Engineering drug delivery systems 9780081025499, 0081025491, 9780081025482

Table of contents :
Content: 1. Drug delivery: A 21st century branch of science2. Novel drug delivery systems3. Formulation design in drug delivery4. Formulation development and characterisation5. Pharmacokinetics of drug delivery systems6. Nano- and micro-particles as drug carriers7. Implantable drug delivery systems8. 3D printed drug delivery systems9. Intelligent drug delivery systems10. Polymers and Hydrogels to Deter Drug Abuse11. Glucose sensitive drug delivery systems

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Engineering Drug Delivery Systems

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Woodhead Publishing Series in Biomaterials

Engineering Drug Delivery Systems

Edited by

Ali Seyfoddin Seyedehsar Masoomi Dezfooli Carol Ann Greene

Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom Copyright © 2020 Elsevier Ltd. 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: 978-0-08-102548-2 For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisition Editor: Sabrina Webber Editorial Project Manager: Mariana Kuhl Production Project Manager: Poulouse Joseph Cover Designer: Greg Harris Typeset by MPS Limited, Chennai, India

Contents

List of contributors

ix

1.

1

2.

Novel drug delivery systems Ahmed M. Faheem and Dalia H. Abdelkader List of abbreviations 1.1 Introduction 1.2 SMART nanocarrier-based drug delivery systems 1.2.1 Mechanisms of nanocarrier transport throughout the systemic circulation reaching the specific target 1.2.2 Types of nanocarrier-based drug delivery system 1.3 SMART extended release drug delivery system 1.4 Novel technique for nanocarrier fabrication: microfluidics 1.5 Summary and conclusion References Formulation design in drug delivery Ghada Zidan, Carol Ann Greene and Ali Seyfoddin 2.1 Introduction to formulation design 2.2 Route of administration 2.2.1 Peroral route 2.2.2 Parenteral route 2.2.3 Pulmonary route 2.2.4 Nasal route 2.2.5 Transdermal route 2.2.6 Implants 2.3 Nonbiodegradable materials for drug delivery 2.4 Biodegradable materials for drug delivery 2.4.1 Biodegradable polymer classification 2.4.2 Mechanisms of degradation 2.5 Material surface properties 2.5.1 Surface morphology 2.5.2 Physiological factors 2.5.3 Surface and interfacial energies 2.5.4 Surface charge 2.5.5 Mechanical effects 2.5.6 Adjusting material surface properties for specific applications

1 1 2 2 4 9 9 11 11 17 17 18 18 18 19 19 19 20 20 22 22 24 25 25 25 26 26 27 27

vi

Contents

2.6 2.7 2.8 2.9

3.

4.

Structural and bulk properties Material size Material shape Smart materials for drug delivery 2.9.1 Temperature-responsive polymers 2.9.2 pH-responsive polymers 2.9.3 Field-responsive polymers 2.9.4 Bio-responsive polymers 2.10 Target cites 2.10.1 Brain targeting 2.10.2 Colon targeting 2.10.3 Cancer tissues targeting 2.11 Conclusion References

29 30 31 31 32 33 33 34 35 35 35 36 37 38

Formulation development and characterization Seyedehsar Masoomi Dezfooli, Lari Dkhar, Abbey Long, Hamideh Gholizadeh, Soniya Mohammadi, Carol Ann Greene and Ali Seyfoddin 3.1 From lab to clinical trials: different stages of developing a drug delivery system 3.2 Technological aspects of a novel drug delivery system 3.2.1 Preparation of drug carriers by emulsion/suspension techniques 3.2.2 Microfabrication and molding of pharmaceuticals 3.2.3 Microfluidic technologies for drug delivery 3.2.4 Solid freeform fabrication of drug delivery systems 3.3 Physiochemical characterization of drug delivery systems 3.3.1 Morphology 3.3.2 Physical properties 3.3.3 Chemical properties 3.3.4 Stability and biodegradability 3.4 Simulating a physiological environment for in vitro drug delivery studies 3.5 In vitro and in vivo toxicity studies 3.5.1 In vitro toxicity studies 3.5.2 In vivo toxicity studies 3.6 In vivo performance evaluations 3.7 Correlation between in vivo and in vitro studies 3.7.1 Correlation levels References Further reading

43

Nano- and microparticles as drug carriers Mo´nica Cristina Garcı´a 4.1 Introduction

43 43 43 46 47 48 51 51 52 53 53 53 55 55 58 58 60 60 61 70 71 71

Contents

Micro versus nanoparticles: physicochemical properties for drug delivery 4.3 Types of carriers for drug delivery 4.3.1 Carriers based on lipids 4.3.2 Carriers based on polymers 4.4 Biomedical applications of lipid-based nanocarriers 4.5 Biomedical applications of polymer-based nanocarriers 4.6 Final remarks and future perspectives References

vii

4.2

5.

6.

7.

73 75 75 81 89 92 98 99

Implantable drug delivery systems Ian Major, Sarah Lastakchi, Maurice Dalton and Christopher McConville 5.1 Introduction 5.2 Nondegradable polymers 5.3 Nondegradable subdermal implants 5.4 Nondegradable vaginal rings 5.5 Nondegradable ocular implants 5.6 Biodegradable implants 5.7 Biodegradable polymers 5.8 Injectable in situ forming implants 5.9 Bioresorbable ceramics 5.10 Biodegradable metal alloys References Further reading

111

Three-dimensional printed drug delivery systems Lilith Mabel Caballero-Aguilar, Saimon Moraes Silva and Simon E. Moulton 6.1 Introduction 6.2 Direct write: pressure-assisted systems 6.3 Fused deposition modeling 6.4 Inkjet printing 6.4.1 Drop-on-solid inkjet printing 6.4.2 Drop-on-drop inkjet printing 6.5 Summary and future perspectives References

147

Intelligent drug delivery systems Sepehr Talebian and Javad Foroughi 7.1 Introduction 7.1.1 Active and passive drug delivery 7.1.2 Hydrogel drug delivery systems 7.1.3 Thermoplastic drug delivery systems 7.1.4 Microdevices delivery systems 7.1.5 Transdermal patches delivery systems

163

111 111 115 116 121 122 123 126 127 129 129 146

147 148 151 154 155 157 159 160

163 166 168 171 173 175

viii

Contents

7.1.6 7.1.7 References 8.

9.

Emerging therapeutic methods based on implantable drug delivery systems Summary

177 178 179

Polymers and hydrogels to deter drug abuse Samaneh Alaei, Niloofar Babanejad, Rand Ahmad and Hamid Omidian 8.1 Introduction 8.1.1 Prescription drug abuse 8.1.2 Routes of abuse 8.2 Abuse deterrent formulations 8.3 Polymer properties in the abuse-deterrent products 8.3.1 Thermal 8.3.2 Rheological 8.3.3 Swelling 8.3.4 Mechanical 8.3.5 Binding 8.3.6 Film-forming 8.4 Conclusion References

185

Glucose-sensitive materials for delivery of antidiabetic drugs Maria Saeed and Amr Elshaer 9.1 Introduction 9.1.1 Types and causes of diabetes 9.1.2 Diabetes management 9.2 Need to redesign insulin delivery systems 9.3 Glucose-sensitive materials 9.3.1 Glucose oxidase 9.4 Advantages and limitations of glucose oxidase-based systems 9.4.1 Concanavalin A 9.5 Advantages and limitations of concanavalin A-based systems 9.5.1 Phenylboronic acid 9.6 Advantages and limitations of phenylboronic acid 9.7 Conclusion References Further reading

203

Index

185 185 185 187 191 191 193 194 195 197 198 199 200

203 203 205 206 207 207 212 212 216 216 219 222 222 227 229

List of contributors

Dalia H. Abdelkader Faculty of Pharmacy, Department, Tanta University, Tanta, Egypt

Pharmaceutical

Technology

Rand Ahmad Department of Pharmaceutical Sciences, College of Pharmacy, Nova Southeastern University, Fort Lauderdale, FL, United States Samaneh Alaei Department of Pharmaceutical Sciences, College of Pharmacy, Nova Southeastern University, Fort Lauderdale, FL, United States Carol Ann Greene Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand; Department of Ophthalmology, New Zealand National Eye Centre, University of Auckland, New Zealand Niloofar Babanejad Department of Pharmaceutical Sciences, College Pharmacy, Nova Southeastern University, Fort Lauderdale, FL, United States

of

Lilith Mabel Caballero-Aguilar ARC Centre of Excellence for Electromaterials Science, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, VIC, Australia; BioFab3D@ACMD, St Vincent’s Hospital Melbourne, Fitzroy, VIC, Australia Maurice Dalton Materials Research Institute, Athlone Institute of Technology, Athlone, Ireland Seyedehsar Masoomi Dezfooli Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Lari Dkhar Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Amr Elshaer Drug Discovery, Delivery and Patient Care (DDDPC), School of Life Sciences, Pharmacy and Chemistry, Kingston University, Kingston upon Thames, London, United Kingdom Ahmed M. Faheem School of Pharmacy and Pharmaceutical Sciences, University of Sunderland, Sunderland, United Kingdom; Faculty of Pharmacy, Pharmaceutical Technology Department, Tanta University, Tanta, Egypt

x

List of contributors

Javad Foroughi Intelligent Polymer Research Institute, University of Wollongong, Wollongong, NSW, Australia Mo´nica Cristina Garcı´a Unidad de Investigacio´n y Desarrollo en Tecnologı´a Farmace´utica (UNITEFA)-CONICET-UNC, Co´rdoba, Argentina; Departamento de Ciencias Farmace´uticas, Facultad de Ciencias Quı´micas, Universidad Nacional de Co´rdoba, Ciudad Universitaria, Co´rdoba, Argentina Hamideh Gholizadeh Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Sarah Lastakchi School of Clinical and Experimental Medicine, College of Medical and Dental Sciences, University of Birmingham, Birmingham, United Kingdom Abbey Long Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Ian Major Materials Research Institute, Athlone Institute of Technology, Athlone, Ireland Christopher McConville School of Clinical and Experimental Medicine, College of Medical and Dental Sciences, University of Birmingham, Birmingham, United Kingdom Soniya Mohammadi Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Simon E. Moulton ARC Centre of Excellence for Electromaterials Science, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, VIC, Australia; BioFab3D@ACMD, St Vincent’s Hospital Melbourne, Fitzroy, VIC, Australia Hamid Omidian Department of Pharmaceutical Sciences, College of Pharmacy, Nova Southeastern University, Fort Lauderdale, FL, United States Maria Saeed Drug Discovery, Delivery and Patient Care (DDDPC), School of Life Sciences, Pharmacy and Chemistry, Kingston University, Kingston upon Thames, London, United Kingdom Ali Seyfoddin Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand Saimon Moraes Silva ARC Centre of Excellence for Electromaterials Science, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, VIC, Australia; BioFab3D@ACMD, St Vincent’s Hospital Melbourne, Fitzroy, VIC, Australia

List of contributors

Sepehr Talebian Intelligent Polymer Research Wollongong, Wollongong, NSW, Australia

xi

Institute,

University

of

Ghada Zidan Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand

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Novel drug delivery systems

1

Ahmed M. Faheem1,2 and Dalia H. Abdelkader2 1 School of Pharmacy and Pharmaceutical Sciences, University of Sunderland, Sunderland, United Kingdom, 2Faculty of Pharmacy, Pharmaceutical Technology Department, Tanta University, Tanta, Egypt

List of abbreviations EPR enhanced permeability and retention FA-DABA-SMA folic acid-2,4-diaminobutyric acid-poly(styrene-alt-maleic anhydride) MF microfluidics NDDS novel drug delivery system NPs nanoparticles PEG polyethylene glycol PLA polylactic acid PLGA polylactic-co-glycolic acid RES reticuloendothelial system

1.1

Introduction

Novel drug delivery system (NDDS) is an expression mainly associated with the formulation of new pharmaceutical forms which have optimized characteristics such as smaller particle size, higher permeability parameters, and selective site targeting. NDDSs can be used to enhance the performance of biotherapeutic agents when compared with their effect in the conventional dosage forms. This chapter will explain the concept of NDDS, the different methods of design, and some of their clinical applications. Generally, the pharmaceutical drug delivery system consists of: 1. a suitable dosage form (pharmaceutical formulations) that carries the drug into the body; 2. the release mechanism of a drug from the dosage form to the organ/cells of targeting after administration; and 3. an optimum medical device/pharmaceutical technique used for manufacturing the dosage form.

Therefore a NDDS could be obtained via: 1. formulation of SMART nanocarrier-based drug delivery systems to improve the cell selectivity for enhanced targeting and Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00001-9 © 2020 Elsevier Ltd. All rights reserved.

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Engineering Drug Delivery Systems

2. superior controlling of the duration of action (SMART extended release drug delivery systems), 3. utilization of novel techniques of manufacturing, for example, microfluidics (MF).

1.2

SMART nanocarrier-based drug delivery systems

A carrier-based drug delivery system implies that drug molecules will be loaded into vesicular and/or polymeric system. Some examples of SMART nanosystems are polymeric micelles, liposomes, dendrimers, and nanoparticles (NPs) [1]. The design of a nanosystem with optimum characteristics such as higher drug-loading capacity, smaller particle size (50
300 nm), and controllable release profile are the main goals for designing a successful pharmaceutical product [2]. The term “SMART” means that the nanocarrier drug delivery system can release the drug in response to physiological stimuli thereby targeting it to the diseased cells/tissue with an extended/controllable manner [3]. After the administration of a SMART nanosystem, there are three possible mechanisms that might occur:

1.2.1 Mechanisms of nanocarrier transport throughout the systemic circulation reaching the specific target 1.2.1.1 Passive targeting Passive targeting is the primary pathway for a colloidal nanosystem via the enhanced permeability and retention (EPR) effect. The EPR effect has been extensively investigated in previous studies. These studies illustrated that the EPR effect highly depends on the degree of vascularity and efficiency of lymphatic drainage at the site of targeting. Increased leakage of blood vessels and inefficient lymphatic drainage might enhance the EPR effect and achieve better accumulation of nanocarriers in targeted tissues [4]. Fortunately, a crucial gain from the EPR effect is that it can be used to maximize the delivery of noncolloidal systems to tumors or cancerous tissues due to their enhanced vascular permeability when compared with healthy tissues (see Fig. 1.1) [1]. One of the biggest limitations of drug transport via passive diffusion or convection is the lack of site selectivity which might lead to several side effects and drug resistance [5]. To overcome this obstacle, some novel techniques have been designed to formulate colloidal nanosystem that can actively and selectively bind with targeted cells after extravasation. This is explained as the active targeting approach [6].

1.2.1.2 Active targeting Active targeting is an advanced strategy used to ensure selectivity and specificity of SMART nanosystem to the targeted site/organ/cells. Commonly, targeting of antitumor drugs to the cancerous tissue has become the main strategy to reduce the

Novel drug delivery systems

3

Figure 1.1 Approaches of nanosystem transport via passive and active targeting.

effects the drugs might have on healthy tissues. The technique of active targeting should be considered during preparation of the SMART noncolloidal system via functionalization of the surface of the carrier with ligands, which specially bind with its corresponding receptor on the surface of the targeted cell. Ligand
receptor attachment can guarantee that the nanosystem will be optimally delivered to the diseased cells rather than the surrounding healthy tissue [7]. The attached ligand on the surface of the nanosystem can be classified as several subtypes including antibodies or parts of their fragments, nucleic acids (aptamers), and various classes of peptides. These ligands bind with their specific receptors that are densely localized on the surface of tumor cells. This approach also ensures higher cellular uptake through the endocytic pathway [8]. The affinity of binding between the ligand and the receptor which is overexpressed on the surface of the targeted cell is the most important factor affecting the delivery of the drug. After ligand
receptor interaction, two possible mechanisms might occur, the nanosystem might start to release part of its encapsulated drug in the close proximity of the target cells and act as sustained release drug reservoir or the intact nanosystem is engulfed via endocytosis and the release begins inside the cell [9]. The second mechanism is desirable to ensure efficient delivery of drug inside the cells. Recently, SMART NPs have been widely used as a nanocarrier drug delivery system for cancer therapy. The surface of SMART NPs is functionalized with specific ligands for active targeting (see Fig. 1.1). These systems take advantage of the fact that tumorous cells highly express specific receptors that can be targeted with their ligands (see Table 1.1). Some limitations regarding the clinical application of SMART NPs due to immunogenicity of the targeting ligands and impaired dose delivery because of lysosomal digestion following endocytosis remain as challenges that need to be resolved. Most studies investigate the optimum methods and designs that ensure higher drug efficacy and delivery to overcome these drawbacks [3].

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Engineering Drug Delivery Systems

Table 1.1 Some examples of SMART nanoparticles functionalized with ligands for active targeting. SMART NPs

Ligand

Receptor

Folate-based NPs

Folic acid

Biotin-based NPs Lectin-based NPs

Biotin (vitamin H) Galactose Jacalin

Hyaluronic acid
based NPs Peptide-based NPs Monoclonal antibody (mAB)-based NPs

Hyaluronic acid H2009.1 peptide IL-13 peptide EGF HER2 mABs

Transferrin-based NPs

Transferrin

Folate receptors [10] Prostate-specific membrane antigen [11] Biotin receptors [12] Asialoglycoprotein receptors [13] Thomsen
Friedenreich carbohydrate antigen [14] Glycoprotein CD44 receptor [15] Integrin αvβ6 [16] IL-13Rα2 receptor [16] EGFR [17] Anti-HER2 monoclonal antibodies [18] Transferrin receptors [19]

NP, Nanoparticle.

1.2.1.3 Responsive to stimuli targeting This is a newer targeting strategy, is still under investigation as several limitations have emerged after its clinical application. The main concept of this approach is that SMART nanosystems start to release their encapsulated drug content after exposure to an external trigger [3]. These triggers might be pH, temperature, light, ultrasound, or magnetic/electric field [10,20,21]. To optimally apply this technique, one or more of the nanosystem components should be sensitive to these triggers. The advantages of responsive to stimuli targeting include: G

G

G

G

G

enhancing of nanosystem internalization and binding to the targeted cells [3]; controlling of the drug release [22]; reducing of unwanted side effects [3]; efficient drug distribution throughout the tumor mass [23]; and improving the bioavailability of some insoluble class II/IV chemotherapeutic agents [3].

1.2.2 Types of nanocarrier-based drug delivery system 1.2.2.1 Liposomes Liposomal structure is based on the presence of internal and external zones which have different affinity to the drug molecules (see Fig. 1.2). A liposome is composed of uni/multilayers of phospholipids organized in vesicular form. The center of this vesicle has a higher affinity to encapsulate hydrophilic drugs, whereas the hydrophobic drug is incorporated in the peripheral zone between the lipid
phospholipid layers [1,24]. Liposomes are classified into niosomes, phytosomes, ethosomes, and transfersomes. The presence of nonionic surfactant with no or very low concentration

Novel drug delivery systems

5

Figure 1.2 Basic structure of classical unilayer liposome.

of phospholipid leads to formation of niosomes, which have good aqueous dispersibility and stability [7]. Transferomes are considered flexible liposomes with higher elasticity due to the presence of single chain surfactant acting as an edge activator. If ethanol is used as a main component in the preparation of liposomes, ethosomes will be produced [25]. When the phospholipids are used to encapsulate an active component that has herbal origin or extracted from plant, they are called phytosomes [26]. The liposomal vesicular shape has become an attractive form for encapsulating different categories of biotherapeutic agents, which have varying physicochemical properties and three-dimensional (3D) structures. The ability to incorporate several peptides and proteins in liposomes makes them suitable for designing vaccines and delivering cancer therapy [27]. More recently, immunoliposomes have been applied as the liposomal surface can be linked with antibodies directly or by covalent bonds with the polyethylene glycol (PEG) chain of PEGylated liposomes. The presence of PEG chain on the surface of liposomes is a new method used to protect liposomes from the reticuloendothelial system (RES), which will be discussed later. Currently, several liposome-based pharmaceutical products are available in the market (see Table 1.2) and the development of various liposome-manufacturing techniques has led to the growth of a large industry.

1.2.2.2 Polymeric micelles The configuration of polymeric micelles is composed of an external hydrophilic shell and central hydrophobic core suitable for incorporating water-insoluble drugs. The principle of this delivery system is that the hydrophilic shell can mask the nanosytem and protect it from being attacked by the immune system. This is known as the so-called Stealth effect. The Stealth effect enables the nanosytem to pass through the blood vessels with less immunogenic reaction and less uptake by macrophages of RES. This results in a longer circulating time and better kinetic

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Engineering Drug Delivery Systems

Table 1.2 Some of the FDA-approved liposome-based pharmaceutical products available in the market [28]. Drug

Route of administration

Clinical indication

Trade name

Vincristine

IV

Marqibo

Verteporphin Cytarabine

IV Spinal

Doxorubicin

IV

Morphone sulfate Amphotericin B

Epidural IV

Acute lymphoblastic leukemia Macular degeneration Neoplastic/ lymphomatous meningitis Breast/ovarian cancer and Kaposi’s sarcoma Severe pain Severe fungal infections

Daunorubicin

IV

Leukemia

Visudyne DepoCyt

Doxil/Lipodox/ Myocet DepoDur Amphotec/Abelcet/ Ambisome DaunoXome

IV, Intravenous.

stability [29]. As a result, the presence of polymeric micelles in systemic circulation will be extended, thereby enhancing the drug bioavailability for Class II and IV candidates (drugs with low solubility) [30]. The major difference between polymeric micelles and the traditional surfactant micelles is that the central core of the polymeric micelle is much larger, which leads to higher power of solubilization [31]. Polymeric micelles can be functionalized to be “SMART” for active targeting as discussed previously via linkage with targeting ligands [7]. SMART polymeric micelles could be localized into the cell, thereby enabling the modulation of cellular functions such as activity of efflux transporters, gene expression, and apoptotic signal transduction. This property of polymeric micelles has led to their wide use in delivering cancer therapy [7].

1.2.2.3 Dendrimers Dendrimers are a less common type of carrier-based nanosystem. It differs from most other vesicular systems in its 3D structure (see Fig. 1.3). It is mainly composed of a central core surrounded by a peripheral zone consisting of densely hyperbranched chains arranged to look tree-like. Dendritic components can be arranged symmetrically as building blocks and the chains can be further functionalized with targeting ligands to form “SMART” dendrimers [32] or attached with imaging contrasting agents for application in diagnostics [33]. Dendrimers are monodispersing macromolecules with nano/microsize. The central cavity of dendrimers might be hydrophilic or lipophilic depending on the nature of its individual units [7]. Dendrimers can be classified into amphiphilic dendrimers, tecto dendrimers, chiral dendrimers, and peptide dendrimers based on the structure [33]. Recently dendrimers have been used for gene delivery and transfection applications.

Novel drug delivery systems

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Figure 1.3 Classical structure of dendrimers.

1.2.2.4 Nanoparticles NPs are classified into several categories of nanomaterials with different configurations such as nanocapsules, nanospheres, nanopores, and nanoshells with particle size ranging from 20 to 250 nm [1]. Polymeric NPs can either encapsulate the drug into their internal phase, or in some cases, free drug molecules might be adsorbed on their surface for initial burst release a short time after administration [34,35]. Polymeric materials such as polylactic-co-glycolic acid (PLGA) and its coblock derivatives (PLGA-PEG-PLGA or PEG-PLGA-PEG) have been widely used in NPs preparation due to their biocompatibility and biodegradability [21]. The surface of NPs can also be bound with functionalized groups that can attach to their specific receptors on the cell membrane. The resulting SMART NPs would be actively transported to the targeted cells with higher cellular uptake capacity and better selectivity and without any negative effects on the surrounding healthy tissues [3]. SMART NPs have been improved and developed through several stages with the aim of increasing site selectivity, specificity, and cellular uptake. The first stage includes NPs transported via passive diffusion with no selectivity such as those encapsulating peptides or anticancer drugs [35,36]. Then, an active targeting approach is employed via binding with the targeting ligand attached to its specific receptors overexpressed on the surface of targeted cells. A recent trend in SMART NPs design is to combine more than one drug release mechanisms for emphasis on internalization into the targeted cells [20]. For cellular targeting, the first step is to activate the surface of SMART NPs with functionalized groups for cell targeting. Then, a linkage is formed between the targeting ligand and receptors on the cell surface. A second mechanism that uses pH-responsive polymers for encapsulating anticancer drugs can also be employed. The polymeric matrix starts to lose its architecture in the acidic medium of

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Engineering Drug Delivery Systems

cancerous tissue, enabling the drug molecules to move freely after localization inside the cancerous cell. The third targeting mechanism is to use densely positively charged polymers that will be attracted to infected cells [37]. PEG chains can also be added on the surface of NPs to extend its presence in the systemic circulation via the Stealth effect (discussed previously) [35]. In recent years, a synthetic polymeric material composed from amphiphilic copolymer polystyrene-alt-maleic anhydride functionalized with folic acid using a linkage of 2,4-diaminobutyric acid (FA-DABA-SMA) has gained widespread attention. This is a multifunction polymer that can potentially achieve maximum targeting with lower or no side effects. This polymer exhibits the three possible strategies of targeting: passive targeting, active targeting, and responsive to trigger [10,22]. FA-DABA-SMA has a hydrophobic core that encapsulates low-soluble drugs which include most of the chemotherapeutic agents and a hydrophilic shell that enable it to circulate in the blood vessels with lower immune interaction or macrophage engulfment (RES effect). The surface of the polymer is linked with folic acid to enable active targeting via binding with folic acid receptors that exist on the cell. This polymer is sensitive to pH change and at acidic pH, the microenvironment of cancer tissue, its interior chains begin to open allowing the drug particles to travel out to the nanosystem (see Fig. 1.4) [3]. The synthesis and clinical applications of FA-DABA-SMA has been widely described in the literature. The initial polymeric material is polystyrene-alt-maleic anhydride that forms the main spherical nanoshape. The ligand (folic acid) is combined with a linker (DABA) by a chemical interaction between folic acid, dicyclohexylcarbodiimide, and hydroxy succinimide to form a stable functionalized group [22]. FA-DABA-SMA has been used to encapsulate curcumin as an example of a hydrophobic anticancer drug and also because of its intense fluorescent property [38,39]. FA-DABA-SMA encapsulates curcumin into its hydrophobic core to form

Figure 1.4 Representative graph showing the configuration of folic acid-2,4-diaminobutyric acid-poly(styrene-alt-maleic anhydride) (FA-DABA-SMA) nanoparticles (NPs) at neutral/ acidic pH.

