis, 15 Idebenone, 16 IL. See Interleukin (IL) Immune modulation using liposomal gene vectors, 181 183 using small-molecule therapeutics, 171 173 Immune response modulation with liposomal delivery immune stimulation with liposomal vaccines, 184 193 immune system, 163 164 liposomal immune modulation with liposomal gene vectors, 174 184 with small-molecule therapeutics, 164 174 LNPs, 159 163 Immune stimulation with liposomal vaccines, 184 193 clinical pipeline of LNP-based vaccines, 185t liposomal adjuvants, 189 192, 190f liposomal vaccine clinical trials, 192 193 liposomal vaccines, 188 189 Immune system, principles of, 163 164 Immunomodulatory LNPs, 164 Immunosuppressants, 181 Immunotherapy, 159, 182 183 Imperfect type of NLC, 232 In vitro liposomal gene delivery, 179 180 In vivo liposomal gene delivery, 180 181 In vivo studies, 84 85 Inclusion complexes of CDs, 31 33, 32f Indomethacin (IND), 16, 107 108 INF. See Interferon (INF) Inflexal V, 192 Innate immunity, 163 Inorganic BPs, 135 136 polyphosphates, 136 polyphosphazenes, 135 136 Insulin, 52 53 Intact vesicles, 78 79 Intercellular route, 11 12, 73 74 Interferon (INF), 77

303

Index

Interleukin (IL), 150 IL-10, 183 Intrafollicular pathway, 12 Intranasal delivery, 238 Invasomes, 86, 87f Ionic gelation, 141 Ionic surfactants, 219 Ionizable cationic lipids, 177 179 Iontophoresis, 72 Isoceteth-20, 16 Itraconazole injectable formulation of, 35 oral solution, 45 46

K Keratinocytes, 73 Ketoconazole-loaded SLNs and NLCs, 14 Ketoprofen-loaded transfersomal gel, 90 91 Ketotifen fumarate fast-dissolving sublingual tablets, 50 51 Korsmeyer Peppas model, 257 259

L Lactic acid fermentation, 101 102 Langerhans cells, 73 Large lipid nanoparticles, 234 Laser diffraction (LD), 225 229 Latanoprost ophthalmic solution, 41 LCST. See Lower critical solution temperature (LCST) LD. See Laser diffraction (LD) Lecithin, 18 Letermovir, 38 Leuconostoc mesenteroids, 262 263 Lidocaine, 75 76, 89 90 Lidocaine-loaded transfersomes, 81 Lightsensitive poly(2-cinnamoyloxyethyl methacrylate), 285 Lilial-functionalized methacrylate (LILMA), 276 Lipid nanoparticles, 213 214, 233 234, 236 lipid nanoparticle-based mRNA vaccine, 189 NLC, 213 214 SLN, 213 Lipid polarity, 217 Lipid vesicles for breaching skin barrier, 74 88 conventional liposomes, 75 77 ethanol-based lipid vesicles, 82 88 transfersomes, 78 82 liposomal formulation in clinics, 88 92

(trans)dermal drug-delivery systems, 71 74 Lipid-based nanoparticle (LNPs), 14, 159 163, 182 183 lipid and liposome compositional landscape, 161f lipids commonly used in liposomal formulations, 162f Lipidic nanoparticles case studies of topical delivery with, 13 14 marketed products, 8t patents, 6t from SLNs to NLCs, 3 5, 4f Lipids, 214 219 polymorphic state, 214 215 presence of liquid lipid, 218 219 proportion, 218 types, 215 217, 215f Lipoplexes, 179 180 preparation method, 177 179 Lipopolysaccharide (LPS), 191 Liposomal adjuvants, 189 192, 190f Liposomal drug delivery, 164, 264 265 Liposomal formulation in clinics, 88 92 Liposomal gene delivery technologies, 189 Liposomal gene vectors, liposomal immune modulation with, 174 184 cationic lipids for gene delivery, 178f clinical pipeline of nucleic acid therapeutics delivery, 175t future directions, 183 184 immune modulation using liposomal gene vectors, 181 183 liposomal gene delivery, 177 179 in vitro liposomal gene delivery, 179 180 in vivo liposomal gene delivery, 180 181 Liposomal immune modulation with liposomal gene vectors, 174 184 with small-molecule therapeutics, 164 174, 165f clinical pipeline of small-molecule therapeutics, 166t future directions in, 173 174 immune modulation using small-molecule therapeutics, 171 173 liposomal pharmacology based on Doxil, 165 171 Liposomal lidocaine medical products, 90 Liposomal pharmacology based on Doxil, 165 171 Liposomal vaccines, 188 189

304

Index

Liposomal vaccines (Continued) clinical trials, 192 193 immune stimulation with, 184 193 strategies, 159 Liposomal-modified release formulations in clinical use, 264 265 Liposomes, 2, 74, 103 104, 107 108, 159 163 for drug delivery, 159 pH-responsive, 104 105 thermoresponsive, 105 106 Lipospheres, 224 Lipovacin-MM, 192 193 Liquid crystals, 106 Liquid lipid, 218 219 compatibility between solid lipids and, 221 LMX4, 88 89 LNPs. See Lipid-based nanoparticle (LNPs) Lobules, 72 73 Local pulmonary therapy, 237 Lornoxicam ODT, 50 51 Low-energy methods, 223 224. See also Highenergy methods coacervation, 224 double emulsion, 223 membrane contractor technique, 224 microemulsion technique, 223 phase inversion technique, 224 Lower critical solution temperature (LCST), 105 106, 274 LPS. See Lipopolysaccharide (LPS) Lybridos, 50

M MA. See Market authorization (MA) MAA. See Methacrylic acid (MAA) Macrolide antibiotic azithromycin, 174 Magnetic resonance imaging (MRI), 103, 280 281 Major histocompatibility complex (MHC). See Human leukocyte antigen (HLA) Market authorization (MA), 90 91 Marqibo, 265 Mathematical models for drug release, 252 260 Matrix tablet systems, 261 mBSA. See Methylated bovine serum albumin (mBSA) MDT. See Mean dissolution time (MDT) ME. See Microemulsions (ME) Mean dissolution time (MDT), 260 Medical diagnosis, 271 272

Melanocytes, 73 Membrane contractor technique, 224 Merkel cells, 73 Messenger RNA (mRNA), 174 177 Methacrylic acid (MAA), 279 280 (3-(Methacryloylamino) propyl) trimethylammonium chloride, 276 Methotrexate, 16 17, 171, 173 Methoxy-functionalized PEG methacrylates (mPEGMA), 273 274 Methyl methacrylate (MMA), 274 Methylated bovine serum albumin (mBSA), 173 174 Micelles, 271 272 Miconazole (MN), 14 Miconazole-loaded SLNs (MN-SLN), 14 Microemulsions (ME), 2, 22, 223 Micrometer-range carriers, 2 Microneedles, 72 MicroRNA (miRNA), 177 immunomodulatory potential of, 182 Midazolam, 43 44 Minoxidil, 75 76 miRNA. See MicroRNA (miRNA) MLO. See Monolinolein lipid (MLO) MMA. See Methyl methacrylate (MMA) MN. See Miconazole (MN) MN-SLN. See Miconazole-loaded SLNs (MNSLN) Modified extrusion assay, 80 81 Modified peroral drug delivery biopolymers in, 261 263 nanofibers in, 263 264 Modified release systems, 249 advantages and limitations, 250 251 biopolymers in modified peroral drug delivery, 261 263 classification, 252 historical review of, 251 liposomal-modified release formulations, 264 265 mathematical models for drug release, 252 260 modified release dosage forms formulation, 251 nanofibers in modified peroral drug delivery, 263 264 release profiles comparison, 260 261 terminology, 250 Monolinolein lipid (MLO), 116 Monolithic systems, 252 Mononuclear phagocytic system (MPS), 145

305

Index

Monophosphoryl lipid A (MPLA), 191 Morphine, 43 Mortar, 1 2, 73 mPEGMA. See Methoxy-functionalized PEG methacrylates (mPEGMA) MPLA. See Monophosphoryl lipid A (MPLA) MPS. See Mononuclear phagocytic system (MPS) MRI. See Magnetic resonance imaging (MRI) mRNA. See Messenger RNA (mRNA) mTOR, 173 174 Mucilages, 262 Multilamellar vesicles, 75 76 Multiple oil-in-solid fat-in-water type (O/F/W type), 233 Mycobacterium tuberculosis, 151, 191 192