Novel drug delivery systems

9

an advanced SMART nanocarrier-based drug delivery system [40]. Preclinical studies have shown that FA-DABA-SMA NPs loaded with curcumin can be selectively targeted to pancreatic cancer cells with no toxic effect and good therapeutic outcomes [10]. It is hoped that this promising approach of using SMART FA-DABASMA NPs will be extended to clinical trials in the future.

1.3

SMART extended release drug delivery system

Controlling the release of drugs from conventional dosage forms has been a point of interest for reducing the frequency of administration and increasing patient compliance. Several pharmaceutical techniques such as dissolution matrix/reservoir [41,42], diffusion matrix/reservoir [43,44], floating systems [45], and osmotic systems [46] have been developed in this area. However, several drawbacks have been found after clinical applications including [3]: G

G

G

G

sensitivity to first pass metabolism, high cost of preparation, difficulty to deal with drugs of high molecular weights or low-soluble drugs, and irritation of gastrointestinal tract.

The EPR effect is the favorable pathway for nanocarriers to be saved from immune response and stay in the circulation for a longer time. Nanocarrier-based drug delivery formulations are usually sterically hindered and therefore they are easily detected by RES and engulfed by macrophages [47]. PEGylation is a strategy that can be used to overcome this obstacle. Addition of PEG chains to the surface of a nanocarrier system stabilizes it and enhances its circulation throughout the blood vessels via the Stealth effect. The immune system is unable to recognize the nanosystem and so no immune reaction against the nanocarrier is elicited. This enables the nanocarrier to be directed to the targeting site with an extended time [48]. A combination of PEGylation and active targeting or responsive to stimuli approaches is a novel strategy which can ensure higher selectivity and extended release. For example, immunoliposomes, used for vaccines, are linked with monoclonal antibodies for active targeting. Also, PEG chains could be bound on the surface of immunoliposomes to formulate PEGylated immunoliposomes which can escape destruction by RES and stay for prolonged time in the circulation [27]. SMART PEGylated NPs conjugated with antibodies for active targeting have shown promising results in the treatment of various types of cancers due to their extended availability in the systemic circulation [3].

1.4

Novel technique for nanocarrier fabrication: microfluidics

MF is a novel technique that has been developed to fabricate carrier-based drug delivery system and offers several advantages over the traditional pharmaceutical

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Engineering Drug Delivery Systems

methods of preparation. Nanocarriers DDS prepared by MF technique have optimum in vitro characterizations such as higher drug-loading capacity, smaller particle size, uniform size distribution, typical spherical morphology, controllable release pattern, and less side effects [49]. MF technology can be used to fabricate NPs using a single or multiple emulsion technique and incorporate the bioactive agent depending on its hydrophilic/lipophilic nature in the aqueous or the organic phase [50]. Also, MF technique can combine more than one drugs with different physicochemical properties (solubility and partition coefficient) in the same carrier [49]. The in vitro characteristics of resulting NPs can be well controlled via manipulation of the flow rate of the immiscible solvents through the MF channels and adjusting geometric parameters of the device [51]. T-junction and coflowing rate are the main operating factors that should be optimally adjusted prior to the beginning of fabrication process [52,53]. This means that each single phase containing its soluble drug of nanoemulsion would flow in a separate microchannel till they reach a point of intersection to then complete the flow in a united channel to the end of process. The shape of these channels looks like the letter “T” [52]. Recently, MF has been widely used in the area of nano/micro carrier
mediated formulations. For example, niosomes are fabricated using MF via vagarious mixing of the two immiscible solvents in MF channels producing niosomal structure with smaller vesicular size than those resulting from the traditional methods [54]. Common polymeric materials such as polycaprolactone, PEG, polyvinyl alcohol [55], and PLGA and its initial components of poly-lactic acid (PLA) [56,57] that are already used in nano/micro formulations due to their biocompatibility, biodegradability, and extended/controlled release behavior could be utilized in MF. Optimum NP configurations can be produced via good modulation of the flow of viscous polymeric materials composed of PLGA and PLGA-PEG [51]. Microparticles of PLGA loaded with bupivacaine have been prepared using a modified single emulsion solvent evaporation technique combined with the flow focusing geometry of the MF device [58]. Another technique called cross-flow membrane emulsification has also used the MF device to produce microsphere of haloperidol (psychiatric drug, dopamine antagonist) encapsulated into PLGA [59]. MF techniques might also be used for manufacturing DDS containing sensitive biological agents such as peptides/proteins. Chitosan microsphere encapsulating insulin has been prepared using MF cross-linkage technique combined with the membrane emulsification method. The resulting microspheres possess optimum particle size and maintain peptide integrity [60]. For liposomes fabrication, the flow rate and the ratio of ethanol to water solution are the main controlling factors affecting the features of the resulting liposomes [61]. Liposomes produced from MF devices have smaller particle size and higher entrapment efficiency when compared to those prepared by the thin film hydration method [62]. Functionalizing of nanocarriers with targeting ligands to form “SMART” system can also be implemented using the MF technique. MF could combine two drugs, each one entrapped in outer or inner zone depending on its hydrophilicity/lipophilicity

Novel drug delivery systems

11

properties and the whole system can then be transported via active targeting due to the attached ligand on its surface. Such complicated nanosystems are widely employed in the treatment of cancer [63,64]. For example, a combination of doxorubicin hydrochloride (hydrophilic) and paclitaxel (hydrophobic) anticancer drugs were encapsulated into PLGA NPs using modified nanoprecipitation MF technique [64]. For prostate cancer, docetaxel and prodrug of cisplatin were incorporated into PLA polymer and the surface of SMART NPs was functionalized with 10-Aptameter for active targeting [63]. MF has no found use in clinical applications other than pharmaceutical nano formulations. It is used to model and detect in vitro drug toxicological side effects and its influences on body organs. Each organ is represented as a chamber in an MF system, the drug starts to flow throughout MF channels into these chambers and any pathological/physiological changes occurring in the organs can be detected. Also this method can be applied for an individual drug or a combination of drugs [49]. The implementation and clinical applications of MF have had a great impact on the pharmaceutical industry. The MF technique has resulted in many new innovations in the manufacturing of drug delivery systems. Also MF has been used to screen therapeutic/toxicological effects of bioactive agents via its application at the level of tissue culture studies.

1.5

Summary and conclusion

Multifunctional SMART nanotechnology that combines more than one targeting mechanism and provides extended release behavior results in higher selectivity, fewer side effects, better therapeutic index, and improved patient compliance. Therefore future research should focus on nanomedicine and the clinical implementation of these SMART nanocarrier-based drug delivery systems. Nanomedicine can be optimally used for therapeutic purposes and has the potential to be applied in biomedical and diagnostic objectives.

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Formulation design in drug delivery

2

Ghada Zidan1, Carol Ann Greene1,2 and Ali Seyfoddin1 1 Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand, 2Department of Ophthalmology, New Zealand National Eye Centre, University of Auckland, New Zealand

2.1

Introduction to formulation design

The aim of drug formulation design is to develop systems suitable for the administration of therapeutics that can be maintained at a desired concentration and duration to achieve the optimal therapeutic effect. These systems are referred to as drug delivery systems (DDSs). The rising costs of research and development, together with the difficulty associated with bringing new drugs into the market because of the high clinical testing costs, and the time required for regulatory approvals, make the development of a new formulation for a currently approved drug an easier and cost-effective process. In most cases, it is beneficial to minimize the number of doses administered as this improves patient convenience and enhances dose compliance thereby leading to better clinical outcomes. This can be achieved via DDSs as it can help to control the release of medication over a longer period of time. Furthermore, recent advances in pharmaceutical formulation design can be used to develop tailored systems for the deliver therapeutics to specific cells, tissues, or organs [1]. Polymers are the backbone of DDSs; they are used as drug carriers, reservoirs, matrices, coatings, or membranes. The main role of polymers in drug delivery is to protect the incorporated therapeutics from the physiological environment, improve drug stability, and control the release rate of the drug. In ancient Greek, the prefix “poly” refers to “many,” and “mer” means “parts,” so polymers are many parts of monomers (small molecules) that are attached together. The monomers can be connected linearly or branched producing polymers with various unique properties. Polymers can be natural or synthetic and can be either degradable or nondegradable [2]. Selecting the right material for drug delivery is challenging because of the diversity of the currently available polymers that can be used directly or tailored to suit the desired clinical purpose. It is important to thoroughly understand the required chemical, interfacial, mechanical, and biological properties of the formulation before selecting a suitable biomaterial for drug delivery purposes. Each material has its own surface and bulk properties that differ according to its nature, chemical and physical structure. Although it is easier to use a material without making any alterations, sometimes this might be necessary as the bulk properties of a material Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00002-0 © 2020 Elsevier Ltd. All rights reserved.

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might be mechanically suitable for the purpose, while its surface properties such as adhesion or charge are not. In such cases, it might be necessary to modify some of the material surface properties or even combine more than one polymer, chemically or physically. Further developments in polymer science have introduced “smart” polymeric hydrogel systems that can regulate the delivery of therapeutics in response to a specific stimulus. Moreover, the polymers used in DDSs can be bioactive and provide their own therapeutic benefits. Examples of such bio actives are chitosan that has antibacterial properties and collagen that can assist in wound healing. Polymers, however, are primarily conjugated to therapeutics to modify their transport or circulation half-life characteristics, or to achieve either passive or active targeting to the desired sites [3]. Before developing a pharmaceutical formulation, it is crucial to first understand what is required of the DDS, the route of administration, the target site, type of incorporated therapeutics, and the desired rate of drug delivery.

2.2

Route of administration

2.2.1 Peroral route The most common route of administration of therapeutics is the peroral route, where drugs are delivered via ingestion of either liquid (solution, suspension, emulsion) or solid dosage forms (tablets or capsules) through the mouth. DDSs that are designed for oral delivery pass through the gastrointestinal system where the drugs can be degraded, which makes this route unsuitable for sensitive therapeutics like peptides and proteins. Encapsulating the medication with resistant microspheres can help to protect them from gastric and intestinal degradation. Coating with various pH-resistant materials such as cellulose or vinyl acetate polymers can also help protect the incorporated drugs. For example, gelatine polymeric microspheres can be coated with various concentrations of sodium alginate and cross-linked with calcium chloride to increase their resistance to the gastrointestinal medium [4
6].

2.2.2 Parenteral route The most common parenteral routes include intravenous (IV), intramuscular (IM), or subcutaneous (SC); the drug-carrying vehicle is injected into those sites to reach the systemic circulation. The IV route leads to immediate drug availability in the circulation, which is an advantage if the rapid action of the drug is required; however, this route of delivery has a short drug circulation time. Polymers used for IV delivery should be water-soluble. On the other hand, IM and SC injections usually form a drug “depot” that controls the release of the medication into the systemic circulation for a long period of time. Flupentixol is an antipsychotic medication that is formulated in oil to be administered as IM injections every 1 or 3 months [7]. Some of the most commonly used, FDA-approved biodegradable polymers that can be used to prolong the release of parenteral drugs are poly(lactic acid) (PLA),

Formulation design in drug delivery

19

poly(glycolic acid) (PGA), a copolymer of lactic and glycolic acid named poly(lactic-co-glycolic acid) (PLGA), and poly(lactide-co-glycolide) (PLG). PLGA is widely used due to its excellent biocompatibility, biodegradability, and mechanical strength. PLG can form an in situ gel subcutaneously, which prolongs the release of the entrapped drugs for several days [8]. Various molecules have been incorporated in those biodegradable polymers including vaccines, peptides, proteins, and drug micromolecules [9].

2.2.3 Pulmonary route Pulmonary DDSs are formulated for direct delivery of therapeutics to the lungs via inhalation. Pulmonary DDSs enable the direct application of drugs to the lungs, and therefore require the use of low drug doses thereby minimizing any associated side effects. Controlled DDSs are attractive options for inhalation therapies as they permit the application of inhalers only once or twice daily, making it more patient convenient. Polymer and lipid-based nanoparticles have been widely used in pulmonary DDSs [10].

2.2.4 Nasal route Nasal delivery is where the formulation is inserted through the nasal opening in the form of nasal drops or sprays for either a local or systemic action. The nasal delivery route is especially useful for delivering immunization vaccines as the nasal mucosa can filter the pathogens before entering the body. Moreover, the nasal epithelium has relatively high permeability and low enzymatic activity that facilitates absorption of the applied vaccines into the circulation. Nanocarrier DDSs can be good vehicles for delivery of vaccines as they offer protection and efficient transport of the enclosed antigens. Nanocarriers can also be designed to offer improved and effective antigen recognition by immune cells [11]. Nasal and pulmonary routes are currently used to deliver systemic drugs to the blood as an alternative to parenteral drug delivery, especially in the case of peptide and protein therapeutics. Formulations including liposomes, microspheres, and gels have been exploited for drug delivery via nasal and pulmonary routes. Factors that must be considered when formulating a DDS for these routes include biocompatibility, the ability to be transferred into an aerosol, stability against the generated forces during aerosolization, and degradation within an acceptable period of time [10].

2.2.5 Transdermal route Transdermal DDSs adhere to the skin surface to allow the passage of the active ingredients across the skin layers for either a local or systemic effect. This can be achieved via transdermal sprays, creams, patches, or implantable devices. Transdermal patches usually have the drug entrapped in a reservoir with a porous membrane covering it or enclosed within an adhesive matrix that melts with body temperature and releases the embedded medication. Transdermal delivery is limited

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to small molecules that can penetrate the skin. The most famous transdermal patch is the nicotine patch that releases small amount of nicotine in the blood to help with smoking cessation.

2.2.6 Implants Implants can be tailored to deliver therapeutics into a specific location. Implants are useful in cancer therapy as they control the delivery of drugs into the targeted tissue over a long period of time. This method of delivery is advantageous as it offers greater convenience to patients compared to conventional chemotherapy approaches. Implants having a rod, wafer, or mesh-like structure, can be inserted during surgery, whereas others can be injected into the cancerous tissue to form an “in situ” implant at body temperature [12].

2.3

Nonbiodegradable materials for drug delivery

Various nonbiodegradable polymers such as silicones, polypropylene, polyvinylpyrrolidone, polyvinyl alcohol (PVA), and ethylene vinyl acetate (EVA) are currently used in drug delivery applications. These polymers do not degrade in the body with time and must be biocompatible to be able to control the delivery of therapeutics over a long period of time. As mentioned earlier, surface modifications can be useful in improving their long-term biocompatibility. Moreover, nonbiodegradable polymers can act as drug dispersion matrices for the release of multiple drugs from microelectro mechanical systems in a preprogrammed manner [13]. Nondegradable polymers are used in vaginal or intrauterine contraceptive devices, ocular implants or inserts, transdermal implants, dental restorations, and several other long-term drug delivery devices. These devices are meant to be removed and replaced, by means of a minor operation, after the drug has been depleted over a period of weeks, months, or years. Elvax is a widely used, thermoplastic copolymer of ethylene and vinyl acetate monomers (EVA) that is safe, with excellent flexibility and film-forming properties. EVA membrane acts as a permeation barrier prolonging the release of the entrapped therapeutics. The alteration of the monomer ratios of EVA is used to control the release of various molecular weight drugs [13]. Furthermore, varying the vinyl and acetate contents in EVA can affect the crystallinity and melt temperature and thus optimize the softness of the polymer without the need for a plasticizer. This is an extra advantage that improves the safety of EVA in medical devices [14]. EVA is used in several devices including ocular inserts and intravaginal rings. Nondegradable polymers are used in several ocular inserts and implants. OCUSERT is an oblong ring that is inserted into the glaucoma patient’s lower eyelid (conjunctival sac) to deliver pilocarpine in a controlled manner to lower ocular pressure. It is composed of alginate gel and two rate-controlling membranes made of soft EVA. This method of drug delivery allows for a lower and safer dose of the

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drug is administered and is better than the traditional eye drops that suffer from short eye transit time and very low bioavailability. Moreover, inserts are more convenient for patients who are not able to frequently apply eye drops, thereby improving patient compliance which in turn leads to better clinical outcomes. The drawback of OCUSERT is that it can get expelled from the conjunctival sac during sleep and some patients may find them uncomfortable. Nonbiodegradable ocular implants aim to deliver therapeutics to the back of the eye for several years as an alternative to the repeated intraocular injections and thus offer a steady drug concentration and minimize plasma levels. However, the insertion and removal of those intraocular implants are quite invasive and can be associated with complications such as retinal detachment and intravitreal hemorrhage [15]. Nondegradable polymers are used in transdermal drug delivery devices to control the release of the enclosed medication over a long period of time. Transdermal drug delivery can be via patches applied on the skin or implants under the skin. Researchers have used various nondegradable polymers and other materials for transdermal drug delivery. Multiwalled carbon nanotube (MCNT) chemically modified with carboxymethyl guar gum has been used as a carrier for transdermal delivery of diclofenac sodium. The incorporation of the MCNT increased the entrapment efficiency and prolonged the release of the drug [16]. The Norplant system is a famous transdermal implant used to deliver the contraceptive drug Levonorgestrel in a steady state rate for a period of 5 years. Silastic membrane, a silicone elastomer made of cross-linked polydimethylsiloxane (PDMS) with a silicon dioxide filler, acts as a drug dispersion matrix and the rate-limiting barrier membrane. Transdermal drug delivery implants are usually inserted by means of a simple surgical procedure and require removal after the consumption of the loaded drugs [17]. The field of contraceptive drug delivery has benefitted from the advantages of nondegradable polymers due to the convenience and efficacy of the drug delivery devices and the safety profile of the polymers used. Moreover, avoiding the first pass of orally administered steroids in the liver protects the drugs from degradation, and thus enables the administration of a much smaller and safer dose. Transdermal implants, vaginal rings, and intrauterine devices (IUDs) are some of the drug delivery devices that are used for long-term delivery of contraceptive medication. The NuvaRing delivers etonogestrel/ethinyl estradiol via a vaginal ring made of Elvax. The ring is easily placed by the user into the vagina and lasts up to 4 weeks. Progestasert is a T-shaped IUD that delivers progesterone for 1 year. A membrane made from EVA polymer controls the delivery of the enclosed contraceptive drug from the hydrophobic silicone reservoir [14]. Nonbiodegradable polymers are very useful in the field of dental and oral medicine as they allow the delivery of growth factors from the dental restorative sealer resins and denture crowns in order to enhance tissue regeneration. Dental resins are routinely applied as part of dental treatments such as root-end filling, perforation sealing, and adhesion of fractured roots [18]. It was found that the drug delivery particles made of the nondegradable 2-hydroxyethyl methacrylate (HEMA) and a suitable cross-linker successfully delivered growth factors to the injured site and were able to adsorb proteins and water-soluble antimicrobials due to their

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hydrophilic nature thereby inhibiting bacterial infection [19]. In another study, poly-HEMA-based hydrogel particles controlled the release of bovine serum albumin and fibroblast growth factor-2 for up to 14 days. It is worth noting that nonbiodegradable polymers are more appropriate and preferred over biodegradable polymers in making restorative materials due to their ability to maintain their structural integrity in wet environments [18].

2.4

Biodegradable materials for drug delivery

The numerous advantages of biodegradable polymers have led to their wide exploitation in the drug delivery field. Biodegradation is the breakdown of materials into intermediate or end products via hydrolytic or enzymatic activity after being in contact with biological fluids. The use of biodegradable polymers eliminates the need for the removal of the DDS after the release of all the incorporated drugs as the biocompatible or nontoxic degradation products can be removed from the body by the normal physiological clearing mechanisms. These polymers are safe, biocompatible, and can control the release rate of the loaded therapeutics. The polymers chosen for a particular DDS should have the appropriate permeability for the desired application and be able to tolerate the sterilization procedure, if applicable, without any negative impact on the polymer’s physicochemical properties [1,20]. To formulate a successful DDS using biodegradable polymers, a balance between the mechanical and the degradation properties of the polymer must be maintained since polymer degradation can cause loss of its mechanical properties, and therefore affect its functionality. To assess the mechanical properties of the DDS formulation, the tensile strength, elastic modulus, and matrix integrity should be characterized. Moreover, the polymer swelling properties should be analyzed as these have a significant effect on the release rate of the drug and the degradation process of the DDS matrix [21]. To maintain a balance between the degradation and the mechanical properties of DDSs, it can be advantageous to formulate a two-component system. This system can be a combination of a hydrophilic, swellable polymer and a hydrophobic polymer with strong mechanical properties. This approach can be achieved by physical mixing of various ratios of the two polymers or by chemical synthesis of a copolymer containing fragments of both polymer chains. Chemical modifications can alter the physicochemical properties of the polymers used by enhancing the polymer mechanical properties or altering its degradation rate [21]. A widely used copolymer in DDSs is PLGA, which is made up of PLA and PGA [22].

2.4.1 Biodegradable polymer classification 2.4.1.1 Natural polymers Natural polymers are polysaccharides or proteins derived from plant, animal, marine, or microbial origins. They can differ in molecular weight, physical and

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chemical properties depending on the source, isolation, and purification techniques. They are biodegradable, biocompatible, and usually have some bioactivity. Some natural polymers have special recognition for receptor-binding ligands in cells which can help in drug targeting. However, the bioactive properties can be disadvantageous, as they have the potential to trigger immunogenic responses. Natural polymers have the advantage of forming reversible bonds with other molecules, and some also have the property of self-assembly. Examples of natural polysaccharide biodegradable polymers of plant origin are dextran and pectin. Other commonly used biodegradable polymers are hyaluronic acid which is of animal origin, xanthan gum, and alginic acid which are of microbial origin and chitosan which is of marine origin [22]. Dextran is a widely used drug carrier as it can easily undergo chemical transformation. Furthermore, the abundance of hydroxyl groups in dextran makes the system easily modifiable for drug loading and controlled release. Alginic acid is another safe, biodegradable, natural widely used polysaccharide obtained from marine brown algae. This natural polymer can be easily converted to a gel in the presence of divalent cations through ionic cross-linking [1]. Chitosan has some unique properties that make it one of the famous and widely used natural polymers. Apart from its nontoxicity, biodegradability, and good biocompatibility, it has a mechanical film-forming ability and antimicrobial activity. Chitosan microspheres are effective in protein delivery owing to their positive charge which allows the formation of polyion complex with negatively charged proteins or DNA, thereby prolonging their release. Moreover, chitosan can be chemically cross-linked to sustain the release of loaded drugs [1].

2.4.1.2 Synthetic polymers Synthetic polymers have the advantage of being biologically inert. Furthermore, their properties are more predictable than natural polymers which allows for batchto-batch uniformity. Their properties can be tailored for specific applications. Some synthetic polymers contain hydrolysable linkages such as esters, amide, peptide, urea, or anhydride in their backbone that allow their degradation in the human body. Examples of synthetic polymers include PLA, PGA, PLGA, polyethylene glycol (PEG), and PVA [22]. PLGA is an FDA-approved copolymer of PLA and PGA. It is known for its biocompatibility and strong mechanical properties and thus has been extensively used in the delivery of drugs, growth factors, proteins, and other macromolecules such as DNA, RNA, and peptides. The higher the molecular weight of PLGA, the slower its degradation and the more prolonged the release rate of the incorporated drugs. The sustained release rate of therapeutics can also be achieved by increasing the lactide content, which decreases the polymer degradation rate and results in slower drug release. However, PLGA has low entrapment efficiency for hydrophilic drugs as drugs usually get expelled from the hydrophobic polymer into the dispersing aqueous phase during mixing. To increase the hydrophilicity of PLGA, the number of glycolic acid units in the polymer must be increased. PLA and PLGA

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provide a degradation rate ranging from months to years, depending on their composition and molecular weight and therefore are suitable for use in implanted DDSs [1,23].

2.4.2 Mechanisms of degradation The two main degradation mechanisms that occur in biodegradable polymers are hydrolytic or enzymatic degradation.