N N,N-dimethylacrylamide (DMAAm), 276 N-acetyl-glucosamine, 127 N/P ratio. See Nucleic acid phosphates ratio (N/P ratio) NABs. See NP albumin-bounds (NABs) Nano colloidal systems, 3 Nanocarriers, 72, 74 Nanoemulgel-based SLN, 15 Nanoemulsions, 2 Nanofiber-based capsules, 264 Nanofibers in modified peroral drug delivery, 263 264 Nanolithography, 271 272 Nanomaterial carriers, 251 Nanometer-range carriers, 2 Nanoparticles, 215, 223 224, 271 272 morphology, 229 for oral delivery, 150 151 stability during storage, 217 for vaccines and gene therapy, 151 Nanoprecipitation, 141 Nanostructured lipid carriers (NLCs), 4, 213 214, 235 case studies in humans for medical applications, 239 241 DL models, 232 233 amorphous type, 233 imperfect type, 232 O/F/W type, 233 formulation procedures, 221 225 lipidic nanoparticles from SLNs to, 3 5

structural types, 233f Nanosystems pH-responsive, 101 102 thermoresponsive, 102 103 Nanotechnology, 99, 261 Naproxen-loaded SLNs, 16 Nasal formulations, CDs in, 43 45 Natural polymers, 125 130, 261 approved and investigational drugs with, 148 151 biopolymers of bacterial origin, 129 130 polysaccharides, 126 128 proteins, 128 129 Neem oil, 17 Neisseria meningitides, 192 Neurodegenerative diseases, 238 Niemann Pick disease, type C (NP-C), 56 60 CDs in control of obesity and hyperlipidemia, 57 60 NPC1 and NPC2, 56 57 Nile red-loaded ethosomes, 82 84 Nimesulide/βCD complex, 47 Niosomes, 2 NLCs. See Nanostructured lipid carriers (NLCs) NOD-like receptors, 163 Non-Fickian diffusion, 259 Non-Fickian models, 258 Nondeformable liposomes, 81 Nonionic emulsifiers, 219 Nonionic surfactants, 219 Nonsteroidal anti-inflammatory drugs (NSAIDs), 88 Nonviral gene delivery, 177 179 Noyes and Whitney equation, 254 NP albumin-bounds (NABs), 149 NP-C. See Niemann Pick disease, type C (NP-C) NSAIDs. See Nonsteroidal anti-inflammatory drugs (NSAIDs) Nuclear magnetic resonance, 230 Nucleic acid phosphates ratio (N/P ratio), 179 Nucleic acids, 177 180

O O/F/W type. See Multiple oil-in-solid fat-inwater type (O/F/W type) O/W emulsions. See Oil-in-water emulsions (O/ W emulsions) Obesity, CDs in control of, 57 60

306

Index

Occlusion, 5 10 Ocular administration of drug, 237 238 Ocular formulations, CDs in, 40 43 ODTs. See Orally disintegrating tablets (ODTs) OEOxMA. See Oligo(2-ethyl-2-oxazoline methacrylate) (OEOxMA) Oil-in-water emulsions (O/W emulsions), 223, 225 Oleic acid, 240 Oleth-20, 16 Oligo(2-ethyl-2-oxazoline methacrylate) (OEOxMA), 283 284 Optoelectronics, 271 272 Oral delivery, nanoparticles for, 150 151 Oral drug delivery, 263 264 Oral formulations, CDs in, 45 48 Orally disintegrating tablets (ODTs), 48 49 Organic solvent-based approaches, 225 solvent emulsification diffusion method, 225 solvent emulsification evaporation method, 225 solvent injection method, 225 Oromucosal formulations, CDs in, 48 51 ORX-301, 57 Osmotic effect, 10 11 Osmotic systems, 253 254

P PAA. See Poly(acrylic acid) (PAA) PAC. See Poly alkyl cyanoacrylates (PAC) Paclitaxel-loaded SLNs, 18 PAMPs. See Pathogen-associated molecular patterns (PAMPs) Parabens, 220 Parenteral formulations, CDs in, 34 40 Partial glycerides, 216 Particle size reduction, 225 Passive targeting mechanism, 143 Pathogen-associated molecular patterns (PAMPs), 163 164 Pattern recognition receptors (PRRs), 163 Pazenir, 149 PBS. See Phosphate-buffered saline (PBS) PCL. See Poly(ε-caprolactone) (PCL) PDI. See Polydispersity index (PDI) PDLLA. See Poly-DL-Lactide (PDLLA) PDMAEMA-b-PLMA. See Poly(2(dimethylamino)ethyl methacrylate)-b-poly (lauryl methacrylate) (PDMAEMA-bPLMA) Pectins, 262

PEG. See Poly(ethylene glycol) (PEG) Penetration, 80 81 Penetration enhancer-containing vesicles (PEVs), 74 75, 87 Peptides, 44 45 Perturbation of SC barrier, 85 86 Pevaryl Lipogel, 75 76, 88 89 PEVs. See Penetration enhancer-containing vesicles (PEVs) Peytonie’s disease, 90 PGA. See Poly γ-glutamic acid (PGA) pH-responsive cubosomes, 114 115 liposomes, 104 105 physicochemical characterization of, 110 111 thermal analysis on, 106 109 liquid crystalline nanosystems, 115 116 nanosystems, 101 102 Pharmaceutical formulations development, 261 Pharmacokinetic analysis, 241 PHAs. See Polyhydroxyalcanoates (PHAs) 3-PHB. See Poly 3-hydroxy butyrate (3-PHB) Phase inversion technique, 224 Phenoxy-functionalized PEG acrylate (phPEGA), 275 Phosphate-buffered saline (PBS), 106 107 Phosphatidic acid, 160 Phosphatidylcholine, 160, 219 Phosphatidylethanolamine, 160 Phosphatidylglycerol, 160 Phosphatidylinositol, 160 Phosphatidylserine, 160, 180 Phospholipids, 219 Photon correlation spectroscopy, 225 229 phPEGA. See Phenoxy-functionalized PEG acrylate (phPEGA) phPPGA. See Poly(propylene glycol)-4nonylphenyl ether acrylate (phPPGA) Phytantriol (PHYT), 112 Piroxicam, 46 47 PLA. See Polylactic acid (PLA) Plain drug solution, 81 Plasmodium falciparum surface antigen (PfMSP-119), 81 PLGA. See Poly lactic-co-glycolic acid (PLGA) PnBA. See Poly(n-butylacrylate) (PnBA) PnBA-b-PAA. See Poly(n-butylacrylate)-b-poly (acrylic acid) (PnBA-b-PAA) PNIPAM. See Poly(N-isopropylacrylamide) (PNIPAM)

307

Index

PNIPAM-b-PAA. See Poly(Nisopropylacrylamide)-block-poly(acrylic acid) (PNIPAM-b-PAA) POE. See Poly orthoesters (POE) Poloxamer 188, 18 Poly 3-hydroxy butyrate (3-PHB), 130 Poly alkyl cyanoacrylates (PAC), 135, 149 Poly lactic-co-glycolic acid (PLGA), 125, 133, 149 150 Poly orthoesters (POE), 133 134 Poly γ-glutamic acid (PGA), 130 Poly(2-(dimethylamino)ethyl methacrylate)-b-poly (lauryl methacrylate) (PDMAEMA-bPLMA), 106 107 Poly(3-azido-2-hydroxypropyl methacrylate), 285 Poly(acrylic acid) (PAA), 102 Poly(ethylene glycol) (PEG), 135, 149, 273, 287 288 and biodegradable polyester, 282t nonlinear amphiphilics, 281 283 carriers based on PEG grafted poly(meth) acrylates, 275 281, 275f DOX, 280f carriers based on PEG poly(meth)acrylate brushes, 273 275, 273f in Doxil, 171 Poly(L-lysine-iso-phthalamide), 114 115 Poly(methoxymethyl acrylate), 285 287 Poly(n-butyl acrylate), 285 Poly(n-butylacrylate)-b-poly (acrylic acid) (PnBAb-PAA), 108 Poly(n-butylacrylate) (PnBA), 108 Poly(N-isopropylacrylamide) (PNIPAM), 102 103 Poly(N-isopropylacrylamide)-block-poly(acrylic acid) (PNIPAM-b-PAA), 108 109 Poly(propylene glycol)-4-nonylphenyl ether acrylate (phPPGA), 284 Poly(t-butyl acrylate), 285 Poly(t-butyl methacrylate), 285 Poly(ε-caprolactone) (PCL), 133 134, 149 150, 281 Polyanhydrides, 134 Polydispersity index (PDI), 16 Poly-DL-Lactide (PDLLA), 133 Polyglycolic acid, 132 Polyhydroxyalcanoates (PHAs), 130 Polylactic acid (PLA), 123, 132 133, 149, 281 Polymer, 114 115, 123, 125