2.4.2.1 Hydrolytic degradation Hydrolytic degradation occurs when a polymer absorbs water, swells, and dissolves upon further water uptake. The degree of swelling and solubility of hydrophilic polymers depends on the polymer hydrophilicity and molecular weight. Examples of these hydrophilic polymers are dextrans, poly(ethylene oxide), and PVA [20]. Some polymers have hydrolytically labile chemical bonds that break in contact with water. Polymers having functional groups that are susceptible to hydrolysis include esters, anhydrides, amides, and ureas. The degradation process is affected by polymer crystallinity, the degree of hydrophilicity, and molecular weight. Examples of such polymers are PLA, PGA, and polylactones. Hydrolytically degradable polymers have minimal variation from site-to-site and patient-to-patient compared to enzymatically degradable polymers and is the main reason why they are used in drug delivery implants [20,21]. Degradation of water-insoluble polymers can occur via polymer ionization at a specific pH. This pH-sensitive solubility phenomenon that can be achieved by the use of polyacid or base polymers can be utilized to control the release of the drug in certain pH environments. Polymers such as hydroxypropyl methylcellulose phthalate and cellulose acetate phthalate used for enteric coatings undergo ionization degradation. Some polymers can form soluble macromolecules via the ion-exchange process. Insoluble divalent metal salts of polyanions such as calcium salts of alginates form water-soluble components after ion-exchange with monovalent ions in the surrounding media [21].

2.4.2.2 Enzymatic degradation Polymers that undergo enzymatic degradation can only degrade upon contact with enzymes in the body. Such polymers can be adjusted to be organ-specific to only release the loaded drug in a specific tissue site. Adjustments to the degradation profile of the polymer can be achieved by controlling the cross-linking density of the hydrogel. Given the sensitivity and specificity of the enzymatic degradation process, temperature, pH, and ionic strength are key parameters to be considered. Moreover, highly cross-linked networks might cause steric hindrance to enzyme penetration and thus reduce the degradation rate and release of the drug. Most natural polymers such as proteins (gelatine, collagen, albumin, fibrin) and

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polysaccharides (sugar and dextran) undergo enzymatic degradation. Enzymatic degradation can also occur in some synthetic polymers polydiols and PVA [21].

2.5

Material surface properties

The natural properties of a material, whether metallic, polymeric, or ceramic, determine its applications. Where a material is in contact with body tissues, it is of paramount importance to consider its surface properties. Surface properties are related to the nature of the material, morphology, porosity, chemistry, and surface free energy.

2.5.1 Surface morphology Morphology, whether it is smooth, rough, or patterned, plays an important role in the adhesion and cellular behavior of biomaterials. Smooth and rough surfaces have different contact angles and therefore differ in how they adhere to cells. Usually, cells adhere to materials that have rough surfaces since they have a larger surface area, which in turn enhances cell proliferation. For this reason, biomaterials with rough surfaces are used when the drug delivery implant is designed to be in contact with tissues. On the other hand, in materials such as those that are used in orthopedic joints where tissue integration is not required, smooth surfaces are preferred [24]. Researchers have created model synthetic surfaces with various surface topographies to test the effect of surface morphology on cell adhesion. The adhesion properties of surface topographies of various shapes and sizes vary with the type of material and cells used [25]. For example, it was found that bone cells adhered better to the submicron compared to the nanometer-sized features of a titanium surface [26]. Moreover, the space between features can also affect the adhesion properties. It was shown that an increased spacing between features deposited with fibronectin results in the improved adhesion of cells [27]. However, adhesion on patterns of gold nanoparticles was greater when the spacings were less than 100 nm [27].

2.5.2 Physiological factors Physiological properties of biomaterials can also have an impact on the adhesion of cells to their surfaces. Lectin is one example of a selective bioadhesive carbohydrate-binding protein which is found in both animals and plants. C-type lectin is of animal origin and is found to have specificity to liver cells [25]. Some plant lectins are selectively bound to particular arrays of sugar molecules in glycosylated cell membrane components and can be used for drug delivery to the gastrointestinal tract (GIT) [28,29]. Tomato lectin-conjugated nanospheres fed to rats have shown a 50-fold increase in intestinal uptake compared with controlled nanospheres owing to the selective bioadhesive properties of tomato lectin to the intestinal mucosa [30]. Lectin-conjugated polymers have also been investigated for the selective drug

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delivery to cancerous colon tissues owing to their selective bioadhesion properties [31]. Another type of selective bioadhesion is also found in pathogenic bacterial proteins (fimbrins) that attach to the surface of the gut or bind and then enter the cells (invasins) [30,32].

2.5.3 Surface and interfacial energies The surface free energy is a type related physical property of the material that can affect its wettability and therefore its adhesion to the surrounding tissues. Metal and ceramic materials have strong chemical bonds (covalent, ionic, or metallic) holding the atoms and molecules together. Such materials have a high surface free energy as large amount of energy is released when they are dissociated into smaller pieces. On the other hand, polymers and hydrocarbons have low surface free energy as the molecules are held by the weak van der Waals forces or hydrogen bonds that can be easily broken. Biomaterials with high surface free energy have high wettability since their high surface energy creates strong attractive forces that pull liquid droplets down to their bulk, thereby enhancing their spreadability [33,34]. Once the polymer surface is wetted, and there is no contact with air molecules, as in the case of adhesion of an implant to a biological membrane, the energy between the surface of the implant and the surface of the adjacent biological membrane (interfacial energy) plays an important role in the adhesion process. Furthermore, the surrounding liquid environment affects the adhesion process between those two surfaces since it can influence the spreading of one surface over the other [25].

2.5.4 Surface charge Material chemistry (polar or nonpolar) and charge type (positive, negative, zwitterionic, or neutral) can affect the properties of a material. Biomaterials with polar/ ionic groups on their surface have either acidic or basic properties depending on the charge of the ions. Those polar surfaces can have either attractive or repulsive forces with the surrounding environment due to the formation of ionic or hydrogen bonds. Moreover, these materials tend to be hydrophilic in nature due to the ability of their surface ions or dipoles to bind with water molecules. On the other hand, materials with nonpolar surface tend to be hydrophobic due to the weak van der Waals interaction of their surface molecules with water [35,36]. Surface charges can affect the biocompatibility and cellular affinity of biomaterials. At physiological pH, the cell surface is negatively charged, which enables the adherence of positively charged polymers like chitosan to tissues. This bioadhesion property of chitosan has been used and studied in various tissues, for buccal, intestinal, intranasal, and ocular drug delivery routes. Positively charged chitosan-coated nanostructured lipid carriers loaded with acyclovir were able to extend precorneal residence time and therefore showed higher drug penetration compared with the neutral ointment and eye drop formulations [37]. Furthermore, it has been found that incorporating positive charges on HEMA gels enhanced the adhesion and spreading of cells [38]. A study performed by Andrade et al. showed that the

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strength of protein adsorption to biomaterials depends on the net charge and polarity of both the protein and the biomaterial substrate surfaces [39].

2.5.5 Mechanical effects For initial stages of biomaterial adhesion, interfacial energy and chemical interaction are required. However, to maintain the adhesion between the biomaterial and the tissue, interpenetration or mechanical interlocking between the molecules of the two surfaces is also important. Surface etching of polymers to create micropores or a rough surface facilitates adhesion of materials via an interlocking mechanical effect. This interpenetration or interdiffusion is also related to the mobility of the individual chains of biomaterial and their entanglement in the membrane they are adhered to. The swelling capacity, which is the capacity of the dried polymer to hold water, affects the chain mobility. The higher the swelling ratio of the wet-todry weights, the greater the chain mobility, and thus the higher the interpenetration tendency [40]. In general, it is important to know that the material surface properties will not only affect its biocompatibility with tissues, but also its permeability and degradability. The more hydrophilic a material is, the higher its sorption capacity, and swelling ratio. A highly hydrophilic material will therefore undergo hydrolytic degradation faster than less hydrophilic materials and can be used for the quick release of active ingredients such as in orally dispersible tablets or effervescent granules. On the contrary, hydrophobic materials are water repellent and have low degradation and erosion properties which are advantageous in orthopedic or dental implants or any other types of nonbiodegradable implants [36].

2.5.6 Adjusting material surface properties for specific applications Change to the surface properties of a biomaterial are made in order to enhance its biocompatibility or to adjust its properties which allows its use in various applications. Surface modification can be implemented to obtain a cleaner surface, promote or reduce cell adhesion, and reduce protein absorption, or bacterial adhesion. Material surface properties can be improved or even altered by chemical, physical and/or biological means [25].

2.5.6.1 Physical techniques The surface properties of a material can be altered using physical approaches such as plasma or ultraviolet ozone (UVO). These treatments can enhance surface wettability, spreadability, promote bioadhesion, and help reduce surface friction. Plasma is generated by applying an electric field to a low-pressure gas under vacuum, causing some of its electrons to leave their atoms and resulting in the generation of

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highly reactive free radicals and charged ions. Oxygen, carbon dioxide, nitrogen, argon, and helium are some gases used in this treatment process [41]. Plasma treatment can render the surface hydrophilic or hydrophobic depending on the type of gas used. Oxygen and carbon dioxide gases generate polar groups on the surface (C
OH, COOH, and C 5 O), thereby making the surface hydrophilic, which in turn enhances adhesion properties [42]. On the other hand, the use of fluorinated gases like CF4 enhances surface hydrophobicity as they introduce a polytetrafluoroethylene (PTFE)-like structure that has low surface energy. This hydrophobic property is useful in preventing the adhesion of proteins on poly (methyl methacrylate) contact lenses. Grafting of molecules on newly formed activated surface can be achieved via plasma etching. For example, PEG can be added to promote cell adhesion or a dithiol group can be used to enable attachment to gold nanoparticles [43]. Another way to enhance surface hydrophilicity is via UVO, which is generated by the dissociation of ozone (O3), under 254 nm UV light, to molecular oxygen and atomic oxygen with high surface oxidizing properties. It has been shown that PDMS UVO-treated surfaces were rougher, more hydrophilic and had better cell spreadability compared to nontreated surfaces [44]. The simplest and most cost-effective way to physically modify biomaterial surfaces is sandblasting, shot-peening, or laser-peening in order to generate porous and rougher surfaces [45]. If a surface is not hard enough to withstand mechanical roughness, the application of a rough mesh surface using polymer fibers on the nonporous surface can be another option. The generation of the tissue’s extracellular matrix on the porous mesh will eventually connect it to the nonporous biomaterial over time [46].

2.5.6.2 Chemical process Another approach for altering the surface properties of biomaterials is to chemically change the surface functional groups. Heparin immobilization on polymeric biomaterials, used in stents and vascular grafts, is an effective way of enhancing the compatibility of materials that are in contact with blood to prevent coagulation and clot formation [47]. Immobilization of heparin can be achieved on ethylene vinyl alcohol (EVAL) polymer surface via covalent bonding, using bifunctional reagents such as adipoyl chloride and hexamethylene diisocyanate (HMDI) as linkers [48]. Ionic bonding can be also used for heparin immobilization on the EVAL surface through the incorporation of N,N-diethylethylenediamine with acryloyl chloride and HMDI as linkers [49]. Linker-free covalent immobilization of heparin can be achieved via plasma treatment. This is especially used if the surface of the biomaterial is chemically inert as in the case of PTFE vascular grafts [50]. Thermal oxidation and alkali acid heat surface treatment are other types of chemical surface treatments that are performed on biomaterials to improve their surface properties. Thermal oxidation leads to the formation of a thin layer of oxide on the material surface under high temperatures. It has been useful in treating alumina and zirconia alloys coatings for orthopedic implants to overcome the problems

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associated with overlay coating techniques that result in microcracks or pores [51,52]. It has also been shown that the exposure of porous titanium alloy hip implants to an alkali acid chemical heat treatment improves its biofunctionality through modifying the surface chemistry and creating nanotopographical features on the implant surface that is in contact with the body tissues [51].

2.6

Structural and bulk properties

Material micromorphology, pore size, swelling behavior, and biodegradability control the structural and bulk properties of the polymer matrix. These properties have an important role in the mass transfer of water and drugs into and out of the matrix and thus affect the release kinetics of the incorporated drugs. In general, drug release from polymers depends on the overall effect of various factors including water and drug diffusion, drug dissolution, polymer swelling, and polymer degradation. The overall release rate depends mainly on the slowest process of those listed above, which is referred to as the “rate-limiting step” [53]. For nonbiodegradable polymers, drug release from the matrix will either be diffusion controlled or water penetration controlled. Diffusion-controlled drug release depends on the movement of drug molecules within the matrix from the higher drug concentration to the adjacent regions with lower drug concentration. Low permeability peptide drugs will be released from the nonbiodegradable matrix through the pores and channels created by the dissolved drug phase [5]. On the other hand, the release of drugs from water penetration
controlled systems depends on the swelling properties of the matrix. Swelling will increase the aqueous solvent content within the matrix and increase the mesh size, thereby facilitating the diffusion of the drug from the matrix through the swollen network [54]. In the case of biodegradable polymers, the release of drugs occurs either via surface or bulk degradation/erosion of the drug
polymer matrix, depending on the chemical structure of the polymer backbone. Erosion and degradation can occur as a surface or bulk process. Erosion is the dissolution of chain fragments in noncross-linked polymers without any chemical change in the molecular structure, while degradation is caused by the chemical cleavage of polymer covalent bonds [3]. Surface degradation/erosion occurs when the rate of erosion exceeds the rate of water permeation into the bulk of the polymer and is desirable because of its zeroorder kinetics. On the other hand, bulk degradation/erosion occurs when water molecules permeate into the bulk of the matrix at a faster rate than erosion, leading to matrix breakdown and exhibiting complex degradation/erosion kinetics. Most of the biodegradable polymers used in drug delivery undergo bulk degradation/erosion. The degradation/erosion process can be manipulated by modifying the surface area of the polymer material used or by including hydrophobic monomer units in the polymer. The microstructural design and chemical composition of materials can be used to modify polymeric matrices for DDSs [55].

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Engineering Drug Delivery Systems

Material size

The effect of particle size on the drug carrier has been widely investigated in the development of drug delivery biomaterials. Methods such as emulsification, selfassembly, and jet breaking are used in the fabrication of micro- and nanoparticles. Particle size is controlled by adjusting the polymer and surfactant concentrations, or by changing or fine-tuning the fabrication method which can involve changes in the mixing method or temperature and time of the process. The commonly used mixing methods are vortexing, homogenization, and sonication. Within the preparation technique, the nozzle/capillary diameter and the material flow rate can have an effect on the particle size produced [56]. The most important characterization tests performed on formulated nanoparticles are the particle size, zeta potential (surface charge), entrapment, and encapsulation efficiencies, as well as the drug release profile. The particle size of the carrier has a significant impact on the pharmacokinetics of the carried drug since it affects the drug’s extravasation, circulation time, intracellular trafficking, uptake, degradation, and clearance. Moreover, the transport and adhesion of the drug carrier in blood vessels, airways, or GIT is affected by its size. There are known size-dependent clearance mechanisms within the body tissues and vessels, so it is of paramount importance to identify the pathway of the drug carrier from the administration route until it reaches the target site, to formulate a suitable particle size. Nanoparticles are very well suited for IV delivery as they can easily pass through the smallest capillaries (5
6 μm) in diameter without forming aggregates or leading to embolism [57]. The size is important if a carrier is to escape phagocytosis via the mononuclear phagocyte system (MPS). The MPS is part of the immune system located in the reticular connective tissue of the body. Particles larger than 500 nm can be phagocytosed by white blood cells that are programmed to eat bacteria and foreign bodies with sizes varying from 2 to 3 μm. Microparticles can be captured by Kupffer cells of the liver or physically trapped in the capillary bed, whereas nanoparticles smaller than 100 nm can easily leave the blood vessels through the endothelial lining due to their small size. In such case, smaller particle sizes extend the circulation of the carried therapeutics and prolong their effect. Peyer’s patches is another part of the MPS that is located in the small intestine. The optimum size for Peyer’s patch uptake is usually less than 1 μm, and therefore while microparticles remain trapped in the Peyer’s patches, nanoparticles can circulate systemically [57]. Size is not the only factor protecting drug carriers from macrophages in the MPS. Hydrophobic particles are recognized as foreign and thus are rapidly taken up by the MPS, finally ending up in the liver or in the spleen. This makes hydrophobic nanoparticle drug carriers a good choice for delivering drugs to the liver. However, if sustained systemic circulation is required, the surface of the hydrophobic nanoparticles must be modified to prevent phagocytosis. The most common hydrophilic moiety used for surface modification is PEG, a hydrophilic, nonionic polymer that has excellent biocompatibility. PEG can be added to the particles either via mixing, covalent bonding, or surface adsorption [58].

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Various types of nanocarriers have been used in formulating DDSs, including metals (silver and gold), metal oxides, magnetic particles (Fe2O3 and F3O4), carbon nanotubes, nanosuspensions, micelles, liposomes, polymeric, or solid lipid nanoparticles. Liposomes have shown low encapsulation efficiency, poor storage stability, and rapid leakage of water-soluble drugs in the blood compared with solid, polymeric, or lipid-based nanoparticles. Solid lipid nanoparticles have shown the ability to control the release of therapeutics for a longer period of time compared with liposomes [59].

2.8

Material shape

The transport of particles within the body, which strongly influences their efficiency as drug carriers, is affected by particle size and shape. Spherical-shaped particles are not always the best for formulating drug delivery carriers as they are more susceptible to macrophage uptake than other shapes. Although spherical particles need to be less than 200 nm in diameter to pass through the spleen, red blood cells that are 10 μm in size succeed in passing through owing to their flexible disk shape. Several approaches such as soft lithography, self-assembly, mechanical stretching, microfluidics, or template-assisted self-assembly can be used to control the shape of the carrier [56]. It has been shown that the local geometry of the particle at the point of cell attachment determines its internalization by macrophages. A macrophage attached to a flat region of the elliptical disk particles did not succeed in internalizing particles for over 12 hours, whereas the same particles when attached to the macrophage at the pointed ends were fully internalized in a few minutes [60]. Moreover, studies show that endothelial targeting efficiency of micrometer-scale disks is better than spheres, even those of nanometer dimensions [61]. On the other hand, spherical particles would be a better choice for evading receptor-mediated endocytosis by cancer cells. It was found that the gold nanospheres with a diameter of 14 or 74 nm were taken by HeLa cells three times more frequently compared to 74 nm 3 14 nm rods [62]. Nano- or microfibers have been successfully used in drug delivery formulations due to their easy and cost-effective manufacturing technique. Electrospinning is used to produce ultrafine fibers with nano- or micro-sized diameter by applying a strong electric field on the polymer solution or melt. Compared with other DDSs formulation techniques, electrospinning offers great flexibility in the selection of materials and drugs. The electrospun fibers have a high loading capacity and high encapsulation efficiency making them efficient drug carriers. Their fibrous shape helps in wound dressing and drug delivery applications [63].

2.9

Smart materials for drug delivery

Advancements in material science have led to the development of smart polymeric hydrogel systems that can regulate the delivery of drugs in response to specific

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Engineering Drug Delivery Systems

stimuli. The ability to control the release of the incorporated therapeutics at a desired time and place offers a significant advantage in pharmaceutical drug formulation development. Although smart polymers have a tunable sensitivity to various stimuli, they suffer from a slow response time. However, given the current diversity of polymer sources and various novel synthesis techniques, tuning of polymer response time has also become achievable. Stimuli-sensitive polymers undergo an abrupt change in their physical properties in response to small environmental stimuli, and subsequently return to their original shape after removal of the stimulus. These reversible changes in the polymer properties can cause alterations in their shape, state, hydrophilic and lipophilic balance, solubility, conductivity, or affect the way the polymer interacts with solvents. Stimuli that cause modifications in smart polymers can be physical (temperature, light, magnetic and electrical fields) or chemical (pH, chemical agents, and biomaterials). Physical stimuli directly alter the energy level of the polymer/solvent system inducing a polymer response at some critical energy level, while chemical stimuli act on the molecular interactions between the polymer and the solvent (adjusting hydrophobic/hydrophilic balance) or between the polymer chains (influencing the cross-linking, backbone integrity, or electrostatic repulsion) [3,64].

2.9.1 Temperature-responsive polymers The solubility of thermo-responsive polymers abruptly changes in response to a slight change in temperature. This can enable an aqueous polymeric solution to acquire a reversible sol
gel transition near the body temperature and controls the release of the incorporated drugs. Those thermo-responsive polymers usually have a hydrophobic moiety such as methyl, ethyl, or propyl groups. They also have a critical parameter which is either a lower critical solution temperature (LCST) or an upper critical solution temperature. LCST is the temperature above which the one phase polymeric system becomes hydrophobic, leading to phase separation. These critical temperatures can be shifted with some alterations in the polymer used to achieve drug release at the desired site in a specific temperature range. Increasing the hydrophobic area in the polymer chain will lead to stronger hydrophobic interactive forces, and thus the polymer undergoes a phase collapse at a lower temperature, thereby lowering the LCST. On the other hand, increasing the hydrophilic content of the polymer chain will increase the LCST [3]. Thermosensitive polymers can be categorized as negatively thermosensitive, positively thermosensitive, or thermoreversible according to the phase response generated as a result of change in temperature. Examples of thermosensitive polymers include pluronics, polysaccharide derivatives, chitosan, N-isopropylacrylamide (NIPAAm), and PLGA
PEG
PLGA triblock copolymers. Thermo-responsive polymeric systems can effectively control the delivery of hydrophilic and lipophilic drugs in specific sites, which helps reduce the drug systemic side effects and prolong its action. However, those systems have some limitations including high-burst drug release and low mechanical strength of the gel, which lead to uncontrolled release rates of the drug. Thermosensitive polymers can be used to formulate

Formulation design in drug delivery

33

long-acting injectable formulations via forming an in situ depot of therapeutics with prolonged drug release rates [64].

2.9.2 pH-responsive polymers pH-responsive polymers are polyelectrolytes with either pendant acidic or basic groups that can accept or release protons in response to changes in pH. The ionizable groups, including carboxylate, sulfonate, or amino groups, cause a change in the ionization state in either acidic or basic environments that alters the solubility, surface activity, or the chain conformation of the hydrogel. pH-sensitive polymers are classified into weak polyacids and weak polybases. Weak polyacids accept protons at low pHs and release protons at neutral and high pHs. Examples of pHsensitive weak polyacids are poly(acrylic acid) (PAA), poly(methacrylic acid) and chitosan. These polymers have pKa values ranging from 3 to 11; they are soluble in media with a pH lower than their pKa values, while they undergo phase separation if the pH is raised above their pKa value [3]. pH-sensitive polymers are used in oral DDSs owing to the fact that the GIT has regions of varying pH [65].

2.9.3 Field-responsive polymers Field-responsive polymers are sensitive to light, electrical, magnetic, electromagnetic, or sonic fields. The polymer’s response to the applied field is quick and the release rate of the loaded drug can be controlled via altering the position and intensity of the applied signal.

2.9.3.1 Light-responsive Light-responsive polymers undergo an instant phase transition in response to light exposure. A photosensitizer absorbs the light, and this results in an increase in the local temperature. This change in temperature changes the swelling behavior of the hydrogel which then leads to the release of the loaded drugs. Light-sensitive polymers have the advantage of being water-soluble, biocompatible, and biodegradable. The light source is applied outside the body and at the site where the polymers injected, and results in the incorporated drug being released following the phase change. However, light-sensitive polymers suffer from a high initial burst release, rapid release rates, and toxicity of the unreacted monomers. Moreover, the low penetration depth of the irradiated light, the long induction periods required, and the need to use photosensitive initiators at high concentrations are limitations to their use in DDSs.

2.9.3.2 Electric field-responsive Electric field-responsive polymers change their physical properties in response to a small change in electric current. These polymers contain a relatively large concentration of ionizable groups and are also pH-sensitive. The applied electric current

34

Engineering Drug Delivery Systems

leads to anisotropic swelling or deswelling of the polymer as the charged ions move toward the cathode or the anode. This movement causes degradation or bending of the polymer chains resulting in the release of the drug. The most important consideration when using technique of drug delivery is the careful selection of the type of electric current that can release the drug without any stimulation to the nerve endings in the surrounding tissues. Examples of electro-responsive polymers are chitosan, alginate, and hyaluronic acid [64].

2.9.4 Bio-responsive polymers Bio-responsive polymers can respond to biological stimuli inherently present in the body. These polymeric systems have functional groups that can interact with biological species. The interactions lead to structural polymeric changes that cause the release of the incorporated therapeutics. The interacting functional groups can be originally present in the polymer or can be synthetically conjugated to help achieve the desired biological reaction. Bio-responsive polymers can be sensitive to glucose, enzymes, or antigens.

2.9.4.1 Glucose-responsive polymers Glucose-responsive polymers can release insulin in response to glucose stimulation, mimicking the way insulin is normally secreted in healthy bodies. Controlled release of insulin from glucose-sensitive polymers minimizes the diabetic complications that are often associated with conventional insulin injections. PAA is an example of a glucose-sensitive polymer that is sensitive to gluconic acid (a byproduct of the enzymatic degradation of glucose). The presence of gluconic acid reduces the pH of the blood resulting in the protonation of the PAA carboxylate moieties and thus insulin can be released [64].

2.9.4.2 Enzyme-responsive polymers Enzymes have an important role in maintaining physiological homeostasis; their dysregulation is associated with many diseases and pathological conditions such as inflammation, cardiovascular disease, osteoarthritis, Alzheimer’s, and cancer. Biological sites that release enzymes can be targeted using synthesized enzymeresponsive polymers for site-specific drug delivery. These polymers can be designed to integrate recognition elements that can be specifically recognized by the targeted enzymes. The most common enzymes targeted via these smart polymers are the protease and glycosidase enzymes. For proteolytic enzymes, the recognition elements of the polymers may include peptide chains/linkers or polymer
peptide conjugates with specific amino acid sequences that have specific affinity toward the enzyme of interest. Therapeutic agents can be incorporated within the polymeric material by physical mixing or via a chemical reaction. The loaded drugs are released via diffusion or following the degradation of the polymer by the enzyme of interest [66].

Formulation design in drug delivery

35

2.9.4.3 Antigen and antibody-responsive polymers Hydrogels of antigen-responsive polymers harbor a specific antibody and undergo abrupt volume changes when they come into contact with their specific freely moving antigen. Similarly, antibody-responsive hydrogels contain a specific antigen which enables them to respond when in contact with the matching free antibodies. Antigen
antibody interactions are highly specific, and hence such hydrogels are not sensitive to the temperature or pH of the environment. Several approaches have been adopted to prepare antigen- or antibody-responsive hydrogels, including physical trapping of the antibody or antigen in the hydrogel network, chemical conjugation of the antibody or antigen to the polymer network or by grafting the antigen and the antibody within the polymer matrix. In the latter technique, the binding between the antigen and its corresponding antibody causes cross-linking of the hydrogel network. In the presence of the competitive free antigen, the bound antigen is displaced and this disrupts the gel volume and causes the release of the drug [67,68].

2.10

Target cites

2.10.1 Brain targeting It can be challenging to deliver pharmaceuticals to some sites in the body, including the brain, colon, and cancer cells. Drug targeting to the brain is challenging due to the presence of the endothelial cells lining the cerebral microvasculature which forms a blood
brain barrier (BBB). This barrier protects the brain from the fluctuations in plasma composition, and from the circulating neurotransmitters and xenobiotics capable of disturbing neural function. The lipid BBB is closely packed with very tight junctions that prevent the passage of molecules and ions, and this makes it hard to deliver various Alzheimer and antiepilepsy medications to the brain. Several drugs have properties such as high lipid solubility, low molecular size, and positive charge which do not allow them to pass the BBB. A common noninvasive approach for delivering water-soluble drugs through the BBB is by formulating a lipid-soluble prodrug DDS or loading the drug into a polymeric or lipid-based nanocarriers that can pass the BBB due to their small size [69].

2.10.2 Colon targeting The colon is an organ that requires special DDS formulation design that ensures the safe arrival of therapeutics to the colon following their passage through the GIT. The long transit time and varying pH levels that the DDS faces following its ingestion until it reaches the colon lead to the degradation of the drugs before reaching their target site. Drugs targeted to reach the colon can be designed to either act locally in the colon or to get delivered to the systemic circulation. Locally acting colon therapeutics are designed for treatment of conditions such as ulcerative

36

Engineering Drug Delivery Systems

colitis, chronic inflammatory bowel disease (Crohn’s disease), and colorectal cancer. Colonic systemic absorption is also a way of delivering peptide and protein drug molecules and other drugs that are poorly absorbed from the upper GIT to the blood. Recent developments in pharmaceutical technology formulations designed for colon targeting include coating drugs with pH-sensitive polymers, embedding in bacterial degradable matrices and prodrug formation. Another common way for colon targeting is enteric-coated drugs with polysaccharide polymers that dissolve in the colon and not in the upper GIT. Polysaccharide coatings such as guar gum are pH-resistant and thus prevent the degradation of the DDS in the GIT. Entericcoated films degrade by colon-specific bacteria to release the drugs at their target site [70]. Some DDSs have recently been developed for colon drug targeting and include pressure-controlled colon delivery capsules, colon targeted drug delivery system (CODESTM), and osmotic-controlled drug delivery [71].

2.10.3 Cancer tissues targeting Most anticancer drugs have poor aqueous solubility and the potential to adversely effect on healthy tissue, thereby limiting their clinical efficacy and therapeutic safety. Cosolubilizers can be used to enhance the drug solubility, but this might increase the side effects of the medication. Moreover, to achieve therapeutic efficacy for cancer therapeutics, it is often necessary to administer high doses of drugs due to their short circulation time in the blood plasma, low specificity, and poor pharmacokinetics. Several approaches are used to achieve drug targeting of polymeric carriers in order to overcome these limitations [72].

2.10.3.1 Passive cancer targeting The passive targeting approach is based on the hypothesis that cancerous tissues have pathological properties, such as leaky vasculature, lower pH, higher temperature, and a negative surface charge. Cancerous cells promote the generation of new blood vessels resulting in leaky vessels with gap sizes of 100 nm to 2 μm. Moreover, tumor tissues exhibit poor lymphatic drainage due to the high interstitial pressure at the center of the tumor compared to the periphery. The combination of leaky vasculature and poor lymphatic flow results in enhanced permeation and retention effect, which helps nanoparticle penetration into the leaky tissues during blood circulation and preferentially localize in the cancerous tissues owing to their higher retention abilities [73]. Cancerous cells are also characterized by a high metabolic rate due to their rapid proliferation, and consequently and increased temperature of 40 C
44 C. Thermo-responsive polymers can therefore be employed to deliver drugs to these cells. Temperature-sensitive polymers exhibit a fine hydrophobic
hydrophilic balance in their structure that is disrupted when they encounter tumor tissues that have a higher temperature. The conformational structural changes can lead to phase separation (precipitation) after the polymer’s LCST is reached, resulting in the release of the entrapped drugs. Localized hyperthermia in cancerous

Formulation design in drug delivery

37

tissues can also be induced by using external physical methods such as ultrasound radiation or photothermal means [65]. Cancerous tissues have lower pH compared with healthy tissues. To maintain adequate supply of nutrients and oxygen, cancer cells use more lactic acid compared to other cells, thereby lowering the pH of the surrounding extracellular region. Utilization of pH-sensitive nanoparticulate systems that are designed to be stable at a physiologic pH of 7.4 and either swell or shrink to release the drug cargo in tumor tissues at lower pH can be a useful approach in passive cancer drug targeting [73].

2.10.3.2 Active cancer targeting Active drug targeting to cancerous cells can be achieved via carrier functionalized coatings or ligands. Nanoparticles can be coated with or attached to a hydrophilic polymer such as PEG which protects the drug from macrophages ingestion and prevents particle adsorption by blood serum proteins, and thus prolongs their circulation time. The incorporation of ligands over the nanoparticle surface is another common approach for active drug targeting. Targeting moieties include antibodies, antibody fragments, specific molecules, small peptides, and RNA aptamers that have affinity to specific receptors that are exclusive to or are overexpressed in cancer cell surface [74,75]. Bio-responsive polymers can also be used for cancer drug targeting as cancer cells express and release unique enzymes such as matrix metalloproteinases (MMPs). Formulating an albumin-bound form of doxorubicin with a MMP-2specific octapeptide sequence between the drug and the carrier that can be specifically cleaved by the MMP-2 in the cancer cells, releasing the free drug, has been shown to be effective [76].

2.11

Conclusion

The good safety profile, biocompatibility, and ease of modification of polymeric materials have enabled scientists to take great strides in the development of these materials for biomedical and drug delivery applications. Also the collaboration between polymer and pharmaceutical scientists has improved design flexibility and led to the development of novel DDSs capable of delivering sensitive therapeutics such as peptides and proteins and targeting specific tissues. Despite the progress made so far, there remains the need for developing new materials that match the specific and unique requirements of newly developed therapeutics and medical applications. For example, for applications that require materials with a certain level of biological activity, strategies to incorporate biological motifs onto synthetic polymers in the form of hybrid materials have been developed. Before selecting the appropriate material for drug delivery applications, thorough understanding of all the requirements from the DDS is important. The properties of the drug, molecular weight, stability, targeted organ, and route of

38

Engineering Drug Delivery Systems

administration need to be considered. Furthermore, the limitations associated with loading capacity, encapsulation efficiency, and release kinetics including the degradation rate of the polymer for each delivery system must be addressed and, where possible, optimized.

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Formulation development and characterization

3

Seyedehsar Masoomi Dezfooli, Lari Dkhar, Abbey Long, Hamideh Gholizadeh, Soniya Mohammadi, Carol Ann Greene and Ali Seyfoddin Drug Delivery Research Group, School of Science, Auckland University of Technology, Auckland, New Zealand

3.1

From lab to clinical trials: different stages of developing a drug delivery system

The first step in the development of a new drug delivery system (DDS) is the identification of difficulties concerning current traditional drug formulations. Once these problems are clearly defined, research involving potential new formulation options is carried out. Innovative solutions for developing new formulations are then employed to develop novel DDS which helps solve the problems with traditional formulation approaches. The DDS is thoroughly characterized for its in vitro properties, including the drug release property. In vitro characterization phase is then followed by pharmacokinetic studies in small animal experiments. Finally the new DDS must enter a translational process with the aim of producing the final formulations for clinical use. All drug delivery products are required to proceed through a stringent clinical development pathway in order to secure Food and Drug Administration (FDA) approval.

3.2

Technological aspects of a novel drug delivery system

3.2.1 Preparation of drug carriers by emulsion/suspension techniques 3.2.1.1 Pharmaceutical suspensions Suspensions are liquid preparations of uniformly dispersed solid drug particles that have poor solubility. Liquid formulations are preferred to solid dosage forms as they are easier to administer, have better bioavailability, and fewer side effects. The size of particles varies for different drugs and primarily depends on the rheological properties of the formulation and the physiochemical characterization of the drug. Particle sizes of colloidal suspensions range from 1 nm to 0.5 μm [1]. Suspensions Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00003-2 © 2020 Elsevier Ltd. All rights reserved.

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improve drug stability, are safe and appropriate for administering to children and infants, and they can incorporate large dosages of drugs. In general, suspensions are used for the following applications in drug delivery: 1. For formulating insoluble or partially soluble drugs in a liquid form that is appropriate, safe, and convenient for elderly patients, children, and infants. 2. To mask the unpleasant taste of drugs such as chloramphenicol palmitate and paracetamol. 3. To increase the stability of specific drugs. 4. For formulating X-ray contrast media such as rectal and oral propyliodone. 5. For formulating topical formulations such as calamine lotion which require the active agent to be left as a light deposit after the evaporation of the dispersing agent.

3.2.1.2 Pharmaceutical foams New foam-based carriers for active agents offer several advantages over traditional carriers. Foams, either liquid or solid, are colloids that contain two phases and are produced by supersaturation of the liquid phase with gas or via mechanical means. Foams are formed when a gas (dispersion phase) is dispersed in solid or liquid phase (continuous phase). Foams can be produced by several methods including the pressurized aerosol system, the bag-in-can system, in situ gas generation, whipping, shaking, bubbling, and air spray foam pump line [2]. After formation of foam, the liquid (continuous) phase changes into solid/gel phase and results in solid or dry foam. This form is also known as sponge or xerogel. Dry foams are used to deliver steroids, antibiotics, or disinfectants topically. Gelatine or collagen sponges with high absorbance of ichor are used after cosmetic surgery or after skin-grafting surgery. Solid foams such as polyurethane and foam rubber are important commercial products [3]. Foams are mechanically and thermodynamically unsteady systems (elastic system) that can shrink as the gas confined in the foam bubbles is compacted [4]. Liquid foams have widespread applications in the pharmaceutical industry, and patients find them more pleasant than traditional topical formulations as they do not leave a residue, are nonviscous and easy to remove [5].

3.2.1.3 Emulsions An emulsion is a dispersion system which can be used to deliver either oil or watersoluble drugs. This system allows for the formation of droplet drug carriers which result due to the shear stress and the interfacial tension between immiscible fluids. The flow rate, solution viscosity, and surface tension are key parameters that need to be optimized in order to obtain effective emulsion systems [6]. Several types of emulsions, including oil in water (o/w), water in oil (w/o), multiple emulsions [water in oil in water (w/o/w) and oil in water in oil (o/w/o)], micro- and nanoemulsions have been used in various applications and are discussed briefly in this section. In order to obtain stable dispersed particles, an ionic or nonionic surfactant is usually used to reduce interfacial tension between the different phases within an emulsion system. Nonionic surfactants are commonly used in pharmaceutical

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preparations as they are less toxic. Surfactant molecules help to form a hydrophilic layer around oil particles in o/w emulsions, and their hydrophobic tails protrude into the oil phase thereby helping to stabilize w/o emulsions. In order to produce highly stable emulsions, either a single surfactant or a mixture of different surfactants can be used. Higher stability has been reported for emulsions formed using a mixture of surfactants. In order to facilitate the selection of a suitable emulsifier for a given emulsion, a hydrophilic/lipophilic balance (HLB) system was established in 1948 by William C. Griffin [7]. This system can provide a guide to determine a value for a specific emulsion and then relate it to an emulsifier value on HLB system. For example, a lipophilic emulsifier may have an HLB number less than 9, and a hydrophilic one is recognized with a HLB value greater than 11 [8]. Since the invention of HLB system, many attempts have been made to improve calculations and to incorporate more emulsifiers within the HLB scale [9]. Several guides for selecting a proper emulsifier and related calculations have been published [8]. A shortcoming of the HLB system is the inaccuracy of estimated HLB values when other additives are present in emulsion. Moreover, the calculated HLB values do not take temperature effects into account. Various synthetic and natural oils such as paraffin oil, corn oil, and sunflower oil have been used in the preparation of emulsions. The type of oil used can affect the mean particle diameter [10] and the overall stability of the emulsion. Gart and Remon proved that the stability of emulsions is directly related to the degree of saturation of both the emulsifiers and the oils and that as the percentage of unsaturation in the oil increases, more stable emulsions are formed [11]. The type of oil used also influences the final cost of emulsion-based pharmaceutics. For pharmaceutical applications, natural oils are a better and safer option when compared to synthetic oils. Sunflower and sesame oil are excellent options for cosmetic emulsions as they contain a high percentage of vitamin E. O/w emulsions have been used to entrap lipophilic drugs such as prednisolone, hydrocortisone, betamethasone, testosterone and its esters, and progesterone [12,13]. Although w/o emulsion systems have also been used for drug delivery, their use is limited by the fact that their oral and parenteral delivery are associated with complications arising from their destabilization upon dilution in the biological aqueous phase. It is possible to use w/o emulsions for drug delivery routes such as intramuscular injection where the risk of dilution by the biological aqueous phase is low [14]. W/o emulsion systems can be explored for oral delivery of hydrophilic drugs such as peptides or oligonucleotides. These drugs can be dispersed in the aqueous phase to produce w/o droplets in order to protect them against enzymatic degradation [15]. Microemulsions, another type of colloid used to form drug delivery carriers [16
18], exhibit a better thermodynamic stability and optical clarity and are simpler to prepare when compared to normal emulsions [13]. Nanoemulsions can produce droplets as small as 25
50 nm. Several studies have reported the use of nanoemulsions for ocular drug delivery [19
22]. Multiple emulsions have also been used for the targeted delivery of drugs. A multiple emulsion is a colloidal system of a combination of o/w and w/o emulsions which is stabilized by lipophilic and hydrophilic surfactants [23]. In

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multiple emulsions, the particles of the dispersed phase are contained in another droplet within the continuous phase. This forms an additional compartment which helps to prolong the release of a drug and can be used to deliver multiple drugs simultaneously [24]. In order to acquire a stable multiple emulsion, the ratio of hydrophilic and lipophilic surfactants must be optimized. Furthermore, the stirring speed and time of the primary and secondary emulsions should be finely adjusted. A 70% prolongation in release of naltrexone hydrochloride has been obtained using w/o/w and o/w/o emulsions [25]. Several studies have utilized w/o/w emulsions for delivering hydrophilic bioactives such as immunoglobulins [26], insulin [27], proteins [28], and amino acids [29,30]. Furthermore, an encapsulation efficiency as high as 90% has been reported for w/o/w systems [31]. However, several disadvantages such as the potential for the leakage of encapsulated bioactive, the difficulty in large-scale production, and low cost-effectiveness and storage stability are associated with these emulsion systems.

3.2.2 Microfabrication and molding of pharmaceuticals Injection molding (IM) is a fast manufacturing technique that is widely used in the plastics industry for producing parts that require a defined dimension, size, and shape [32,33]. This technique uses a ram or a screw-like plunger under highpressure conditions to force melted material into a mold where they are shaped to the contour of the mold. The completed product solidifies inside the mold and is ejected at the end of the manufacturing cycle. IM machines generally consist of the plastinating injecting unit (PIU) and clamping unit (CU). PIU contains a hopper for feeding the material into the heating barrel, where the heating, mixing, compression, and melting procedures happen. The temperature can be controlled for different materials by the heater bands. CU is the mold part, where the melted material is injected and shaped, and generally consists of two halves that combine to form a cavity of the desired three-dimensional (3D) shape. The main benefits of IM in the pharmaceutical field are related to its scalability, versatility, and favorable processing conditions [32,34
36]. As IM is a continuous cyclic process, it can be easily scaled up to industrial levels with large throughput. Each single cycle of IM can be completed in a few seconds and this enables concurrent production of multiple parts, thereby reducing process time and increasing effectiveness for scalability. The versatility of IM brings pronounced benefits and possibilities for manufacturing pharmaceutical devices with defined shape and dimensional features [32
39]. The IM process does not require any solvents and therefore simplifies the manufacturing process, reduces cost and time, and preserves the stability of drugs [32,40]. Furthermore, the pressure and heat involved in the IM process can help to reduce the microbial contamination and promote the interactions between drug and polymer to improve the bioavailability of poor-soluble drugs [41,42].

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3.2.2.1 Materials used for injection molding Polymers, including thermoplastics, thermosets, and elastomers, are the basic materials used for IM [32]. Since 1995 there have been approximately 18,000 diverse materials available for IM and the total number of available materials has been steadily increasing [33]. Currently, the plastic industry supplies a wide range of polymers for IM, including alloys or blends of previously developed materials. For the selection of a suitable material for an IM application, the various characteristics that need to be taken in account include impact strength, melt viscosity, elasticity, density, mold shrinkage, flow temperature, tensile strength, elongation, modulus of elasticity, hardness, water absorption, luminous transmittance, brittleness temperature, and coefficient of linear thermal expansion [43]. Additives such as fillers, colorants, and antishrinkage agents can be used to improve the thermos behaviors and plastic properties of the materials [32].

3.2.2.2 Manufacturing parameters There are several critical parameters that may affect the quality of the finished product. Huang and Tai [44] enumerated the following input parameters for the surface quality of the IM products: mold temperature, melting temperature, packing pressure, packing time, and injection time [44]. In addition to these parameters, the effects of injection speed and injection acceleration on the width of the segregation line have also been shown to be important [45]. The effects of melting temperature, injection pressure, packing pressure, and packing time on the polymer shrinkage have also been previously investigated [46]. It is important that the specific parameters in the IM process be optimized based on the individual properties of the different polymers, equipment, and the final applications.

3.2.2.3 Application of injection molding technology IM technology has been applied in the pharmaceutical industry primarily to reduce the cost and time of manufacturing and/or improve the performance of existing drug devices [32]. Oral capsules, as an alternative system to gelatin dip-molded capsules, have been produced by IM using natural materials such as potato starch or gelatin [35,36,38,40,47
53]. Oral devices such as an oral multilayer device [54
59], oral capsular device [60
62], and oral magnetic depot capsular device [62] have been produced using the IM technique. Production of implant matrices made of polylactic acid, polyanhydride copolymer, PLC, and poly(lactic-co-glycolic acid) to deliver drugs such as vapreotide pamoate, gentamicin sulfate, fluconazole, praziquantel, and 5-fluorouracil has also employed the IM technique [34,37,42,63
66].

3.2.3 Microfluidic technologies for drug delivery In recent years, the development of novel drugs and new and improved methods of delivering them to the body has interested many researchers. Drug delivery allows the targeting of a drug in dosage form to a specific site in the body while offering

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advantages such as better drug stability, improved drug release profiles, and better patient compliance. A DDS can either release the drug immediately or can help to lengthen the drug release based on specific requirements [67]. Drugs can be transferred through different routes such as via the skin (dermal), via inhalation, via oral administration, and hypodermic injections [68]. Most drugs do not reach the destination zone where they can work locally, thus they usually fail to provide optimum clinical results. This occurs because the active pharmaceutical agents enter the body and are distributed throughout the body including healthy nontarget sites. A large amount of the drug is wasted due to this widespread distribution of the drugs within the body. The development of new drugs is time consuming, typically taking an average of 15 years, and is also an extremely costly process which is estimated to increase by 7.4% every year [69]. Conventional delivery methods such as intramuscular, intravenous, sublingual, oral, subcutaneous, and rectal deliveries are associated with problems such as inflammation, interaction with foods, and degradation by enzyme due to pH changes in the body, low permeability, low solubility, and toxicity [70]. Furthermore, hypodermic injections are associated with pain and anxiety [71]. New technologies, such as microfluidics, that can control drug release at specific sites by using site-specific proteins and antibodies are being explored for their utility in the field of drug delivery. Such an approach would improve patient comfort and would have fewer side effects [72]. Microfluidics or lab-on-a-chip (LOC) use only a microscopic amount of fluid that is operated with micro-scale structure within a device. LOC devices are created by microfabrication systems that incorporate various components. Microfluidic devices contain microcapillaries or microchannels where the fluid flows, with a filter/heat exchanger, valves, pumps, actuator, and mixers as their basic components [73]. Microfluidic devices can be used to deliver therapeutic compounds and analyze biological fluids. They are portable and cost-effective and provide well-controlled microenvironments [74]. Microfluidics have applications in cellular analysis, diseases and cancer detection, synthesis of drug carriers and particles, and gene and drug delivery [75]. Particle production in microfluidic systems is categorized into droplets/emulsions, supraparticles synthesized by assembly of colloids and photolithographybased methods. The most popular carrier synthesis is droplet-based microfluidics, which can produce extremely homogeneous and reproducible drug-loaded particles, microgels, microbubbles, and microcapsules. The size of the particles can be controlled by adjusting the surface tension, viscosity of the solution, and flow rate. Droplets produced using this system are large and can deliver high dose drug loading and provide a stable drug release for long periods of time. In the production of supraparticles by the assembly of colloids and microfluidics is used for changing, manipulating, or assembling of microparticles into superparticles or other complex structures. In photolithography, an irradiated mask with UV light within a photo polymerizable flowing through microchannel-based device is used [6].

3.2.4 Solid freeform fabrication of drug delivery systems Solid freeform fabrication (SFF) is the production of freeform solid objects directly from a computer model. SFF technology builds 3D objects using a layer-by-layer

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manufacturing method. A computer-aided design (CAD) software first produces a computer-generated design. This CAD data are then transferred to the SFF machine, which produce the physical structure [76]. Additionally, computerized tomography or magnetic resonance imaging medical scans can also be used to create a customized CAD model which can be printed to achieve the exact external shape required [77]. The several types of SFF-based fabrication techniques are discussed in detail below.

3.2.4.1 Three-dimensional printing Three-dimensional printing (3DP) uses inkjet printing technology to eject a binder from a jet head. 3DP is also a layer-by-layer process and the predefined CAD structure is printed on a first layer of powder via an inkjet printing head. When the piston chamber is lowered, it fills another layer of powder, and the process is repeated until every layer is printed. Additionally, the unbound powder acts as a support that can be removed after the structure is completed. After the binder dries in the powder bed, the finished structure is retrieved, and the unbound powder is removed [76,78]. 3DP is fast, affordable, easy to use, and offers the advantage of being able to use several biomaterials for fabrication when compared to other SFF technologies. However, structural design can be restricted by low positional accuracy and nozzle size of the inkjet printer [79]. Bioprinting technology can be used to perform computer-assisted printing of natural polymers and viable cells [80].

3.2.4.2 Stereolithography Stereolithography (SL) uses an ultraviolet (UV) laser beam to solidify successive thin layers of photo-sensitive polymers, built layer-by-layer. SL irradiates the surface of a UV-curable liquid photopolymer and hardens the photopolymer. A crosssectional structure with solidified lines is produced when scanning UV lasers are overlapped. The successive stacking of these cross-sectional structures creates the desired 3D printed structure. Microstereolithography (MSTL) is another form of SL that provides micronscale precision in SFF techniques. MSTL enables freeform fabrication of 3D printed structures at the micron level and customization of the inner structural design by using CAD/computer-aided manufacturing technology just as in SL. In MSTL, a laser beam is used to solidify minor areas of the photopolymer by using a focusing lens. This technology was introduced in the early 1990s and has been used in LOC, microactuator, and prototype fabrication [81,82]. SL fabricates structural designs through the absorption of photon energy by a photo initiator via a series of monomer chain reactions which are referred to as initiation, propagation, and termination. The properties of the photopolymerization reaction control the outcome of the final structure and its characteristics such as the yield strength, elastic modulus, and shape accuracy. Thus when using SL as a fabrication method, the use of biomaterials with good photopolymerization capabilities is vital. Currently, there is a focus on developing new photocurable biomaterials,

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as the existing biomaterials have limited properties and have hindered the widespread application of SL technology. Newer biodegradable materials have been developed for SL applications by modifying existing biomaterials in order to improve photopolymerization. These new materials include polypropylene fumarate-based materials, gelatin-based materials, and trimethylene carbonatebased materials [83
85].

3.2.4.3 Fused deposition modeling Fused deposition modeling (FDM) is a manufacturing technology that uses a moving nozzle to extrude fibers of polymeric material layer-by-layer in order to build a structural design. FDM is mainly used for mechanical system modeling, fabrication, and production. FDM printing also creates external support structures that must be manually removed [86,87]. FDM technology uses the melt extrusion method to construct a structural design from a thermoplastic polymer [88,89]. When a layer in the x
y plane is completed, the base platform (z axis) is lowered, and the procedure is repeated until the complete structure is produced [90]. In this process, the filament materials are heated and extruded the through a nozzle onto the base plate following a path as set by the data implemented in CAD [91]. It is possible to control pore size and porosity by changing the material deposition amount, the spacing between the material paths, and the height interval (z axis) [92].