Polymeric nanoparticles, 74, 136 145, 140f. See also Solid lipid nanoparticles (SLN) approved and investigational drugs with, 148 151 drug release mechanisms, 142 143 preparation, 139 142 coacervation, 140 141 emulsion solvent evaporation, 139 140 ionic gelation, 141 nanoprecipitation, coprecipitation, and dialysis, 141 spray drying, 142 surface adsorption and encapsulation, 139f properties, 137 139 targeting, 143 145 active, 144 passive, 143 tumor, 144f, 145, 146f Polymorphism, 229 230 Polymorphs, 214 215 Polyphosphates, 136 Polyphosphazenes, 135 136 Polypropylene glycol (PPG), 287 288 Polysaccharides, 126 128, 261 alginate, 128 cellulose, 128 chitosan, 127 hyaluronic acid, 127 starch, 128 Posaconazole intravenous formulation, 35 36 Postpolymerization, 276 279 Power law model, 258 PPG. See Polypropylene glycol (PPG) Precirol ATO5, 18, 240 241 Prednisolone, 173 Progesterone, 16, 37 38, 75 76 Prolonged action/release systems, 250 Propylamino-βCD, 41 Proteins, 44 45, 128 129 albumin, 129 collagen, 129 gelatin, 129 Protonable/deprotonable groups, 102 PRRs. See Pattern recognition receptors (PRRs) Pseudomonas, 130 Psoriasis, 15, 240 241 anthralin for, 173 Pulmonary tuberculosis treatment, 237

308

Index

Q Quercetin, 173 174, 287 288 encapsulation of, 80

R Raloxifene-loaded ethosomes, 84 85 Randomly methylated β-cyclodextrin (RAMEB), 30 31, 43, 51 53 Recrystallization processes, 215 Repartition, 71 72 Repeat action systems, 250 RES. See Reticulo endothelial system (RES) Resveratrol (RSV), 15 RSV-loaded NLC, 21 RSV-loaded SLN, 15 Reticulo endothelial system (RES), 103 104, 138 Retinol, 11 Rigid vesicles, 80 81 Ring-opening polymerization (ROP), 271 272 RISC. See RNA-induced silencing complex (RISC) RNA interference (RNAi), 174 177 RNA-induced silencing complex (RISC), 177 RNAi. See RNA interference (RNAi) ROP. See Ring-opening polymerization (ROP) RSV. See Resveratrol (RSV)

S S1P. See Sphingosine-1-phosphate (S1P) SAXS techniques. See Small-angle X-ray scattering techniques (SAXS techniques) SBEβCD. See Sulfobutylether-β-cyclodextrin (SBEβCD) SC. See Stratum corneum (SC) SC lipid liposomes (SCLL), 76 Scanning electron microscopy (SEM), 229 SCF technology. See Supercritical fluid technology (SCF technology) SCFEE. See Supercritical fluid extraction of emulsions (SCFEE) SCLL. See SC lipid liposomes (SCLL) Sebaceous glands, 13 Sebum, 13 Self microemulsifying drug delivery system, 2 Self nanoemulsifying drug delivery system, 2 SEM. See Scanning electron microscopy (SEM) Semisolid vehicle, 13 Sequessome vesicles, 90 91