3.2.4.4 Selective laser sintering The selective laser sintering (SLS) technology uses a high-power carbon dioxide laser emitting infrared radiation to selectively heat materials like plastics, ceramics, metals, or composite powders just beyond their melting point for the fabrication of 3D structures [93]. During an SLS operation, the laser fabricates the cross-section of the model to be built by sintering powder in a thin layer [94]. The interaction of the laser beam and powder increases the temperature to the glasstransition temperature thereby fusing neighboring powder particles while also bonding the new layer to those previously sintered. After each cross-section is scanned and solidified, the piston retracts to a new position to supply a new layer of powder using a mechanical roller and the process is repeated until the design is completed [95]. The powder that remains unaffected by the laser acts as a natural support embedded in the surrounding and remains in place until the structure is completed [95]. SLS can create intricate designs, such as anatomically shaped structures and offers the advantage of customisation of pore sizes, porosity, and topology. SLS can be used to manufacture parts from an extensive range of materials such as bioceramics and titanium. Furthermore, large numbers of structural designs can be fabricated within the powder bed, thereby allowing mass production. However, this method may be limited in the production of structures from biopolymers as the operating temperature is very high [96,97].

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3.2.4.5 Bioplotting Bioplotting is an innovative technique that has been developed for soft tissue applications at the Freiburg Materials Research Centre, Germany. The characteristic of this process is 3D dispersion of a viscous material into a liquid medium having a matching density. It involves a moving extruder head (x-, y-, and z-axis control) and uses compressed air to force out a paste-like plotting medium. The extruder head can be heated to the required temperature. The medium solidifies when it comes in contact with the substrate or the previous layer [98]. This technology provides great flexibility as structures can be created without the need for support and a variety of materials including melts, pastes, reactive resins, or hydrogels can be used.

3.2.4.6 Phase-change jet printing This system includes two inkjet print heads. Each of the print heads delivers different materials, one for constructing the model and the other for providing support for any unconnected features. The jet heads make molten microdroplets that are heated above the melting temperature of the material and deposit it in a drop-on-demand which solidifies on impact to form a bead [99]. The overlapping of adjacent beads forms a line and the overlapping of adjacent lines forms a layer. A horizontal rotary cutter arm flattens the top surface and is used to control the layer thickness. This process is repeated to build a layer-by-layer design until the entire structure is completed. The structure is then submerged in a solvent which dissolves only the support material, so as to leave the physical model in its desired shape [100]. SFF technologies have great potential as they offer a high degree of freedom for the design of structures with regard to parameters such as pore size, pore geometry, orientation, and interconnectivity.

3.3

Physiochemical characterization of drug delivery systems

Many of the pharmacological and therapeutic properties of conventional drugs can be enhanced by using novel DDSs which are designed to improve the pharmacokinetics and biodistribution of drugs and to function as drug reservoirs with customizable release behavior capability [101]. The physiochemical features of DDS include morphology, physical and chemical properties, stability, and biodegradation and can be customized according to the drug(s) being delivered, their therapeutic purposes, and their pharmaceutical applications.

3.3.1 Morphology 3.3.1.1 Size The size of DDS should be customized according to the required application. For small-scale DDSs, such as colloidal systems, nanoparticles, and microparticles, the

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main aim is to develop systems with optimized drug loading, release properties, and a long shelf life [102]. For large-scale DDSs, such as implants and scaffolds, the size and shape should be personalized based on the therapeutic applications.

3.3.1.2 Porosity The porous structure of DDS is crucial for achieving controlled-release profiles. Generally, a small-scale DDS with porous matrices can offer the following advantages: (1) an ordered pore network which enables finely controlled drug loading and release kinetics; (2) a high pore volume to enable incorporation of the desired amount of drug(s); and (3) a high surface area, to provide high potential for drug adsorption [103]. For large-scale DDS, such as hydrogels, the porous structure can be organized by different levels of crosslinking. Similarly, the release rate and period of associated drug in the hydrogel/polymer systems can be controlled by the pore sizes [104
106].

3.3.1.3 Microstructure In a polymer DDS, it is important to study the microstructure of the matrix in order to understand the mechanism of drug loading and controlled-release behaviors, so as to optimize the performance of the DDS. The microstructure of DDS can be visualized using scanning electron microscopy (SEM) and transmission electron microscopy (TEM) [102,107
116]. Both SEM and TEM use an electron beam to interact with the samples to create visual images of the topography and composition of the material. The difference between SEM and TEM is that the electron beam in SEM is focused on the surface of the sample while in TEM it is transmitted through the cross-section of the sample material.

3.3.2 Physical properties 3.3.2.1 Mechanical properties The mechanical properties of DDS include tensile, strain, and compressive strengths and are particularly important features of films and hydrogel systems. The mechanical strength of a hydrogel/polymer system can be modified by adjusting the crosslinking interactions. For example, introducing chemical functionalities to the polymer backbone or increasing physical crosslinking can result in improved mechanical strength, modification of the drug
polymer affinity, and slowing of the gel degradation [117].

3.3.2.2 X-ray powder diffraction X-ray powder diffraction technique is used to produce a unique pattern for determining crystal phase. The peaks from samples at particular diffraction angles indicate the crystalline or amorphous states of a material. Each drug or component also

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has their own specific crystallinity peaks which can be detected as evidence of their presence or successful incorporation into the polymer matrix [117,118].

3.3.2.3 Differential scanning calorimetry Differential scanning calorimetry is used to study phase transitions by detecting the free energy changes arising due to an alteration in the enthalpy (ΔH) or entropy (ΔS) of a system [102]. The enthalpy changes can result in both endothermic or exothermic signals depending on whether the transition is an energy-consuming process such as the melting of a solid, or an energy-releasing process such as recrystallization [102].

3.3.3 Chemical properties 3.3.3.1 Fourier-transform infrared spectroscopy Fourier-transform infrared spectroscopy (FTIR) is used to measure the infrared spectra of samples [119]. FTIR study of drug-loaded polymer systems can help to provide information regarding the chemical structure of each component. FTIR can be used to detect the presence of the drug(s) in the polymer matrix and to reveal any polymer
polymer or polymer
drug interactions, which result in the generation of new peaks or shifting of existing peaks.

3.3.4 Stability and biodegradability One of the critical reasons for developing DDS for a conventional (“free”) drug is to enhance its stability and control its degradation rate. The stability and biodegradation of a DDS can be tailored to match different systems and pharmaceutical applications. For example, in the polymeric micelles system, the drug stability, loading capacity, and the release patterns can be tailored by adjusting the chemical composition, molecular weight, and the architecture of the amphiphilic block copolymers [120]. Biodegradability and the degradation rate are also important for developing successful formulations [121
123]. Biodegradable materials are commonly used in DDS as they are nontoxic, and their self-degradation in vivo obviates the removal of the device. Additionally, the release behavior of the incorporated drug is also partly dependent on the degradation rate of the polymer matrix.

3.4

Simulating a physiological environment for in vitro drug delivery studies

Drugs can be administered via various routes, and the selection of a route depends on the drug characteristics, site of action, duration of action, condition of the patient, and environmental factors such as pH, presence of enzymes, and

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Figure 3.1 The various routes used for drug administration.

metabolization of the drug. The various routes that are used for administering drugs are shown in Fig. 3.1. Drug release is the process by which a loaded drug is discharged in the body through dissolution of the matrix, diffusion of the drug in a solution, or both with time [124]. Drug dissolution is defined as a dynamic process in which a material is transferred from a solid state to solution. This is a two-step process that involves the initial release of molecules from the surface into the surrounding dissolution medium. In the second step, the drug diffuses into the rest of the solvent via a concentration gradient [125]. Drug release studies play a significant role in the development of DDS. The effectiveness of a DDS is influenced by parameters such as particle size of the drug and its carrier, active ingredients and their solubility, desorption of the confined drug by degradation, erosion or combination of both processes, and environmental conditions of the release medium [126]. In vitro drug release is a discriminatory method used to simulate the biological environment and the variables that impact the release rate. In vitro testing of drug release can help predict bioperformance of the pharmaceutical substance under controlled conditions [127]. Such studies play a significant role in preformulation studies and drug selection, screening of excipients, supporting of scaleup and preapproval changes, quality control procedures, investigation of in vitro
in vivo correlations, and act as a surrogate for in vivo studies. If an in vitro study is appropriately designed, the release profile obtained can reveal fundamental information on the dosage form and its behavior, and details on the release mechanism and kinetics, thereby enabling a rational and scientific approach to drug development [128]. A variety of apparatuses for dissolution testing have been proposed and tested over the last few decades. The most common apparatus used for in vitro drug release studies and their recommended applications are listed in Table 3.1. Currently, most of the tests and the recommended apparatus along with other key specifications are available as standards in pharmacopeia and are routinely used in pharmaceutical analysis. These apparatuses are periodically checked using mechanical and chemical performance verification tests to ensure the collection of reliable and reproducible data. For dosage forms such as orally administered mucoadhesive buccal tablets, chewable tablets, and sublingual preparations, product performance tests are adapted from existing procedures and have been well characterized in the United

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States Pharmacopeia (USP). USP dissolution apparatuses are principally designed for oral and transdermal dosage forms and are not optimized for studying the release profiles of novel and complex pharmaceutical forms. Testing these new DDSs require customized apparatus or custom-made setups. The Franz cell and chewing apparatus have been specially designed for topical and semisolid dosage forms and chewing gums, respectively. When developing a drug release test, besides equipment setup, the selection and volume of the dissolution medium, membrane type, mechanical simulation such as agitation speed, temperature, pH, contact surface, sink condition, and assay analysis technique need to be carefully defined in order to mimic the physiological environment of the intended release site in the body. Phosphate buffer solution is the most common dissolution medium used for in vitro drug release testing, but it can also be customized based on the drug delivery route and proposed release site by the addition of enzymes, modification of pH, or addition of minerals. Also, the temperature of the in vitro drug release system must be adjusted to simulate the temperature of the release site in the body; for example, 37 C for oral, 34 C for ocular, and 32 C for nasal drug delivery are recommended. The systems for studying in vitro drug release can provide important information about the drug release behavior of current DDSs; however, the limitations associated with each of the approaches need to be considered. Further refinement and improvement of the current in vitro systems as well as the development of new systems for existing and novel drug delivery carriers is imperative to the advancement of the drug delivery field.

3.5

In vitro and in vivo toxicity studies

3.5.1 In vitro toxicity studies In vitro studies refer to the technique of performing an experimental process in a controlled environment outside of a living organism or cells. Carrying out tests in animals and humans is restricted by accessibility to test subjects, feasibility of testing procedures, and ethical concerns. Additionally, animal models may not provide enough data to predict the clinical efficacy of therapeutic drugs for certain human tissue types. This has led to the need for developing in vitro models for the study of disease mechanisms and drug efficacy. In vitro models are advantageous as they can be controlled and are less expensive and time consuming when compared to animal models. In vitro models also provide systematic, repetitive, and quantitative investigation of cell physiology in the field of therapeutic drug discovery and development. Such testing is often not feasible with animal-based models [129]. The various assays that are used for analyzing in vitro cytotoxicity and other cellular responses include the 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide, cell proliferation, adenosine triphosphate, neutral red, and enzyme-linked immunosorbent assay. The two types of systems generally used for these in vitro tests are the static well plate system and the multicompartmental perfused system. Mammalian cells are one of the most commonly used models in in vitro studies.

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Table 3.1 An overview of common apparatuses used in in vitro drug release studies. Apparatus

Application

United States Pharmacopeia (USP) apparatus 1 (Basket) This apparatus consists of a vessel, a motor, a metallic drive shaft and a cylindrical basket. This apparatus permits observation of the specimen and stirring element during the test. 1-, 2-, and 4-L vessels are available. The vessel sits on a stirring element, and the stirring speed can be regulated. Adjustment of temperature can be controlled by a heating jacket or a suitable heating device. USP apparatus 2 (Paddle) This apparatus is very similar to UPS 1, except that the paddle formed from a blade and a shaft is used as the stirring element. USP apparatus 3 (reciprocating cylinder) The assembly consists of a set of cylindrical, flat-bottomed glass vessels, inert fitting, and screens designed to fit the tops and bottoms of the reciprocating cylinders. USP apparatus 4 (flow through): This apparatus is made of a reservoir and a pump for the dissolution medium, a flow-through cell, and a water bath that maintains the temperature of medium at 37 C. USP apparatus 5 (paddle over disk): This version uses the paddle and vessel assembly from apparatus 2, with the addition of a disk designed to hold the specimen at the bottom of the vessel. USP apparatus 6 (rotating cylinder): Similar to USP1, except to replace the basket and shaft with a cylinder-stirring element. The dosage is placed in the basket at the beginning of test. USP apparatus 7 (reciprocating holder): This apparatus consists of a set of calibrated solution containers. A motor and drive assembly reciprocate the system vertically.

Conventional tablets, chewable tablets, controlled-release solid dosage

Orally disintegrated tablets, chewable tablets, controlled-release solid dosage, suspensions

Controlled-release solid dosage, chewable tablets

Extended-release dosage, poorly soluble active pharmaceutical ingredients, powder, granules, microparticles, implants

Transdermal and topical systems

Topical and transdermal dosage form

Transdermal and topical dosage form

(Continued)

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Table 3.1 (Continued) Apparatus

Application

Franz cells: This system comprises of two parts: sample holder containing 250 or 450 mg of sample and a reservoir of the diffusion cell. PhEur: Chewing apparatus A simple masticatory movement employed to simulate the chewing action on a piece of gum placed in a small chewing chamber containing a known volume of buffer solution at a given temperature.

Drug release from creams, ointments, and gels (topical and semisolid)

Chewing gums

Two-dimensional (2D) models are mostly used in biological research but the limitations associated with them has led to the development of 3D models. 3D models provide a cellular microenvironment that mimics the microenvironment present within the tissues. This is important for drug testing since environmental signals can have a profound effect on the properties, behaviors, and functions of cells which in turn can also affect cellular responses to drugs [130
132]. The microenvironment may send signals to a cell through soluble factors in the interstitial fluid, cell/cell or cell/extracellular matrix (ECM) adhesion, or mechanical forces. At the cellular level, environmental signals may cause changes in cell shape and motility, influence proliferation, differentiation, and apoptosis, and affect the morphogenesis of cellular structures [133,134]. 3D models closely mimic in vivo microenvironments, the control of concentration gradients of signaling molecules and therapeutic agents, composition and structure of ECM around the cells, as well as morphology and arrangement of individual cells [135]. 3D models can also be used to investigate morphogenesis of cellular structures and specially engineered cells [132] which can delay the transport of drugs and genes to target cells, thereby imitating the physiological barriers of in vivo drug delivery. 2D models do not incorporate these barriers and therefore do not represent a realistic model for drug delivery testing. Recently, new in vitro models produced by the engineering of 3D constructs of bone, liver, skin, and cardiac tissues, among others have been developed [132,136,137]. This technology has allowed the development of 3D in vitro models that mimic liver, breast, cardiac, muscle, bone, and corneal tissues in normal organs, and malignant tissues for investigating microenvironmental effects and spatial and temporal constraints on cells. These models are complementary to cell monolayer and animal models in drug screening as well as pharmacokinetic and pharmacodynamic studies [138]. Currently, 2D culture systems dominate current cell-based assays or biosensors even though they are associated with contact inhibition and a loss of native cell morphology and functionality. 3D cell models, on the other hand, provide a more

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realistic representation of human tissues, which is critical to many important cell functions such as morphogenesis, cell metabolism, gene expression, differentiation, and cell
cell interactions. Thus for understanding cytotoxicity and testing of drugs, the maintenance of cells in their 3D environment is crucial as it would improve predictions and reduce clinical trial failures. While designing 3D models is much more complicated than designing 2D models, studying cell- and tissue-based assays with 3D models are much more superior and are therefore the assays of choice for the high-throughput screening of drug cytotoxicity [139
141].

3.5.2 In vivo toxicity studies In vivo studies use a whole, living organism for experimentation and is carried out to observe the overall effects of an experiment on a living subject. In vivo data can be collected from either animal studies or human clinical trials. In vivo studies allow metabolic tests to be carried out to investigate drug pharmacokinetics in order to reveal how therapeutic drugs are absorbed, metabolized, and excreted by the living subject when introduced via the oral, intravenous, intraperitoneal, intramuscular, or transdermal route. In vivo animal studies also allow acute, subacute, and chronic toxicity to be tested. Acute toxicity is investigated by using an increasing dose of the therapeutic agent until the signs of toxicity become obvious. At present, European legislation insists that acute toxicity tests be carried out in two or more mammalian species covering at least two different routes of administration. In subacute toxicity, the animals are administered with the therapeutic drug for 4 to 6 weeks in doses below the level at which it can cause rapid poisoning in order to observe if there is any toxic therapeutic drug metabolite build up over time. Lastly, chronic toxicity tests are longer term and can take up to 2 years and must be carried out in at least one nonrodent species. In vivo studies using various disease models also enable the efficacy of therapeutic drugs to first be tested in animals. For efficacy testing in humans, the therapeutic drug is often administered in a double-blind controlled trial to determine the effect of the drug and the dose
response curve. Additionally, specific tests on reproductive function, embryonic toxicity, or carcinogenic potential can be carried out depending on the result of other studies and the type of therapeutic drug being tested [142
145].

3.6

In vivo performance evaluations

Drug absorption from the gastrointestinal (GI) tract is determined by various factors such as the physicochemical properties of the drug, characteristics of the formulation, and interplay with the underlying physiological properties of the GI tract. Accurate prediction of the in vivo biopharmaceutical performance of oral drug formulations is important for ensuring efficient drug development. Conventionally,

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in vitro assessment of oral drug formulations for quality control purposes are mainly focused on disintegration and dissolution testing and the importance of in vivo biopharmaceutical performance has largely been ignored. More recently, new and innovative technologies for the evaluation of drug products in a more biorelevant and mechanistic manner, have led to a better understanding of drug formulation behavior. However, predicting the in vivo biopharmaceutical performance of formulations that rely on complex intraluminal processes (e.g., solubilization, supersaturation, precipitation) remains extremely difficult. Concurrently, the need for overcoming low drug solubility or controlling drug release rates has led to the development of novel in vitro tools. Hence, the amalgamation of physiochemical measurements, in vitro and in vivo methods, and physiology-based pharmacokinetic modeling is expected to greatly improve the whole process of drug development [146,147]. Physiologically based pharmacokinetic (PBPK) models employ a concept that describes the concentration profile of a drug in various tissues as well as in the blood over time, based on the drug characteristics, site, and means of administration, and the physiological processes to which the drug is exposed. PBPK modeling takes into account the factors affecting the absorption, distribution, and elimination processes [148,149]. In PBPK modeling, parameters are determined from in vitro experiments and physiology by utilizing in silico predictions to predict in vivo data. A Swedish physiologist and biophysicist, Teorell, developed a five-compartment scheme that consisted of a drug depot, fluid volume, kidney elimination, and tissue inactivation in order to imitate the circulatory system [150]. PBPK modeling enables the plasma concentration
time profiles to be predicted from preclinical in vitro and in vivo data, thereby providing support at various stages of the drug development process. PBPK models have identified key issues in the development of new drugs; however, there are still many aspects that need to be optimized in order to maximize the utility of the PBPK models for predicting drug absorption, understanding conditions in the lower small intestine and colon, the influence of disease on GI physiology whilst taking into account reasons behind population unpredictability. Notably, there is a vital need to create more in vitro models for testing dosage form performance and to streamline data input into the PBPK models [151]. Currently, the first commercially available software, GastroPlus, developed a comprehensive description of the GI tract in the context of a PBPK model. The first version, released in 1998, used a mixing-tanks-in-series approach to describe the movement of drug from one region in the GI tract to the next, with simple estimations of dissolution based on aqueous solubility and absorption rate constants based on existing pharmacokinetic data. Even at this stage, a reading obtained on the absorption (uptake across the GI mucosa) or solubility/dissolution would be limiting to the drug’s bioavailability. This was already a very significant advance as it enabled to show that solubility and dissolution problems are much more cooperative to formulation solutions than permeability limitations [152]. Additionally, several other commercial PBPK models such as Simcyp and PK-Sim have advanced descriptions of the GI tract. Furthermore, there are software programs such as

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MATLAB and STELLA that are adapted to predict in vivo performance of oral formulations [153]. All these programs attempt to account for all processes relevant to the GI absorption of drugs, including release from the dosage form, decomposition/ complexation in the GI tract, the various mechanisms of drug uptake and efflux, and first pass metabolism in the gut wall or liver. These programs describe the interaction of the various factors in determining the rate and extent of drug absorption from the GI tract [149]. Currently, biopharmaceutical profiling at different stages of drug development is still a trial and error process, where quality control traditional dissolution methods and in vivo testing in animals are the primary instruments to guide drug development. It is expected that “innovative” incorporation of in vitro data from more appropriate in vitro models with the development of the GI physiology component of PBPK models, will lead to an improvement in the ability of PBPK models to successfully predict oral drug absorption. Such advancements in technology will strengthen their role in preclinical and clinical drug development, as well as regulatory applications [154,155].

3.7

Correlation between in vivo and in vitro studies

An important aim of developing pharmaceutical formulations is to understand the relation between in vitro and in vitro performance. One of the challenges is to correlate in vitro release behaviors of the formulations to in vivo performance. Hence, researchers have been motivated to explore the in vitro
in vivo correlation (IVIVC) to create reliable models between the in vitro dissolution data and in vivo pharmacokinetic parameters [156]. The term IVIVC has been widely employed in pharmaceutical and related areas, and two definitions have been proposed by the USP and FDA [156
158]. According to the USP, it is defined as the establishment of a rational relationship between a biological property, or a parameter derived from a biological property produced by a dosage form, and a physicochemical property or characteristic of the same dosage form [157]. The FDA defines IVIVC as a predictive mathematical model describing the relationship between an in vitro property of a dosage form and a relevant in vivo response. Generally, the in vitro property is the rate or extent of drug dissolution or release while the in vivo response is the plasma drug concentration or amount of drug absorbed [158].

3.7.1 Correlation levels There are five correlation levels defined by the FDA based on the ability of the correlation to reflect the complete plasma drug level time profile resulting from administration of the given dosage form [158]

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3.7.1.1 Level A correlation This level of correlation is the highest category of correlation which represents a point-to-point relationship between in vitro and in vivo release [158]. The Level A correlation is used to define a direct correlation that the measurement of in vitro dissolution is sufficient enough to determine the in vivo biopharmaceutical rate of the formulation [156]. Hence, in such a case, the in vitro study can act as surrogate for in vivo performance as it is predictive in a point-to-point level.

3.7.1.2 Level B correlation This level utilizes the principles of statistics moment analysis, in which the mean in vitro dissolution time (MDTvitro) of the product is compared to either the mean in vivo residence time or the mean in vivo dissolution time (MDTvivo) [156]. Level B correlation also utilizes all the in vitro dissolution data and in vivo plasma level data. However, it is not considered as a point-to-point correlation and therefore an in vitro study in Level B cannot represent the actual in vivo performance [156,158].

3.7.1.3 Level C correlation This level represents a single point correlation instead of all the in vitro and in vivo data and uses one dissolution time point compared to one mean pharmacokinetic parameter such as AUC, tmax, or Cmax [156]. Therefore it is the weakest level of correlation and is very limited in predicting in vivo drug performance. It is typically used to evaluate the correlation in the early stages of formulation development [156].

3.7.1.4 Multiple-level C correlation In this level, one or several pharmacokinetic parameters of interest (Cmax, AUC, or any other suitable parameter) correlate to the amount of drug dissolved at several time points of the dissolution profile [156,158]. Level C correlation should be based on at least three dissolution time points covering the early, middle, and end stages of the release period. Based on this, if a Level C correlation is reachable, then Level A correlation can also possibly be achieved [156].

3.7.1.5 Level D correlation This level is a rank order and is not considered useful in formulation correlation; however, it can contribute to and aid in formulation development [156,158
160].

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Further reading E. Sachlos, J. Czernuszka, Making tissue engineering scaffolds work. Review: the application of solid freeform fabrication technology to the production of tissue engineering scaffolds, Eur. Cell Mater. 5 (2003) 39
40.

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4

Mo´nica Cristina Garcı´a1,2 1 Unidad de Investigacio´n y Desarrollo en Tecnologı´a Farmace´utica (UNITEFA)-CONICETUNC, Co´rdoba, Argentina, 2Departamento de Ciencias Farmace´uticas, Facultad de Ciencias Quı´micas, Universidad Nacional de Co´rdoba, Ciudad Universitaria, Co´rdoba, Argentina

4.1

Introduction

Advanced micro- and nanomaterial therapeutic systems hold several advantages over conventional therapeutic systems. Targeted delivery of drugs to specific sites of action comprises a big challenge for researchers in the field of pharmaceutical sciences [1], who have been witness of the great progress in the development of improved drug delivery systems (DDSs), including lipid-based and polymer-based carriers (Fig. 4.1). DDSs attempt to improve the therapeutic effect of drugs mainly by controlling their release, which allows modifying the pharmacokinetics of drugs and diminishing their toxic and side effects. Ultimately, these carriers may improve the biodistribution and boost intracellular penetration, enhancing the efficiency of drugs [2]. Micro- and nanoparticles (NPs) are structures that can be used as delivery devices for therapeutically active agents and/or imaging agents, namely antitumoral drugs, proteins, vaccines, and biotechnology-based therapeutic drugs [3,4], in which the active compounds (drug molecules or biologically active materials) are dissolved, entrapped, or encapsulated, or to which the active compound is adsorbed or attached (Fig. 4.2) [2]. Within all the DDS, NPs present numerous advantages mainly associated with their unique characteristics namely small particle size, large surface area, and versatility to customize their surface. NPs may acquire different chemical, physical, and/ or biological properties and functions that are remarkably different compared to those observed in conventionally scaled counterpart materials [2]. These properties can be used to achieve specific purposes with the aim of improving human and animal health [5]. To better address the many issues in nanomaterial design and expedite progress, scientists have fundamentally changed their approach to the research. Particularly, in the last 10 years, there has been a shift in paradigm from researchers working independently on narrow research goals to collaborative teams that facilitate solving greater aims. The combination of scientists and researchers with expertise and knowledge in different areas such as chemistry, biology, materials, engineering, Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00004-4 © 2020 Elsevier Ltd. All rights reserved.