Sesamol-loaded SLNs, 18 Sex hormone binding globulin (SHBG), 50 SHBG. See Sex hormone binding globulin (SHBG) Short hairpin RNA (shRNA), 177 Shunt pathway. See Intrafollicular pathway siRNA. See Small-interfering RNA (siRNA) Skin, 71 73 diseases, 1 delivery of agents for other, 15 22 hydration, 12 13 lipid vesicles for breaching skin barrier, 74 88 penetration enhancers, 78 79, 86 with SLNs, 11 12, 12f retention, 77 skin-penetrating peptides, 72 SLN. See Solid lipid nanoparticles (SLN) Small lipid nanoparticles, 234 Small-angle X-ray scattering techniques (SAXS techniques), 114 115 Small-interfering RNA (siRNA), 174 177, 182 184 lipoplexes, 81 Small-molecule therapeutics immune modulation using, 171 173 liposomal immune modulation with, 164 174, 165f Small-sized carriers, 2 Smart nanobombs, 103 “Smart” molecules, 99 SmPC. See Summary of Product Characteristics (SmPC) Sodium cholate, 219 Sodium dodecyl sulfate, 219 Soft carriers, 82 84 Solid lipid nanoparticles (SLN), 2 3, 213. See also Polymeric nanoparticles administration routes, 233 238 intranasal delivery, 238 ocular administration, 237 238 oral administration, 235 236 parenteral administration, 234 235 pulmonary administration, 236 237 topical administration, 233 234 case studies in humans for medical applications, 239 241 oral administration, 241 topical administration, 239 241

309

Index

characterization techniques coexistence of different colloidal structures, 230 degree of crystallinity and polymorphism, 229 230 EE and DL, 230 morphology, 229 particle size and size distribution, 225 229 surface charge, 229 cosmetic and topical applications, 5 11 for cosmetic applications, 20 22 drug loading models, 231 232 drug-enriched core, 232 drug-enriched shell, 232 homogeneous matrix, 231 drug penetration mechanism with, 12 13 formulation aspects, 2 3 formulation components, 214 220 lipids, 214 219 surfactants or emulsifiers, 219 220 formulation procedures, 221 225 high-energy methods, 222 223 low-energy methods, 223 224 organic solvent-based approaches, 225 lipidic nanoparticles from SLNs to NLCs, 3 5 physiological aspects, 3 preformulation studies compatibility between solid lipids and liquid lipids, 221 partitioning analysis, 221 solubility studies, 220 221 skin penetration with, 11 12, 12f Solubility of CDs, 29, 33 Solvent emulsification diffusion method, 225 Solvent emulsification evaporation method, 225 Solvent injection method, 225 methods of lipid nanoparticle production, 226t Solvent-controlled systems, 252 Sonophoresis, 72 SP. See Spironolactone (SP) Span 80, 240 Spherical geometry, 147 148 Sphingosine-1-phosphate (S1P), 160 163 Spironolactone (SP), 240 Spray drying, 142 Starch, 128 Statistical methods, 260, 265 266 Stearic acid, 240

Steroids, 216 Stimuli-responsive liposomes physicochemical properties of, 110 111 particle size of chimeric nanosystems DPPC, 112f physicochemical characterization of pHresponsive liposomes, 110 111 physicochemical characterization of thermoresponsive liposomes, 111 thermotropic behavior of, 106 110 thermal analysis of thermoresponsive liposomes, 109 110 thermal analysis on pH-responsive liposomes, 106 109 Stimuli-responsive lyotropic liquid crystalline nanosystems, 112 116 pH-responsive cubosomes, 114 115 liquid crystalline nanosystems, 115 116 using polycation of PDMAEMA, 113 114 hydrodynamic radius, 114f thermoresponsive lipid-based liquid crystalline nanosystems, 116 Stimuli-responsive nanocarriers, 99 100 chimeric stimuli-responsive liposomes, 103 106 stimuli types, 100 103 pH-responsive nanosystems, 101 102 physical and chemical stimuli that originate, 101f thermoresponsive nanosystems, 102 103 stimuli-responsive lyotropic liquid crystalline nanosystems development, 112 116 Stimuvax, 192 193 Stratum corneum (SC), 1 2, 12, 71 73 Streptococcus mutans, 192 193 Stress adaptability of deformable liposomes, 80 81 Subcutaneous tissue, 72 73 Sublingual CD-based testosterone, 50 Substituted CDs, 33 34 Sugammadex, 53 56, 54f Sulfobutylether-β-cyclodextrin (SBEβCD), 30 31, 34 35, 37 39 Summary of Product Characteristics (SmPC), 34 35 Super Case II model, 259 Supercritical fluid extraction of emulsions (SCFEE), 222