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Figure 4.1 Carriers based on polymers and lipids commonly used as drug delivery systems.

Figure 4.2 Schematic representation of micro-/nanoparticles carrying active compounds (A) within its matrix, (B) on its surface, and (C) in its internal cavity (particles with core-shell structures).

and clinical practice, and research on nanomaterials has been able to advance more rapidly in the past few years [1]. Nanomedicine is not only an extremely explored area of research but also has become clinical in the last few years since the remarkable development and production of nanosized DDSs. Since the mid-1990s, various nanomedicines have been

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approved by the US Food and Drug Administration (FDA) for specific clinical indications including, for example, the treatment of several types of cancer, lymphomatous meningitis, chronic kidney disease, analgesia, macular degeneration, ocular histoplasmosis, fungal and protozoal infections, respiratory distress syndrome, and menopausal therapy. The list includes approval for novel materials along with the use of approved materials for new clinical indications. It is important to stress that liposomal and polymeric NPs are the most frequently FDA-approved drugs until the present [6]. In this scenario, this chapter aims to cover the main aspects regarding lipidbased and polymer-based nanocarriers. Information related to their development and applications as therapeutic carriers for controlled drug release will be highlighted.

4.2

Micro versus nanoparticles: physicochemical properties for drug delivery

One of the most evident aspects when comparing microparticles and NPs is their different surface area and volume ratio [7,8]. For instance, it has been reported that there is a major probability of loose payload during preparation in the case of larger particles in comparison to smaller ones. In addition, drug efflux may be faster from smaller particles than from larger ones. Thus size affects almost every characteristic of particle function including its degradation, mechanical properties, clearance, uptake mechanisms, etc. [9,10]. In particular, for intravenously injected systems it is widely accepted that the size of particles is a key physical parameter that can be tuned to dramatically alter their biological function [11,12]. The size is an important factor in the design of DDS to increase the half-life in vivo and bioavailability after administration because it has a remarkable effect on particle distribution throughout the body [13,14]. Microparticles usually possess sizes from 1 to 5 μm, and because of its size they are typically removed by the reticuloendothelial system (RES), whereas larger microparticles are usually trapped in the capillary beds. In addition, particles of at least 10 μm and more were shown to produce embolization in the liver and lungs [15]. Always depending on the route of administration, particles with 500
1000 nm can be phagocytosed by macrophages and smaller particles can be endocytosed by phagocytic or nonphagocytic cells [9]. It has been well established that NPs greater than 200 nm are likely to be mechanically filtered in the spleen and ultimately removed by the cells of the phagocytic system; thus generally results in short circulation times in the bloodstream. Particles in the range size from 100 to 200 nm show longer circulation times as they are large enough to avoid liver uptake, but small enough to prevent leakage in the spleen [16,17]. It is important to consider that particles smaller than 100 nm leave the blood vessels through fenestrations in the endothelial lining [9,18]. NPs between 10 and 50 nm are small

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enough to penetrate the very small capillaries within the body tissues, being able to offer a more efficient distribution in certain tissues. However, these particles can be sequestered by the liver. Finally, NPs smaller than 5
6 nm are eliminated fast for extravasation and renal clearance and are consequently not interesting in therapy due to their low residence time in blood circulation [16,17]. Drug delivery literature usually highlights the advantages of NPs over microparticles such as relatively higher intracellular uptake [13,14]. For instance, internalization by targeted nonphagocytic cells could be desirable; thus, as can be seen different strategies can emerge from the size influence in the behavior of particles. It is always necessary to execute pertinent designs depending on the route of administration and the therapeutic goal [8,19]. Not only particle size is important in the design of DDS but also the particle shape, which is a parameter with impact on the performance of drug carriers. Even though the role of particle shape in drug delivery is still in discussion, there is enough evidence to establish its effect over DDS [9]. It has been reported that shape can have dramatic effects on targeting, circulation, internalization, immune cell association, and adhesion [20
22]. Nonspherical particles (disk, cylindrical, and hemispherical forms) are usually more efficient than spherical ones by avoiding phagocytic cell retention for the RES, because they can more easily prevent absorption of opsonins on the surface of the NPs. They may also have increased ability to flow through thin capillaries and to adhere to the walls of blood vessels [23]. There are several aspects that the researchers in the field of nanomedicine should take into consideration since NPs formed into nonspherical shapes might exhibit improved properties over similarly sized spherical nanostructures. Even when some advances in the effects of particle shape have been reported [8,24
27], this parameter opens the door for future studies of its influence in degradation, transport, targeting, internalization, drug release, etc. from DDS [9]. For instance, several years ago, it was reported that particle shapes affect the nanocarrier degradation and drug release properties [28]. Several DDS attempt to get zero-order drug release profiles and this has been achieved by using hemispherical particles that only allowed degradation on the face. Interestingly, for nonspherical particles that have areas with different thicknesses, the shape of the particle may change over time and this behavior results in unique degradation profiles depending on the surface area and diameter [29,30]. The influence of shape on biological interactions is not as clear as the degradation process; however, there is evidence that indicates an intimate relation between phagocytosis and particle shape and between orientation of nonspherical particles and cellular uptake [9]. In addition, different cellular behaviors, including cell adhesion or cell penetration, depend on mechanical processes that originate from the cellular microenvironment. Therefore for the enhanced permeability and retention effect, the penetration of particles is enhanced if they exhibit good flexibility because the elasticity of NPs of any shape increases the possibility of penetration between endothelial cells. Moreover, for active targeting the flexibility of the NPs is a key factor in the

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interaction of them with cellularreceptor-mediated endocytosis since the interaction points of NPs with the cell receptor increase [31]. On the other hand, surface chemistry mainly influences the interaction of the particles with the cells and tissues of the body. The chemical functionalization of particle surfaces aims to diminish recognition by the components of the endothelial system of the reticulum, also called opsonization process, and therefore to increase circulation time in the bloodstream. One of the most reported strategies have involved the surface modification of the particles by functionalization with polyethylene glycol (PEG) to reduce opsonization and adsorption of antibodies on the particle surface, avoiding or minimizing its elimination. It is well-known that PEGylation of NPs increases its blood circulation and biocompatibility [32,33]. Besides, surface charge or zeta potential is another crucial parameter that has to be considered since it is highly related to the NPs toxicity. In general terms, negatively charged NPs exhibit higher biocompatibility than cationic ones [34]. However, positively charged NPs show improved interaction with the negative charge of the cell surface, thus they are more internalized into the cells. The impact of this parameter on in vivo biocompatibility of NPs in relation to size, zeta potential, and hydrophobicity has been studied over 130 nanosystems [35]. As stated above, aspects related to the size, shape, mechanical properties, surface chemistry, and surface charge of the particles need to be considered for obtaining suitable carriers for biomedical applications since they are directly associated to the residence time of particles in bloodstream, their biodistribution as well as the cell internalization for different mechanisms [36].

4.3

Types of carriers for drug delivery

Advancement in nanotechnology has a considerable impact on several industrial sectors, including pharmacotherapy and pharmaceutical technology, together with materials technology and biotechnology. Within this arena NPs in the role of nanocarrier are steadily being applied in biomedical applications, for example, the delivery of macromolecules as therapeutic agents (drugs and genes). The functions of carriers for drug delivery and stimuli exploited for triggering the smart NPscontrolled release of drugs are presented in Fig. 4.3.

4.3.1 Carriers based on lipids 4.3.1.1 Liposomes In 1964, Alec Bangham described that phospholipids in aqueous systems can aggregate in closed bilayered structures called liposomes [37]. Since then, these nanostructures have risen as one of the most useful tools for drug delivery in the field of pharmaceutical science [38]. Liposomes are phospholipid vesicles composed of single or multiple concentric lipid bilayers enclosing aqueous spaces [39] (Fig. 4.4). Most commonly used

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Figure 4.3 Schematic illustration for stimuli-responsive drug delivery systems and summary of different types of internal and external stimuli for triggered delivery of drugs from them.

Figure 4.4 Schematic representation of the different types of liposomal drug delivery systems. Source: Reproduced with permission from L. Sercombe, T. Veerati, F. Moheimani, S.Y. Wu, A.K. Sood, S. Hua, Advances and challenges of liposome assisted drug delivery, Front. Pharmacol. 6 (2015) 286. Copyright 2015 Frontiers Media S.A.

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components for their preparation are phosphatidylcholine and cholesterol [40]. As DDS, liposomes offer several advantages including biocompatibility, capacity for self-assembly, ability to carry large drug payloads, and a wide range of physicochemical and biophysical properties that can be modified to control their biological characteristics [38]. Liposomes can encapsulate small biomolecules, such as glucose, as well as larger ones, including peptides or proteins. Hydrophilic molecules are encapsulated in the internal aqueous compartment of liposomes while hydrophobic molecules are composed of bilayers [40]. These vesicles can be understood as closed lamellar bilayers. Hence critical geometrical requirements are necessary for their self-assembly. The shape and amphiphilic profile of phospholipids or surfactants are crucial parameters to form vesicles in aqueous media, according to their critical packing parameter. The smaller the value of this parameter, that is, the larger the head group area in relation to the surfactant volume, higher is the curvature of the aggregate in aqueous media [41]. Liposomes may have one or more bilayer membranes. Thus they are classified in large unilamellar vesicles and small unilamellar vesicles or multilamellar vesicles. Unilamellar liposomes are formed by a single phospholipid bilayer surrounding the aqueous media. Multilamellar liposomes are formed by concentric bilayers separated by narrow aqueous channels. Both size and number of vesicle layers can affect the drug-loading capacity and also the half-life of these nanostructures in the bloodstream or the permeation profile through the skin [42]. As DDS, the desirable size of liposomes ranges from 50 to 200 nm. However, the decrease of size and lamellarity is typically obtained by subjecting them to sonication, extrusion, or filtration methods [43]. In the last few years, another classification of liposomes has been considered, which include three different types. The first group includes regular liposomes without any modification and mainly consists of phospholipids and cholesterol. They are called large unilamellar liposomes and their sizes are in the range of 50
450 nm. These liposomes are composed of neutral, cationic, and/or anionic phospholipids. These vesicles were first used in the 1980s for the delivery of drugs such as doxorubicin and amphotericin to reduce clinical toxicity. However, the rapid elimination from the bloodstream limits their therapeutic efficacy. The second group includes modified liposomes by coating their surface with inert polymeric molecules, such as PEG, oligosaccharides, polysaccharides, glycoproteins, and synthetic polymers. Insertion of this hydrophilic component allows this “stealth” liposomes that are sterically stabilized with polymers to circulate for days, while non-PEGylated liposomes are cleared within hours [44]. The third group is an alternative targeting strategy developed for liposomes that involves the attachments of antibodies or peptides to either the surface of liposomes or the terminus of PEG molecules, enhancing the efficacy of drugs used to treat several diseases [42,45]. It should be pointed out that both the interior cavity and the surface can carry drugs and/or contrast agents. Various methods have been developed and optimized to prepare liposomes. Authors classify the conventional preparation techniques into [1] lipid film hydration method, which consists in hydrating, with an aqueous solution, a previously

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dried lipid film [46]; [2] reverse phase evaporation method, which consists in evaporating the organic phase of initially prepared reversed micelles; thus the reversed micelles are converted into a gel that collapses resulting in liposomes; and [3] solvent injection technique, which refers to the injection of phospholipids dissolved in an organic phase into an aqueous phase. In addition, new methods using sonication or homogenization were being developed to enable the preparation of large unilamellar vesicles with improvements in trapping efficiency with diameters less than 50 nm. Microfluidic mixing techniques now allow scalable production of large unilamellar vesicles in the size range of 20
50 nm, and supercritical reverse phase evaporation has been successfully used to form monodisperse liposomes. Moreover, all the processes can be optimized by the incorporation of an extrusion step through polycarbonate filters with pore sizes of 100 nm to size control of liposomes in aqueous media [47
50].

4.3.1.2 Solid lipid nanoparticles The solid lipid nanoparticles (SLNs) are one of the newest members of the lipidbased nanocarrier family and they made their first appearance almost thirty years ago [51]. The SLNs are nanostructures prepared from a lipid matrix (Fig. 4.5) that is solid at body and room temperature. SLNs are stabilized by suitable surfactants and their mean diameter is in the range of 40
1000 nm [52,53]. A great number of

Figure 4.5 Schematic representation of a solid lipid nanoparticle (SLN). Only one phospholipid layer is observed because the interior of the particle is solid. Some molecules such as antibodies, targeting peptides, and drugs can be bound to the surface of the SLN. Also, hydrophobic drug molecules can be loaded into the solid lipid.

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reports have described different SLN formulations since the early 1990s [52], and since then these nanostructures have received significant attention in the field of pharmaceutical technology research [53]. SLNs are interesting lipid-based carriers used as DDS because they can provide physical stability, protection of the incorporated drug from degradation, controlled release, and low cytotoxicity, if well-tolerated excipients are used [54]. Moreover, they preserve particle size on the nano- to submicron scale after drug incorporation; are composed of biocompatible and biodegradable components (i.e., physiological lipids or lipid molecules), do not require the use of organic solvents for their preparation, and the fabrication process can be performed at a lower cost and is easily scaled up [52]. SLNs are composed of 0.1%
30% solid lipid matrix, including one or more of the base ingredients, namely trimyristin, tristearin, trilaurin, stearic acid, and glyceryl caprate as Capmul MCM C10, theobroma oil, triglyceride coconut oil, 1-octadecanol, glycerol behenate as Compritol 888 ATO, glycerol palmitostearate as Precirol ATO 5, cetyl palmitate wax, among others, which are dispersed in an aqueous media of 0.5%
5% surfactant as a stabilizing agent, and 5% of the drug to be incorporated [52,54]. For longer circulation time in vivo, curdlan and PEG molecules have been used to functionalize the surface of the SLN [52]. The methods to incorporate drugs into SLN can be performed using highpressure homogenization, microemulsion formation, emulsification and solvent evaporation (precipitation), solvent injection (or solvent displacement), phase inversion, multiple emulsion technique, ultrasonication, and the membrane contractor technique [52]. One of the most employed methods for obtaining SLN is the high-pressure homogenization, which can be performed as hot or cold technique. In both cases, the drug is dissolved or solubilized in the lipid being melted at approximately 5 C
10 C above its melting point. For the hot-homogenization technique the drugcontaining melt is dispersed under stirring in a hot aqueous surfactant solution of identical temperature. Then, the obtained preemulsion is homogenized using a piston-gap homogenizer, and the produced hot oil-in-water emulsion is cooled down to room temperature allowing the lipid recrystallizes and leads to SLN. This methodology is mainly used for drugs that show some temperature sensitivity, because the exposure to an increased temperature is relatively short [55,56]. In contrast, for highly temperature-sensitive drug molecules the cold homogenization technique can be used, in which the drug-containing lipid melt is cooled, allowing the solid lipid ground to lipid microparticles which are dispersed in a cold surfactant solution yielding a presuspension. Then this presuspension is homogenized at or below room temperature, and the cavitation forces are strong enough to break the lipid microparticles directly to SLN. This process avoids or minimizes the melting of the lipid, diminishing loss of hydrophilic drugs to the water phase [56]. Some factors may influence the drug-loading capacity in the lipid matrix such as drug solubility in melted lipid, miscibility of drug melt and lipid melt, chemical and physical structure of solid lipid matrix, and polymorphic state of lipid material. To obtain a high-loading capacity a sufficiently high solubility of the drug in

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the lipid melt is required. Generally, the solubility should be higher than required, because it decreases when cooling down the melt and might even be lower in the solid lipid [56].

4.3.1.3 Lipid nanocapsules The recently developed lipid nanocapsules (LNCs) have been reported in numerous literatures as promising DDS for carrying hydrophobic drugs due to their characteristic oily core (Fig. 4.6) [57]. LNC formulation is based on at least three main components: an aqueous phase, an oily phase, and a nonionic surfactant. The proportions of components vary according to the study and each component has different influences on LNC preparation and stability [58]. In general, LNCs present an oily core corresponding to medium-chain triglycerides surrounded by tensioactives providing a cohesive membrane made of lecithin and free PEG 660 and PEG hydroxystearate at high density leading to really weak complement activation and low macrophage uptake. The aqueous phase consists of a solution of sodium chloride salt in MilliQ water. These nanostructures have a lipoprotein-like structure which could also be considered as a hybrid between polymeric nanocapsules and liposomes. As compared with liposomes which are leaky and unstable in biological fluids, LNCs present physical stability up to 18 months [58]. One of the promising features of LNC as carriers is their easy and organic solvent-free preparation technique that can be easy to scaleup for future industrial purpose [57,58]. Their formulation is based on the phase-inversion temperature phenomenon of an emulsion leading to LNC formation with good monodispersion. This method has already been well described [58,59]. In short, all necessary

Figure 4.6 Schematic representation of a lipid nanocapsule, in which the most commonly components used for its preparation are detailed.

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components are mixed under magnetic stirring at room temperature leading to an oil-in-water emulsion. After progressive heating at 4oC/minute rate, a short interval of transparency at temperatures close to 75oC can %be observed, and the inverted phase (water droplets in oil) is obtained at 85% oC. At least three temperature cycles alternating from 60oC to 85oC at the same rate% are applied near the phase-inversion % % zone [60]. An irreversible shock induced by sudden dilution with cold water to the mixture at 72oC
75oC conduce to the LNC formation and a slow magnetic stirring % is then applied% for 5 minutes in order to stabilize the suspension. For the loaded formulation, the drug powder is first solubilized in the lipid under magnetic stirring and heated if necessary. A cosolvent can be used to improve the drug solubility in the oil phase [58].

4.3.2 Carriers based on polymers 4.3.2.1 Nanospheres and nanocapsules Nanospheres or matrix-type nanodevices (Fig. 4.7A) are polymeric NPs where the entire mass is solid and consists of spherical polymeric matrices which have been widely studied as carriers of therapeutic molecules. The biological active agents are commonly distributed evenly throughout the polymeric core and the active compound loaded is released into the environment by diffusion process [61]. Besides, drugs can be adsorbed on the surface of nanospheres [62]. These nanostructures present sustained and controllable drug-release profiles and high-loading capacity for poorly water-soluble drugs [63,64] as they have the ability to incorporate hydrophobic drugs at concentrations greater than their intrinsic water solubility [65]. The time taken to release the drug from nanospheres depends on the composition of the polymer matrix and its ability to uptake fluid [61]. Drug-loading efficiency and nanosphere size are critical factors that influence drug release from these types of nanostructures. In general, at higher drug loading, an important burst effect and faster release rate are observed at the first time in release studies. Moreover, larger nanospheres commonly have a smaller initial burst effect and longer sustained

Figure 4.7 Schematic representation of (A) a polymeric nanosphere and (B) polymeric nanocapsule.

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release than smaller ones [65]. Different mechanisms are involved in drug release, including erosion of polymer matrix or swelling by hydration, cleavage of polymer bounds by enzymatic degradation or hydrolysis, and diffusion of the physically entrapped drug [66,67]. Nanospheres can be prepared by different methodologies [68], such as emulsification-evaporation [62,69] and nanoprecipitation [65,70,71]. The first method can be performed by single or double emulsification. The single emulsification method consists of preparation of an oil-in-water emulsion which contains the drug and polymer in the oil phase and a surfactant in the aqueous phase. This method is appropriate for encapsulation of hydrophobic drugs. In the double emulsification method, a (water-in-oil)-in-water emulsion is obtained with the drug and surfactant in the internal aqueous phase, the polymer in the oil phase, and another surfactant in the external aqueous phase. In general, this method is used for encapsulation of hydrophilic drugs [62]. In both cases, after the emulsification process, solvent evaporation is required for obtaining the nanospheres. Regarding the nanoprecipitation (or solvent displacement) method, it involves polymer and drug solubilization in a solvent followed by addition of this solution to a nonsolvent solution under constant stirring. Then, the solvent is evaporated. After that, purification and recovery of nanospheres usually require ultracentrifugation or freeze-drying [62,65,66]. On the other hand, nanocapsules or reservoir-type nanodevices (Fig. 4.7B) are vesicular systems that consist of a liquid core (water or oil) in which a drug can be loaded surrounded by a polymeric membrane or coating [72]. These polymeric NPs are hollow and can be defined as nanovesicular systems that exhibit a typical coreshell structure in which the drug is confined to a reservoir [65,72], and usually they present average sizes of 100
500 nm [62]. The core can be filled with drugs, vaccines, or genes. The liquid cavity can contain the biologically active substances in liquid or solid form or as a molecular dispersion [72]. The polymer membrane can also entrap some drugs and it can be functionalized to reach drug-controlled release and targeted delivery [73]. Nanocapsules can be prepared by different methodologies, including nanoprecipitation, emulsion-diffusion, double-emulsification, emulsion-coacervation, polymercoating, layer-by-layer, and emulsion
evaporation [72,74]. Nanoprecipitation, also called solvent displacement or interfacial deposition, is the most employed method for obtaining polymeric nanocapsules loaded with bioactive agents. The initial nanocapsule dispersions obtained can be contaminated by solvents, salts, stabilizers, and cross-linking agents that must be removed in order to guarantee the purity required for biomedical applications. Therefore several strategies for purification have been widely studied, including dialysis against water or polymer solution, filtration through 0.45 μm, cross-flow microfiltration and diafiltration, to eliminate surfactants and solvents [72].

4.3.2.2 Dendrimers Dendrimers are a family of synthetic polymers with three-dimensional, highly branched molecular structures that show a high degree of monodispersity and a

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Figure 4.8 Schematic representation of a dendrimer, showing the number of generations (G). Source: Reproduced with permission from D. Huang, D. Wu, Biodegradable dendrimers for drug delivery, Mater. Sci. Eng. C.90 (2018) 713
727. Copyright 2015 Elsevier.

well-defined architecture [61,75] (Fig. 4.8). They consist of macromolecular systems composed and are synthesized beginning with a inner core and adding branching and functional groups to the outside. Every time the synthetic process is repeated, larger dendrimers are created, and this is called a “generation” and described as Generation 1 (G1), Generation 2 (G2), Generation 3 (G3), etc. [76]. They are stable and have surfaces that can be functionalized with ligands or molecules for active targeting [61]. Dendrimers were discovered in the early 1980s, and they have unique characteristics including a branched, layered architecture which confers them a globular structure and internal hollows that enhance sequestration [77
79]. These nanostructures can be defined as ubiquitous type of precisely defined polymers able to be used in different applications [77,79,80]. Because of all their special properties, dendrimers are considered as promising devices for biomedical applications and especially as nanocarriers for drug delivery purposes [80
85]. In fact, dendrimers have been extensively studied as drug carriers performing either covalent or noncovalent association with the drug loaded. In both cases, multivalent

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sites play a crucial role in the drug
nanocarrier interaction [77,78,81]. Drugs may be physically entrapment in or of the dendritic structure depending on the respective size and ratio of the therapeutically active compound and the dendrimer. Thus drug molecules can be encapsulated in their multifunctional core and protected by the extensive branching. Furthermore, some drugs, for instance paclitaxel, can also be attached to the exterior of the dendrimer [61]. In general, dendrimers can be synthesized via a step-by-step iterative coupling method either in a divergent or convergent approach [75,86]. They grow outward from a multifunctional core molecule, which reacts with monomer molecules containing one reactive and two dormant groups for obtaining the first-generation dendrimer. After that, the new periphery of the dendrimer is activated by reactions with more monomers [87]. In the divergent method, the synthesis starts from the core of the dendrimer to which the branches are attached by adding building blocks in an exhaustive and stepwise way, thus the growth of a dendron originates from a core site. This method involves assembling monomeric modules in a radial, branchupon-branch motif [80]. In the convergent method, the synthesis starts from the exterior, beginning with the molecular structure that ultimately becomes the outermost branch of the final dendrimer. In this approach, the final generation number is predetermined, which requires the synthesis of branches of a variety of necessary sizes beforehand for each generation [80,87]. Recently new breakthrough approaches in dendrimer synthesis have been described, based on “lego” and “click” chemistry, by using divergent and convergent synthetic strategies, and over 100 compositionally different dendrimer families have been synthesized and over 1000 differentiated chemical modifications have been reported [80].

4.3.2.3 Nanogels Nanogels are a special representative of nanosized systems, consisting of colloidal hydrogel NPs that are in turn made of cross-linked polymer networks [88,89] (Fig. 4.9). In the three-dimensional networks through physical or chemical crosslinking, they can hold large amount of water without dissolving into the aqueous

Figure 4.9 Schematic representation of chemically (A) core-cross-linked and (B) shell-crosslinked nanogels, and (C) physically cross-linked nanogels.

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medium [90]. They combine beneficial functions of dendritic systems with those of hydrogels such as large encapsulation cavities and the capability of swelling, as well as responsiveness. These novel nanostructures not only fill the size gap but also present a functional link between common dendrimer or polymer scaffolds and macroscopic hydrogels [91]. Nanogels have attracted growing interest in the last five decades because of their potential applications in the field of pharmacy and medicine, including drug delivery and bioimaging, and due to their flexibility and biocompatibility [92
95]. They exhibit different characteristics such as size in the nanometer scale, porosity, amphiphilicity, charge, softness, and degradability that can be regulated by changing their chemical composition [90]. Nanogels are able to stably trap therapeutically active agents such as drugs, proteins, and genetic material inside the polymer networks. Internal and external factors of the microenvironment can stimulate the response of nanogels to control release of the payload [89,92]. They can be prepared by a number of techniques. Traditionally, nanogels have been classified based on the method of cross-linking as either chemically (covalently) (Fig. 4.9A and B) or physically cross-linked nanogels (Fig. 4.9C). Chemical cross-linking involves formation of covalent bonds between the polymer chains during polymerization of low molecular weight monomers or cross-linking of polymer precursors. The most extensively employed methods for obtaining chemically cross-linked nanogels use heterogeneous polymerization reactions in the presence of either bifunctional or multifunctional cross-linkers. Conventional and controlled/living radical polymerization methods allow for preparation of nanogels with different composition, dimensions, and architectures including core-shell and hollow nanogel particles. Moreover, a variety of other cross-linking approaches such as click chemistry, Schiff-base reactions, thiol-disulfide exchange, amide cross-linking, photo-induced cross-linking, etc., have been developed for the fabrication of nanogels from the polymer precursors. Physically cross-linked nanogels though formed under mild conditions, tend to be more fragile than their chemically cross-linked counterparts as they are stabilized by relatively weak interactions between polymer chains including hydrogen bonding, hydrophobic interactions, or ionic interactions. One of the challenges in the preparation of nanogels by such polymers is a control over the particle size, which requires fine-tuning of the polymer concentrations or environmental parameters, namely temperature, pH, and ionic strength. Generally, the advances in polymer chemistry led to the exceptional diversity and control over the composition, architecture, and functionality of cross-linked nanogels, which in turn provide more flexibility to tune their properties to comply with targeted biomedical applications [90]. Also, nanogels can be obtained in the presence of nucleation sites, often inorganic core, including iron oxide NPs or quantum dots, on which the monomers or polymers are adsorbed and polymerization arises. In this technique, nanogels are built on nucleation sites which work as “templates,” so this methodology allows obtaining nanogels with higher monodispersity [96].

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4.3.2.4 Polymeric micelles and polymersomes Both self-assembled nanostructures are composed of amphiphilic block copolymers in aqueous media [2]. The self-assembly procedure is favored by a thermodynamic process since the hydrophilic chains cover the hydrophobic core to avoid their direct contact with water, which confers a reduction of the interfacial free energy of the polymer
water system [97,98]. For self-assembled nanostructures based on block copolymers, not only the critical concentration for self-assembly is an important parameter but also a parameter known as hydrophilic volume fraction (f) has to be considered. f is defined as the relation between the hydrophilic portion of the polymeric chain and the total molecular mass. For copolymers with PEG as hydrophilic chain and taking into account the density of homopolymers, it is possible to predict the type of nanostructure aggregation. Therefore spherical polymeric micelles are favored at values of f . 50%, while vesicular structures or polymersomes are preferentially formed at 25 , f , 40% [99,100]. Polymeric micelles are self-assembled nanostructures than can be obtained when the polymer concentration is above the critical micelle concentration [101]. They consist of an external hydrophilic surface and a hydrophobic core (Fig. 4.10A). These nanostructures can successfully transport hydrophobic drugs and imaging agents in their hydrophobic cores and/or hydrophilic drugs coupled or adsorbed to the hydrophilic corona [102,103]. Furthermore, in case of core-inversible micelles (reverse polymeric micelles, Fig. 4.10B) hydrophilic molecules in a hydrophilic core can be sequestered [103]. Polymeric micelles exhibit higher stability compared with traditional surfactantbased micelles. This behavior can be explained considering that micelles based on amphiphilic block copolymers present a remarkably low critical aggregation

Figure 4.10 Schematic representation of nanostructures based on amphiphilic block copolymers: (A) typical polymeric micelle in a polar solvent, (B) reverse polymeric micelle in a nonpolar solvent, and (C) polymersome.

Nano- and microparticles as drug carriers

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concentration and slow kinetics of dissociation. Moreover, this kind of nanostructures present other advantages in comparison with other nanocarriers, which include high biocompatibility, ease to prepare and load with the drug, and small size (,100 nm) that allows a deep penetration into tissues [6,103]. Regarding cancer therapy, the small size allows them to participate in extravasation through the fenestrations in tumor vessels, limiting their uptake by the RES. Their hydrophilic corona also protects them from immediate recognition and subsequently increases circulation time in bloodstream [61]. Polymeric micelles can be obtained by chemical cross-linking and physical methods. For obtaining core cross-linked polymeric micelles by radical polymerization, bifunctional cross-linkers and disulfide bridges are commonly employed, depending on the characteristics of the micelle core. The chemical reactions allow obtaining polymeric micelles with improved circulation kinetics, biodistribution, and target-site accumulation; however, the chemical methods can be complicated [97]. On the contrary, physical methods are simpler and practical. They include direct dissolution, dialysis, oil-in-water emulsion, solvent evaporation, cosolvent evaporation and freeze-drying [65,97]. Direct dissolution is the simplest method for preparing drug-loaded polymeric micelles, in which the copolymers and the drug are mixed in water at or above the critical micelle concentration to self-assemble; however, this technique is associated with low drug-loading [97,104]. The most frequently used and reported method for encapsulation of poorly water-soluble drugs is dialysis, where the drug along with the block copolymer is dissolved in a watermiscible organic solvent, transferred to a dialysis tube and then dialyzed against water [65,97]. During this process, the organic solvent is replaced by water which induces self-association of copolymers and the encapsulation of drug. Nevertheless, this method is only suitable for lab-scale production [97]. An alternative is tangential flow filtration, which is a fast and simple method that can be used for scalable manufacturing processes of polymeric micelles [105]. In solvent evaporation method the drug and copolymer are dissolved in a volatile organic solvent and a thin film is formed after evaporation of solvent. Then, the film is reconstituted in an aqueous phase by vigorous shaking. This method is used if the hydrophilic
lipophilic balance of the block copolymers is high, which allows the reconstitution of the thin film in an aqueous medium [65,97]. In case of freeze-drying method a freeze-dryable organic solvent which can dissolve the copolymer and the drug must be used. After that, the solution is mixed with water, freeze-dried, and reconstituted with isotonic aqueous media. This technique is suitable for large scale production [97]. In contrast, polymersomes are artificial vesicles that have received significant attention due to their entangled architecture. These nanostructures are composed of amphiphilic block polymers and have spherical forms in which a hydrophilic core, that can encapsulate water-soluble molecules, is enclosed by a hydrophobic membrane, which may incorporate lipophilic molecules [106] (Fig. 4.10C). The membrane thickness of them can be tailored by shifting the hydrophobic ratio of the copolymers. Polymersomes are generally prepared using high-molecular-weight copolymers. Because of their macromolecular nature, the polymeric chains entangle

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between each other allowing an extra interaction between the copolymers. In consequence, their stability and toughness are superior that to their natural lipid counterpart, liposomes [106,107]. Polymersomes possess higher chemical and physical stability than liposomes. This is largely due to the higher molecular weight of copolymers in comparison to lipids, which differs by at least one order of magnitude. Furthermore, polymersomes are chemically versatile, since their physicochemical properties (vesicle size, membrane thickness, polarity, biodegradation, permeability, stimuli-responsiveness, and targeting capacity) can be modified by selecting polymer chemical composition, appropriate molecular weight values, and proportion between hydrophilic and hydrophobic blocks [108,109]. However, polymersomes usually show low encapsulation efficiency, which is limited to the concentration of the solution [110]. The main pathways for self-assembly of copolymers into polymersomes have based primarily on the value of f. However, only this parameter does not guarantee the success of self-assembly of copolymers into polymersomes and the method chosen to develop them is a key factor for their formation [111]. After the self-assembly process, hydrophilic blocks on the outside of the vesicle adopt a brush-like configuration, which increases their colloidal stability and camouflage ability against protein fouling [109]. The unique properties of polymersomes have attracted considerable attention in pharmaceutical nanotechnology. Because of their colloidal stability and tunable and resistant membrane properties, these nanostructures have been proposed as platforms for drug delivery, due to their ability to encapsulate a broad range of hydrophilic and hydrophobic agents, such as anticancer drugs, proteins, and genes [107,109,112]. More recently applications include imaging and theragnostics [113,114]. Regarding the use of polymersomes as imaging platforms, they provide higher resolution than conventional techniques and allow in vivo monitoring of biological pathways and cellular functions, besides being noninvasive [109,115]. On the other hand, polymersomes encapsulating simultaneously both therapeutic and diagnostic payloads, known as theragnostic, have also been studied [115]. Some examples of these applications will be presented at the end of this section. Polymersomes can be developed in different and very distinct sizes. Some methods that allow obtaining micrometer polymersomes are electroformation, lithography, microfluidic platform, and double emulsion [109]. On the other hand, the useful methods to obtain polymersomes in nanometer size range are film and bulk rehydration and solvent-switch. These ones generally need postpreparation to obtain an appropriate size distribution of vesicles, which includes sonication, extrusion through a polycarbonate membrane with defined pores and freeze
thaw cycles [106,109,116,117]. Most conventional methods commonly used for polymersome preparation are adapted from liposome preparation techniques. The film rehydration is a frequently applied method for lab-scale production. This method consists of a block copolymer solution in an organic solvent which is evaporated to a thin polymer film on a round-bottomed flask. The rehydration of the films is achieved by subsequently adding an isotonic aqueous medium, which leads to a detachment of the film from

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the glass surface. The swelling process can be influenced by stirring, shaking, or sonication, which affects, to some extent, the resulting vesicle size. Generally, this method allows obtaining unilamellar and multilamellar vesicles with a rather broad size distribution [106,118]. Some block copolymers also allow a direct dissolution from bulk material; however, a longer and much more vigorous agitation is required to obtain the complete rehydration of the polymer [119]. In the cosolvent method or also called solvent displacement or nanoprecipitation method, a amphiphilic polymer dissolved in a water-miscible solvent is added drop wise into water under vigorous stirring, and the solvent is subsequently removed by dialysis, freeze-drying, or evaporation [106]. For the development of polymersomes most methods still have low reproducibility and the feasibility for upscaling, which stays a big issue with regard to the translation into clinical application [106]. Current studies have focused on obtaining polymersomes of different shapes (tubular, toroidal, and higher genus particles). As well, recent variation refers to patchy polymersomes, prepared by mixing different types of copolymers. Moreover, polymersomes can also be designed to be responsive to environmental stimuli, including pH, temperature, and redox responsive nanocarriers [107,109]. Another rising research field is polymersomes as compartments for in situ reactions at the nanoscale (nanoreactors) [108,115].

4.4

Biomedical applications of lipid-based nanocarriers

Regarding the biomedical applications of liposomes, it is important to highlight that since they were discovered, several uses of them for different therapeutic purposes have been observed. In fact, some commercial products of liposome formulations have the approval of health regulatory agencies. Table 4.1 shows some of those approved formulations for clinical applications of liposomes in the treatment of cancer and infections, which are the most frequent applications of these approved nanocarriers for commercialization. AmBisome was the first liposomal formulation approved for clinical use. It is an injectable liposomal medicine of amphotericin B and it was approved for commercialization three decades after its discovery [43,120]. Doxil, in 1995, was the first clinically approved doxorubicin-loaded stealth liposomal for the treatment of different types of cancer [45]. More recently, in 2012, Marqibo was approved to treat acute lymphoblastic leukemia. This formulation is based on a liposomal medicine loaded with vincristine. Other examples of approved liposomal formulations can be observed in Table 4.1. Moreover, it is important to note that almost 40 liposomebased medicines are in different stages of clinical trials or have already been approved by the FDA [38,121,122]. Biomedical applications of liposomes can also be observed in different scientific reports, which demonstrate the interest of researchers in this type of nanocarriers. For example, regarding brain targeting, it has been developed a transferrin—cell penetrating peptide—sterically stabilized liposome which demonstrated the efficacy

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Table 4.1 Commercial products of approval liposome formulations. Commercial products

Drug loaded

Indications/therapeutic use

Year approved

AmBisome (Gilead)

Amphotericin B

Fungal infections Leishmaniasis

Doxil/Caelyx (Johnson & Johnson)

Doxorubicin

Kaposi’ssarcoma 1995 Ovarian cancer 1999 Breast cancer 2003 Multiple myeloma 1 velcade (Europe, Canada) 2007

1990 (Europe) 1997 (USA), 2000 1995

DaunoXome (Galen) Myocet (Cephalon) Amphotec (Intermune) Abelcet (Enzon) DepoCyt (Pacira) Lipo-Dox (Taiwan Liposome) Marqibo (Talon)

Daunorubicin

Kaposi’s sarcoma

Doxorubicin

Breast cancer 1 cyclophosphamide Invasive aspergillosis

Amphotericin B

1999 2003 2007 (Europe, Canada) 1996 (Europe) 1996 (USA) 2000 (Europe) 1996

Amphotericin B Cytosine Arabinoside Doxorubicin

Aspergillosis Lymphomatous meningitis Neoplastic meningitis Kaposi’s sarcoma, breast and ovarian cancer

1995 1999

Vincristine

Acute lymphoblastic leukemia

2012 (USA)

2001 (Taiwan)

Source: Adapted from M.C. Garcı´a, C. Aloisio, R. Onnainty, G. Ullio-Gamboa, Self assembled nanomaterials, in: R. Narayan (Ed.), Nanobiomaterials: Nanostructured Materials for Biomedical Applications, first ed., Woodhead Publishing, Cambridge, UK, 2017.

of penetration into the brain
blood barrier, long-term blood circulation in vivo and specific tumor targeting [123]. Also, a drug carrier system of apolipoprotein Emodified liposomes conjugated with phosphatidic acid was fabricated to improve brain
blood barrier penetration and release quercetin and rosmarinic acid to inhibit β-amyloid [1
42] (Aβ1
42)-induced Alzheimer’s disease for the effective treatment of this neurodegenerative disorder in the central nervous system (Fig. 4.11) [124]. In addition, there are several reports describing the relationship between liposomes, stem cells, and regenerative medicine. For instance, Rampichova´ et al. [125] demonstrated that liposomes loaded with fetal bovine serum improved both chondrocyte adhesion and proliferation on microfiber scaffold [123]. Monteiro et al. [126] showed that dexamethasone-loaded liposomes induce osteogenic differentiation of human bone marrow-derived mesenchymal stem cells [126]. Furthermore, for imaging purposes, liposomes enhance the concentrations of encapsulated contrast material, allow controlled release of the payload, and provide an

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Figure 4.11 Schematic representation of apolipoprotein E
quercetin
rosmarinic acid
phosphatidic acid liposomes intended for effective treatment of Alzheimer’s disease. Source: Reproduced with permission from Y.-C. Kuo, I.-Y. Chen, R. Rajesh, Use of functionalized liposomes loaded with antioxidants to permeate the blood
brain barrier and inhibit β-amyloid-induced neurodegeneration in the brain, J. Taiwan Instit. Chem. Eng. 87 (2018) 1
14. Copyright 2018 Elsevier.

adjustable coating for passive or active tumor targeting [127]. Moreover, in the development of vaccines, it has been reported a mucosally active subunit liposomal vaccine formulation, in which the authors explained that mucosal immunity induced by the active subunit liposomal vaccine formulation was superior to that induced by a peptide
protein conjugate administered with cholera toxin B. Those results were correlated with the stimulation of an inflammatory response by spleen cells from liposome-vaccinated mice [128]. On the other hand, concerning to the biomedical applications of SLN, there are several reports about their use as nanocarriers for drug delivery. For instance, SLN based on glyceryl monostearate and Lipoid S75 and loaded with acyclovir were obtained by high pressure hot-homogenization method and evaluated for skin drug delivery applications [129]. Also, SLNs as carrier of glibenclamide were developed

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through emulsion of solvent evaporation and hot high-shear homogenization techniques. They were composed by Precirol and Compritol as lipid components, and it was observed that Precirol-based SLN obtained with the emulsion technique showed improved properties, enhancing the oral bioavailability and therapeutic effectiveness of the drug, which demonstrated their potential use in the treatment of type 2 diabetes [130]. Stearic acid and injectable soya lecithin-based SLNs were obtained by the solvent displacement method and coloaded with vincristine and temozolomide for gliomatosis cerebri treatment [131]. The examples provided and others reported in the literature indicate that SLNs have mostly been developed as carriers of poorly water-soluble drugs to improve their performance as therapeutic agents. Regarding the biomedical applications of LNC, they have been developed for the targeted delivery of diagnostic and anticancer agents. For instance, the coadministration of paclitaxel and the immunostimulant CpG-loaded into LNC in the orthotopic GL261 glioma mouse model showed a significantly improved survival rate of mice in comparison to Taxol [132]. Also, a nanovectorized radiotherapy which consists of LNC loaded with a lipophilic complex of rhenium-188 (LNC188Re-SSS) has been reported. They showed a 83% of cure rate in the 9 L rat glioma model [133]. LNCs have been also reported for applications in gene and cell therapy. For example, the efficient gene silencing with LNC loaded with Bcl-2 siRNA and ferrocifens was verified by the specific extinction of Bcl-2 in melanoma cells [134]. In the area of active targeting, immunonanocapsules have been designed by conjugation of LNC to whole OX26 antibody, which can actively transport drugs to the brain parenchyma [135]. It was also reported that these standard LNCs were modified to allow the encapsulation of positively charged lipoplexes, which provide nanocarriers called DNA
LNC. They can efficiently protect the DNA in their lipid core [136]. The emerging application of albendazole-loaded LNC in the veterinary field is a promising alternative for the treatment of cystic echinococcosis [137], which confirms the versatility of this nanosystems and their wide field of biomedical applications.

4.5

Biomedical applications of polymer-based nanocarriers

Regarding the biomedical applications of polymeric NPs, several examples can be observed in the literature. In the next paragraphs, some of them will be presented and for further information the readers can take a look in the original sources. Table 4.2 presents a few examples of the polymeric nanocarriers intended for cancer therapy, detailing carrier used, therapeutic agent loaded, and stage of development. As can be seen in Table 4.2, different types of polymeric nanocarriers have been evaluated as carriers of anticancer drugs aimed to improve the treatment of cancer. Other examples of recently reported polymeric NPs include tamoxifen-loaded poly

Table 4.2 Some examples of polymeric nanocarriers, their main characteristics and stage of development. Polymeric nanocarriers

Carriers

Therapeutic agents loaded

Targeting ligand

Biomedical application

Stage of development

References

Nanospheres

PLGA and PEG
PLGA diblock copolymer

Docetaxel




Paclitaxel

PEG
PLA

Curcumin

TLTYTWS (TS) peptide


Preclinical (BALB/c mice) Preclinical (nude mice)

[138]

PEG
PLA

Cancer treatment. Intravenous administration. Melanoma tumor

In vitro

[140]

PLA

Tamoxifen




Breast and cervical cancer Breast cancer

[141]

PLGA

Short hairpin RNA




Preclinical (Wistar rats) Preclinical (nude mice)

PEG
PLGA

Cisplatin

RGD peptide

[143]

P(MEO2MA-coOEGMA-coDMAEMA)-b-PLGA PCLA
PEG
PCLA





PEG
PLA

Polydopamine (surface coat) Doxorubicin and paclitaxel (encapsulated) Superparamagnetic iron oxide nanoparticles, cyanine dye (IR820) and paclitaxel Gemcitabine

Preclinical (nude mice) Preclinical (BALB/c nude mice) Preclinical (SD rats)

[146]

PPMA
mPEG/PEG

Doxorubicin

RGD peptide

Cancer treatment

Preclinical (NOD
SCID mice) In vitro

PPI

Methotrexate

Folic acid

Breast cancer

Preclinical (Wistar rats)

[148]

Nanocapsules

Dendrimers




Gene therapy. ovarian carcinomatosis Breast and prostate cancer Photothermal, chemo, and gene therapy. Breast cancer Tumor targeting and multifunctional theranostics Colorectal and pancreatic cancer

[139]

[142]

[144]

[145]

[147]

(Continued)

Table 4.2 (Continued) Polymeric nanocarriers

Carriers

Therapeutic agents loaded

Targeting ligand

Biomedical application

Stage of development

References

Nanogels

Carborane bearing pullulan

Rhodamine B




Boron neutron capture therapy for fibrosarcoma

Preclinical (BALB/c)

[149]

PAA and PEG
PAA

Cisplatin and doxorubicin

Breast adenocarcinoma

Doxorubicin

Preclinical (BALB/c nude mice) In vitro

[150]

PNVCL-g-Ch

NH2-PEG and NH2PEG-folate


mPEG-P(LP-co-LC)

Doxorubicin




[152]

BSA-Ch PEG-g-PEI (L-histidine substituted)

Doxorubicin Methotrexate





Preclinical (C57BL/6 mice) In vitro In vitro

Dextran

Curcumin




In vitro

[155]

Au-PAm hybrids

Methotrexate

In vitro

[156]

Glycine-tethered PLGA

Methotrexate

Methotrexate or folic acid


[157]

PEG
PE/vitamin E PEG
pHPMAm

Paclitaxel and tariquidar Doxorubicin-prodrugs

Transferrin


PEG
PE

Pclitaxel and curcumin




Preclinical (Wistar rats) In vitro Preclinical (athymic nude mice) Preclinical (nude mice)

Polymeric micelles

Themochemotherapy of breast cancer Prostate cancer

Gastric cancer Hepatocellular carcinoma and breast adenocarcinoma Cervical cancer theranostics Papilloma

Mammary gland/ breast tumor Ovarian carcinoma Non-small-cell lung cancer Ovarian adenocarcinoma

[151]

[153] [154]

[158] [159]

[160]

Polymersomes

PEG
PE/vitamin E

Paclitaxel and curcumin

Transferrin

PEG-b-PCL

Paclitaxel, cyclopamine and gossypol




Ovarian adenocarcinoma Ovarian cancer

PEG
b
PGlu
b
PPhe

Cisplatin and paclitaxel




Ovarian cancer

PGA-b-PCL

Folic acid

PEEP
PEDP

Fe3O4 (within the coronas), doxorubicin (drug loaded) miR-200c (micro RNA)

PEG
PCL

Doxorubicin

Angiopep-2

Cervical cancer theranostics Paclitaxel resistance human lung cancer Brain tumor

PBC-b-PDMA

Eosin or camptothecin and doxorubicin (coencapsulated) Doxorubicin and paclitaxel




PMPC
PDPA







Photodynamic and combinational cancer therapy Head and neck squamous cell carcinoma

In vitro

[161]

Preclinical (athymic nude-Foxn1 nu mice) Preclinical (athymic (Ncr-nu/nu) mice) Preclinical (nude mice) Preclinical (nude mice)

[162]

Preclinical (Wistar rats) In vitro

[166]

In vitro

[168]

[163]

[164] [165]

[167]

PLGA, poly(lactide-co-glycolide); PEG, poly(ethylene glycol); PLA, poly(lactic acid); PCL, poly(ε-caprolactone); P(MEO2MA-co-OEGMA-co-DMAEMA)-b-PLGA, poly(2-(2-methoxyethoxy) ethyl methacrylate-co-oligo (ethylene glycol) methacrylate)-co-2-(dimethylamino) ethyl methacrylate-b-poly (D L-lactide-co-glycolide); PCLA
PEG
PCLA, poly(ε-caprolactone-co-lactide)b-poly(ethyleneglycol)-bpoly(ε-caprolactone-co-lactide); PPMA, poly(propargyl alcohol-4-mercaptobutyric acid); Mpeg, methoxy poly(ethylene glycol); PPI, polypropylene imine; PAA, poly(acrylic acid); PNVCL-g-Ch, poly(Nvinylcaprolactam)-g-chitosan; P(LP-co-LC), poly(L-phenylalanine-co-L-cystine); BSA, bovine serum albumin; Ch, chitosan; PEI, poly(ethyleneimine); Pam, polyacrylamide; PE, phosphatidylethanolamine; pHPMAm, poly[N-(2-hydroxypropyl) methacrylamide-lactate]; PGA, poly(L-glutamic acid); PGlu, poly(glutamic acid); PPhe, poly(phenylalanine); PEEP
PEDP, Poly[(mPEG)(N-ethyl-4-aminoazobenzene) phosphazene]-poly[(mPEG)(N,N-diisopropylethylenediamine)phosphazene]; PBC, Poly(benzyl carbamate); PDMA, poly(N,N-dimethylacrylamide); PMPC, poly2-(methacryloyloxy)ethyl phosphorylcholine; PDPA, poly 2-(diisopropylamino)ethyl methacrylate. Source: Adapted from M.C. Garcı´a, C. Aloisio, R. Onnainty, G. Ullio-Gamboa, Self assembled nanomaterials, in: R. Narayan (Ed.), Nanobiomaterials: Nanostructured Materials for Biomedical Applications, first ed., Woodhead Publishing, Cambridge, UK, 2017.