310

Index

Supercritical fluid technology (SCF technology), 222 223 Supra-vir cream, 92 Surface charge, 160 163, 229 Surfactants, 219 220 Sustained action/release systems, 250 Synthetic poly amino acids, 135 Synthetic polymers, 124 125, 130 136, 131t aliphatic polyesters, 131 134 approved and investigational drugs with, 148 151 PAC, 135 POE, 134 polyanhydrides, 134 synthetic poly amino acids, 135

T T-tube mixing technique, 177 179 Targeted delivery, 172 173 Tecemotide, 192 193 TEM. See Transmission electron microscopy (TEM) Temoporfin, 86 Ternary graft copolymers, 285 Terpenes, 86 87 Testosterone βCD complexation of, 49 50 ethosomal patches, 84 Tetracaine-loaded transfersomes, 81 3,7,11,15-Tetramethyl-1,2,3-hexadecanetriol. See Phytantriol (PHYT) TH cells. See Helper T cells (TH cells) Thermal analysis on pH-responsive liposomes, 106 109 of thermoresponsive liposomes, 109 110 Thermodynamic parameters, 106 107 Thermoresponsive amphiphilic homopolymer, 109 110 drug delivery nanosystems, 102 103 graft polymethacrylate nanocarriers, 283 284 lipid-based liquid crystalline nanosystems, 116 liposomes, 105 106 physicochemical characterization of, 111 thermal analysis of, 109 110 mPEGMA, 281 nanosystems, 102 103 Thiomersal, 220 Three-dimensional bioprinting, 102 103 Titanium dioxide, 10

TLRs. See Toll-like receptors (TLRs) TNFα. See Tumor necrosis factor α (TNFα) Tocopherol, 11 Tofacitinib, 173 174 Toll-like receptors (TLRs), 163 Topical corticosteroids, 41 43 Topical ocular delivery, 237 TPL. See Triptolide (TPL) (Trans)dermal drug-delivery systems, 71 74 human skin barrier to xenobiotics, 72 74 Transcellular route, 73 74 Transdermal drug delivery, 49 50 Transethosomes, 85 86 Transfersomes, 2 lipid vesicles for breaching skin barrier, 78 82 liposomal formulation in clinics, 90 91 Transfersulin, 81 Transmission electron microscopy (TEM), 229 Transport Case II models, 258 Trehalose monooleate, 17 Tretinoin, 77 tretinoin-loaded SLN, 15 1,4,7-Triazacyclononane- N,N',N-triacetic acid, 283 “Trifunctional monomer” strategy, 285 287 Triglycerides, 216 Triptolide (TPL), 19 Tumor necrosis factor α (TNFα), 181 182 Tumor proteins, 192 193 Tumor targeting mechanism, 144f, 145 Tween 80, 18, 85 86, 240 241

U Ultraviolet (UV), 1 attenuators, 20 radiation, 10 UV-blocking effect, 10 United States Pharmacopeia and National Formulary (USP/NF), 31

V Vaccines, nanoparticles for, 151 Vardenafil ODT, 50 51 Vaxfectin, 189 Viral gene delivery, 177 Vitamin A palmitate, 21 Vitamin E-SLN dispersion, 21 Voltarol, 40 41 Voriconazole (VCZ), 14, 34 35

311

Index

W

“Warburg effect”, 101 102 Water-immiscible organic solvents, 225 Water-in-oil microemulsion (W/O microemulsion), 224 Water-in-oil-in-water double emulsion (W/O/W double emulsion), 224 Water-miscible organic solvents, 225 Waxes, 216 217 Weibull model, 259 Western Ontario and McMaster University (WOMAC), 90 91

X X-ray diffraction pattern (XRD pattern), 214 215, 230 Xerophobia, 78 79

Z Zero-order kinetics, 253 254 Zeta potential values, 229 Ziprasidone, 37 38 Ziprazosine mesylate, 37 38 Zyrtec, 48