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(lactic acid) (PLA) nanospheres as potential nanocarrier for breast cancer therapy [141], nanospheres of poly(ε-caprolactone) (PCL) or poly(lactide-co-glycolide) (PLGA) loaded with phenazine lapazine for potential treatment of tuberculosis [169], and curcumin-loaded biodegradable PEG
PLA nanospheres for treatment of cervical and breast cancer [140]. Examples of nanocapsules include lychnopholide encapsulated in biodegradable polymeric nanocapsules for treatment Chagas disease [170], carvedilol-loaded mucoadhesive nanocapsules for sublingual administration in the treatment of cardiovascular diseases [171], PEGylated PLA nanocapsules loading high percentage of water-soluble drugs for cancer treatment [146]. On the other hand, dendrimers can be used in the field of drug delivery, targeting, diagnoses, such as contrast agents, in phototherapy, as well as in imaging [172]. It was reported the synthesis of amphiphilic linear-dendritic polymeric hybrids obtained for dual drug delivery of both doxorubicin and triptolide intended for breast cancer therapy [86]. Nanofibers of a fast dissolving polyamidoamine dendrimer as alternative to eye drops for topical antiglaucoma drug delivery of brimonidine tartrate was evaluated [173]. PEGylated polyamidoamine dendrimers for photosensitizer agent delivery in the treatment of fungal infections produced by Candida albicans were also studied [174]. Nanogels have been used for different therapeutic purposes. For instance, nanogels based on bovine serum albumin and chitosan to entrap doxorubicin as potential drug delivery system to treat gastric cancer [153]. Endosomal pH-activatable doxorubicin pro-drug nanogels were designed for triggered intracellular drug release in cancer cells [175]. Dual responsive nanogel (temperature and pH) based on poly(Nvinylcaprolactam)-g-chitosan loaded with doxorubicin rendering a combinational therapy of hyperthermia mediated drug delivery for the treatment of breast cancer [151]. Multifunctional nanogels through self-assembly of dextran
curcumin conjugates toward cancer theragnostics [155] and multifunctional self-assembled nanogels of curcumin
hyaluronic acid conjugates that inhibit amyloid β-protein fibrillogenesis and mitigate the amyloid cytotoxicity more efficiently than free curcumin [176] were also developed. Fluorescent nanogels were studied for disease prevention, testing, and treatment because of their fluorescence imaging and drugloading properties [177]. Cationic nanogels for intracellular delivery of proteins and nucleic acids were used to knockdown tumor growth factor expression with the immunological effect [94]. Different biomedical applications of block copolymer-based nanostructures have been reported. For instance, curcumin-loaded polymeric micelles were developed for enhancing cellular uptake and cytotoxicity to FMS-like tyrosine kinase-3 (FLT3) overexpressing eosinophilic (EoL-1) leukemic cells [178]. Polymeric micelles based on PEG
PLGA diblock copolymers containing curcumin acetate were investigated for different endocytic processes in respiratory epithelial cells [179]. Moreover, various formulations based on polymeric micelles are under clinical evaluation or used as commercial products. Some examples include CriPec (Cristal Therapeutics) based on polymeric micelles composed of a copolymer formed by a PEG block and a lactate-based hydrophobic block loaded with docetaxel, which

Nano- and microparticles as drug carriers

97

were studied for the treatment of several solid tumors (preclinical evaluation) [6]. NC-6004 (NanoplainTM) consists in cisplatin-loaded PEG
polyglutamic acid polymeric micelles for the treatment of solid tumors, which is under phase I of clinical trials [102,180]. BIND-014 (Bind Therapeutics Inc.), docetaxel-loaded PEG
PLGA polymeric micelles for the treatment of prostate cancer, is under phase II of clinical trials [6]. Paclical (Oasmia Pharmaceutical), which consists of paclitaxel-loaded poly(L-glutamic acid) polymeric micelles for ovarian cancer treatment, is under phase III of clinical trials [181
183]. As example of a commercial product, Genexol (South Corean company Samyang) is a formulation based on paclitaxel-loaded PEG
PLA micelles for the treatment of breast cancer, pancreatic cancer and small cell lung cancer [184,185]. Polymersomes are other type of nanostructures based on the self-assembly of block copolymers, had gained great attention in the last years. If they present optimal size, they can alter the tissue distribution of anticancer drugs. Moreover, the PEGylation of polymersomes or the development of PEG-containing polymersomes is widely used strategy to avoid the adsorption of blood protein components. The hydrophilic and neutral nature of PEG reduces protein opsonization, allowing the passive accumulation of polymersomes at the tumor site via the enhanced permeation and retention effect [107]. Some examples of reported polymersomes include rifampicin-loaded PEG
PCL polymersomes for pulmonary delivery to the treatment of tuberculosis [186]. Polymersomes encapsulating gadolinium were reported for magnetic resonance imaging [115]. In theragnostics, poly(trimethylene carbonate)
polyglycolic acid polymersomes loaded with doxorubicin and ultrasmall superparamagnetic iron oxide (contrast agent) were studied for diagnostic and treatment of localized tumors, providing contrast for magnetic resonance imaging and as magnetic field-responsive system for controlled doxorubicin delivery at the target site [113]. A special emphasis in the development of stimuli-responsive polymersomes has been observed in the last years and several examples can be read in the literature, which include responsiveness to biological-stimuli such as pH-responsive, redox-responsive, enzymeresponsive, glucose-responsive, and gas-responsive, as well as external stimuli such as temperature-responsive, light-responsive, magnetic field-responsive, and ultrasound-responsive [187]. As an example, reduction-responsive chimaeric polymersomes for highly efficient loading of pemetrexed (PEM) and targeted suppression of lung tumor in vivo were recentrly reported (Fig. 4.12) [188]. Lung cancer-specific CSNIDARAC (CC9) peptide-functionalized reduction-responsive chimaeric polymersomes (CC9-RCPs) loaded with PEM exhibited 22-fold longer circulation time and 9.1-fold higher accumulation in H460 tumor than clinical formulation Alimta which also contain PEM. Moreover, CC9-RCPs showed better tumor penetration than RCPs and PEM-CC9-RCPs effectively suppressed growth of H460 xenografts and significantly prolonged mouse survival time as compared to PEM-RCPs and Alimta controls. As stated above, numerous reports demonstrate the broad research in the development of lipid-based and polymer-based carrier for therapy and diagnoses. The

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Engineering Drug Delivery Systems

Figure 4.12 Schematic representation of the structure and functions of lung cancer specific CSNIDARAC (CC9) peptide-functionalized reduction-responsive chimaeric polymersomes (CC9-RCPs) for efficient encapsulation and targeted delivery of pemetrexed (PEM) to H460 cells in vitro and in vivo. Source: Reproduced with permission from W. Yang, L. Yang, Y. Xia, L. Cheng, J. Zhang, F. Meng, et al., Lung cancer specific and reduction-responsive chimaeric polymersomes for highly efficient loading of pemetrexed and targeted suppression of lung tumor in vivo, Acta Biomater. 70 (2018) 177
185. Copyright 2018 Elsevier.

crucial aim of research in nanocarriers for biomedical applications is to develop effective and safe treatments for clinic use as well as devices for diagnoses or theragnoses.

4.6

Final remarks and future perspectives

Urgent diseases such as cancer keep researchers looking for new alternative treatments that can achieve the therapeutic effect and leads to patient welfare. Even though great advancements in synthesis allow designing and developing nanocarreris with tailorable physicochemical, pharmacological, and biological properties, a lot of work remains. Several scientific and engineering issues as well as the scale up must be addressed for the successful translation of basic research to clinical applications. There is a need of deeper in vivo studies to better understand the complex interactions between the nanostructures and living organisms to

Nano- and microparticles as drug carriers

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comprehensively evaluate their performance as nanomedicines. Moreover, the biocompatibility and safety of nanocarriers should be studied. The field of stimuli-responsive nanocarriers is an area that requires more research and study to understand how to cross the biological barriers present when thinking in intravenous administration of them. Moreover, comparative studies against commercially approved products based on nanocarriers as well as against different types of nanostructure would provide better information regarding the improvements achieved when developing new nanocarriers for therapeutic uses.

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Ian Major1, Sarah Lastakchi2, Maurice Dalton1 and Christopher McConville2 1 Materials Research Institute, Athlone Institute of Technology, Athlone, Ireland, 2School of Clinical and Experimental Medicine, College of Medical and Dental Sciences, University of Birmingham, Birmingham, United Kingdom

5.1

Introduction

Implantable drug delivery systems provide extended release of a drug for the desired duration, usually over timespans of months and years. Placement of implantable devices routinely requires a specific insertion device for placement rather than a surgical procedure. A broad range of materials are currently in use for the fabrication of implants, both nondegradable and biodegradable, but since implants are exposed to tissues for prolonged periods, it is essential that each material is biocompatible to reduce undesirable cytotoxic effects. There are a wide variety of implantable devices in clinical use including subdermal implants, vaginal rings, intrauterine devices, ocular implants, and intracerebral implants. Subdermal implants are placed under the skin to deliver drugs to this large absorption site. Vaginal rings are flexible torus-shaped drug delivery devices that are inserted into the vagina for up to 12 months for the local and systemic delivery of a wide range of drugs. Ocular implants are anchored to the sclera of the eye or injected into the vitreous to maintain at an accurate constant drug dose for an extended period of time. Intracerebral implants encompass a range of clinical indications including chemotherapy and neuronal therapies. This chapter outlines how each implant listed above has been successfully manufactured using a variety of materials. The chapter is divided into two distinct halves—nondegradable and biodegradable materials— with an overview of how each type of material has been utilized in the production of experimental, patented, and/or marketed implantable drug delivery systems.

5.2

Nondegradable polymers

The three most important nondegradable polymers for the production of drug delivery implants are polyethylene vinyl acetate (PEVA), silicone elastomer (SE), and thermoplastic polyurethane (TPU) (Table 5.1). PEVA is a copolymer composed of both ethylene and vinyl acetate (VA) monomers polymerized via free radical Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00005-6 © 2020 Elsevier Ltd. All rights reserved.

Table 5.1 Summary of nondegradable polymers: their properties and functions. Polymer

Properties

Biomedical applications

Thermoplastic polyurethane

Crystallinity: 2.7%
18.3% Glass transition temperature: 244 C to 266 C Melt temperature: 170 C
220 C Density: 1.02
1.12 g/cm3 Tensile strength: 16
66 MPa Elongation: 300%
1500% Young’s modulus: 33
72 MPa Glass transition temperature: 235 C to 240 C Melt temperature: 170 C
220 C Density: 1.1
1.5 g/cm3 Tensile strength: 5.5
8 MPa Tear strength: 5
17 kN/m Elongation: 200%
800% Glass transition temperature 27 C Melt temperature: 72 C Density: 0.92
0.94 g/cm3 Elongation: 200%
900% Young’s modulus: 0.01
0.2 GPa Tensile strength: 7
40 MPa

Orthopedic implants, transdermal patches, catheters, scaffolds for bone and nerve regeneration, subdermal implants, intrauterine devices, vaginal rings

O

C

N

O

N

C

N

N

C

O

Silicone elastomer Si

Si O

Si O

n

Polyethylene vinyl acetate n

m

O O

C

O

Tear strength: 7
30 kN/m

Medical tubing, peristaltic pumps, catheters, heart pumps, ventricular assist devices, cannulas and vascular grafts, subdermal implants, intrauterine devices, vaginal rings

Subdermal implants, intrauterine devices, vaginal rings, biological encapsulation (human nerve growth factor hormone, albumin, monoclonal antibodies, DNA, heparin, and RNA)

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polymerization. The VA monomers are randomly distributed along the polyethylene back bone. Both VA content and distribution determine the final properties of PEVA including melting temperature, tackiness, the degree of crystallinity, and hydrophobicity. The degree of crystallinity effectively influences the mechanical performance of the polymer, including flexibility and hardness. In addition, the crystallinity of the material will determine drug permeability. Thus higher VA content PEVA has higher rates of drug release [1
4]. Two of the leading suppliers of PEVA resin, Celanese [5] and DuPont [6], were the first to patent and then exploit the copolymer for many industrial applications including hot-melt adhesives, cling film, shoe soles, and tubing. Celanese (United States) has become the leading supplier of pharmaceutical grade PEVA resin with their VitalDose grades being utilized in implantables and solid drug dosage forms. TPU is a linear block copolymer synthesized via polyaddition polymerization of diisocyanates and polyols [7
9]. The copolymer consists of hard segments (diisocyanates and chain extenders) and soft segments (polyols) that are immiscible and exist as distinct phases within the matrix. Hard segments are highly ordered and act as physical cross-links between neighboring amorphous soft segments. TPU containing a higher ratio of hard segments will be stronger and more robust, while a TPU with a higher soft segment ratio will be more elastomeric. Softer grades of the TPU are more amorphous and increasing the ordered hard segments leads to a more semicrystalline material. TPU’s bulk properties can be further altered by varying the type of diisocyanate, polyol, and chain extender present in each segment. Diisocyanates are either aromatic or aliphatic. Aromatic-based TPUs are more thermally stable with better mechanical properties but have poorer UV stability compared with the aliphatic-based TPUs. Polyols are long-chain diols, and molecular weight has a significant effect on melt-processability and final mechanical performance. Higher polyol molecular weight increases phase separation in the TPU, providing better elastomeric performance. TPU polyols can be polyether, polyester, polycaprolactone (PCL), or polycarbonate based, each providing different properties and stabilities. Polyester-based TPUs have excellent mechanical properties and have higher resistance to oils and chemicals but are susceptible to hydrolytic degradation. Polyether-based TPUs have low-temperature flexibility. Polyether-based TPUs are susceptible to oxidative attack and must be stabilized with comonomers. PCL-based TPUs have a mix of properties between polyether and polyester-based TPUs. Polycarbonate-based TPUs provide higher tensile modulus but lower elasticity. Some TPUs are cytotoxic while others can be safely implanted for up to five years [10
12]. TPU resins can also be engineered with biodegradable linkages for safe degradation in the body [13]. Polyester-based TPUs are hydrolytically unstable and will degrade in the body. Polycarbonate-based TPUs are the most suitable for long-term implantation as they do not undergo hydrolytic or oxidative degradation. TPUs have been used in the fabrication of medical tubing and other medical devices [14]. The polymer is being extensively investigated in drug delivery applications, and different resins offer various release profiles—zero-order, Fickian diffusion, or hybrid. Drug release is dependent on TPU crystallinity, steric

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hindrance, and hydrophilicity. Since crystalline regions hinder drug diffusion, softer TPUs will more likely display Fickian diffusion. The spatial structure of hard segments further controls drug diffusion as closer packing impedes drug molecule diffusion. More hydrophilic TPUs will permit the ingress of moisture and will swell the polymer, a phenomenon which can both hasten drug diffusion by increasing the free volume and can also slow drug diffusion by increasing the length of the diffusion pathway. DSM Biomedical (United States) and Lubrizol (United States) provide pharmaceutical grade TPUs in the marketplace. DSM Biomedical provides a variety of TPUs of different chemistries [15,16]. The Lubrizol company provides the Pathway resins as bespoke offerings to the pharma industry, and each resin is designed to match the exact requirements of the product. The TPU chemistry of these resins is much narrower than the DSM biomedical resins as each resin is hydrophobic or hydrophilic aliphatic polyether-based. Formulation of the hydrophilic resins can vary the water-swelling capacity to a broad range of 20%
1000% water by weight absorbed. Silicone has a polysiloxane backbone which is an alternating chain of silicon and oxygen atoms along with organic side groups (methyl, vinyl, or phenyl). SE is formed by cross-linking polysiloxane chains through addition or condensation reactions with a silicon-based oligomer cross-linker that joins chains to form a solid three-dimensional network. The physical properties of each SE are strongly dependent on polymer type, cross-linker type, and cross-link density [17]. The hardness of these elastomer systems increases with increasing cross-link density. Pharmaceutical grade SE is cured in the presence of a platinum catalyst, and a silicon hydride cross-linker via an addition hydrosilylation reaction in the presence of a platinum-based catalyst and the chain-to-chain cross-linking involves carbons on respective methyl side groups. Elevated temperatures ( . 100 C) are a requirement for the processing to occur at a suitable manufacturing rate. Other medical grade SE systems are available including condensation tin-catalyzed [18
20] and addition UV-curable [21] but have seen less uptake than the two-part addition platinum catalyst SE systems for a variety of reasons including the presence of byproducts, poorer mechanical performance, and the need for silica-based fillers. The building block of SE systems, the polymer polydimethylsiloxane, consists of two methyl side groups attached to the whole length of the polysiloxane backbone which is also trimethyl terminated. The methyl side groups shield the polysiloxane backbone giving a very hydrophobic character to the polymer with limited intermolecular interactions. These properties mean the polymer is both chemically and biologically stable and thus ideal for long-term implantation. Also, the polysiloxane backbone features strong polar bonds that are of low rigidity, and therefore permits both sterilization (resistance to thermoxidative attack) and drug diffusion (permeability). For these reasons, the material has found widespread use in drug delivery, medical device, and implant application [22,23,24]. Several chemical companies worldwide offer medical grade SE resins. Bluestar Silicones (France) offers Silbione resins suitable for drug delivery and long-term implantation. Dow Corning (United States) provides different silicone grades based on changes in polymer and copolymer chemistries. Wacker (Germany) resins are used for pharmaceutical and

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medical tubing. NuSil (United States) produces liquid silicone rubbers, high consistency rubbers, and low consistency elastomers for general drug delivery and long-term implantation. Shih-Etsu Silicones (Japan) supplies tubing and injection molding grade resins up to USP Class VI compliant and ISO 10993 standards.

5.3

Nondegradable subdermal implants

SE-based subdermal implants were the first such implants to be developed and have been primarily used for the controlled release of steroidal hormones. In the mid1960s, the Population Council was the first to develop a subdermal implant as longterm, reversible contraception [25]. The six-rod device placed in the upper arm was manufactured from 34 mm length of silicone tubing filled with 36 mg of crystalline levonorgestrel sealed at two ends with silicone adhesive [26]. The implant, which was marketed as Norplant, delivered a daily dose of 30 μg levonorgestrel for up to 60 months. A second generation two-rod implant replaced the crystalline levonorgestrel core with a 75-mg solid dispersion and the silicone outer layer with a polysiloxane copolymer outer layer [27
32]. The two-rod implant is marketed as Jadelle (Bayer, Germany). Another SE-based subdermal implant prescribed for long-term contraception is the Sino-Implant II (Shanghai Dahua Pharmaceutical, China) which is also a two-rod implant in which each rod contains 75 mg levonorgestrel. SEbased steroidal implants also find application in veterinary applications for estrus cycle control [33] or growth promotion [33
38]. In the beef industry, the subdermal release of hormones increases muscle tissue and decreases overall body fat in cattle, thus providing improved efficiency for farmers and leaner cuts of meat for the consumer [35]. Estrogen-releasing implants can contain estradiol or progesterone, while androgen releasing implants often contain trenbolate acetate or testosterone propionate [36,37]. An SE-based reservoir subdermal implant, Compudose (Elanco Animal Health, United States), is used as a growth promoter for cattle [39]. The implant provides zero-order release of estradiol and oxytetracycline for at least 200 days. A nonsteroidal application of SE-based subdermal implants has been used as a preexposure prophylaxis strategy for the sustained release of HIV-1 drug, tenofovir alafenamide [39]. PEVA-based subdermal implants have been investigated and marketed for a variety of clinical applications. Pellet-type implants were developed that were loaded with insulin for the treatment of diabetes [1]. In a rat-based animal model, the implants were shown to provide a 73% reduction in blood glucose levels for up to 26 days. Further, preclinical development of this implant system was able to increase the release duration by using a more soluble form of insulin [40] and having a more porous PEVA matrix structure [41]. A further in vivo clinical study using diabetic rats was able to provide a consistent reduction in glucose levels by employing a reservoir pellet system using a coating of PEVA over the insulin PEVA core. Implanon (Merck, United States) is a PEVA-based contraceptive subdermal implant containing etonogestrel [42]. The 4-cm long and 2-mm diameter

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rod is produced via hot-melt coextrusion of PEVA and the drug. Unlike the Jadelle implant, the Implanon implant is a single-rod device. The FDA has recently approved another single-rod PEVA-based implant for the treatment of opoid addiction [43]. The device provides the sustained release of buprenorphine for up to 6 months, and avoids the daily plasma peaks and troughs associated with conventional sublingual administration of the drug. Genina et al. fabricated both subdermal implant and T-shaped intrauterine devices and using 12 different grades of PEVA combined with the model drug indomethacin [44]. The implants were manufactured via Hot Melt Extrusion (HME) and fused-filament fabrication 3D printing, and each PEVA grade assessed from printability. They reported that 5 out of the 12 grades were printable and that one showed superior print quality. The US Army has recently begun investigating the suitability of PEVA-based subdermal implants for the prevention of malaria [45]. Their key aim is to develop implants with sustained efficacy for a minimum of 5 weeks and is safe and well-tolerated. Current chemotherapeutic strategies of malaria prevention is highly dependent on deployed troop compliance to a prescribed regimen which is not achieving desired targets. The research team investigated a piperaquine-infused PEVA-based implant in mice and were able to offer protection for up to 8 weeks after implantation. Endo Pharmaceuticals Solutions Inc patented a TPU-based reservoir subdermal implant for the zero-order release of histrelin to the treatment of central precocious puberty [46]. The implant is marketed as Supprelin.

5.4

Nondegradable vaginal rings

SE was the first material to be employed in the development of a vaginal ring implant in the 1960s [47,48]. The contraceptive device was loaded with medroxyprogesterone acetate, and many other contraceptive hormones have been utilized with SE-based vaginal rings. The first patents in this area describe matrix devices in which the outer layer was medicated [49,50]. The Population Council was the first to develop a reservoir SE-based vaginal ring for zero-order release [51]. The three-layer ring had a nonmedicated sheath layer and middle-core with a medicated sandwich layer. The ring was specified for the corelease of estradiol and either levonorgestrel, DL-norgestrel, or norethindrone. Manufacturing of this ring was a three-step process in which an injection molded inner core was dipcoated in liquid silicone containing the hormone compounds. A second dip-coating process provided the nonmedicated silicone sheath layer. The Population Council also developed an insert-core vaginal ring in which medicated cores could be inserted with any variety of suitable drug compounds [52]. Aktiebolaget Leo utilized a stepwise process to manufacture a reservoir SE-based vaginal ring to release 17β-estradiol as a menopause treatment [53]. Galen Pharmaceuticals entered the same marketplace by developing reservoir rings for hormone replacement therapy (HRT) for menopause [54] and testosterone therapy for premenopause [55]. Dow Corning used a coextrusion process to manufacture a reservoir vaginal ring [56]. Enhance Pharmaceuticals

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utilized a novel twin-injection molding process for fabrication of reservoir rings [57]. Three currently available vaginal ring products are SE based—Estring, Femring, and Progering. Estring (Pfizer, United States) is a reservoir ring prescribed as a hormone replacement therapy with the zero-order release of 7.5 μg per day for up to 90 days. Femring (Allergan, Ireland) is also a 90-day ring for HRT with estradiol acetate. Progering (Laboratorios Silesia, Chile) is a contraceptive for lactating women and provides a daily dose of 10 mg per day for 90 days. A 2017 study by Bayer AG investigated the effect of levonorgestrel content on the in vitro and in vivo properties of SE-based vaginal rings [58]. The company compared two rings that were designed to provide a daily release of 40 μg. The design of the ring was of a two-segment structure in which the main body was nonmedicated, while a short section was a drug-loaded reservoir. Two different drug-loaded segments were investigated, one containing 50% (by weight) drug and one containing 5% (by weight) drug. Both in vivo and in vitro studies showed that although total drug content differed greatly between both rings, zero-order daily release rates were the same, which underlines the importance of implant design in controlling release rates. Despite the long history of hormone-releasing SE-based vaginal rings, recent studies have described polymer-drug interactions. Murphy et al. highlighted an issue of levonorgestrel chemically and irreversibly binding to SE during the curing process in the production of SE-based vaginal rings [59]. The binding occurs via a hydrosilylation reaction between the drug’s ethinyl group and the hydrosilane groups in the cross-linker. A number of parameters were shown to influence the extent of binding, including cure time and temperature, drug particle size and loading, and SE resin type. A further study by the group used solid-state 13C nuclear magnetic resonance spectroscopy to also demonstrate the covalent, irreversible binding of another hormone compound (ethinyl estradiol) to SE during the curing reaction [60]. Beyond hormone release, SE-based rings are being developed as part of HIV prevention strategies. Nel et al. reported a Phase I randomized crossover trial to assess the safety of placebo SE-based vaginal rings [61]. Silicone placebo vaginal rings were used for 12 weeks followed by a 12-week safety observation period. Results reported from the study, which enrolled 170 African women (age, 18
35 years), that the vaginal ring possessed no safety concerns or adverse side effects. SE-based vaginal rings have been utilized for the release of a range of microbicides [62
71]. Vaginal microbicides are chemical compounds or biological agents that can reduce the transmission of HIV from sexual intercourse through a number of mechanisms: (1) virus destruction, (2) protective vaginal flora maintenance, (3) HIV virus CD4 receptor binding, (4) disrupting HIV replication process, and (5) preventing other sexually transmitted infections that create lesions that increase HIV transmission rates. Dapivirine is the most investigated microbicide, and an SE-based vaginal ring containing the drug is currently in Phase IIIb testing [20]. Dapivirine is a nonnucleoside reverse-transcriptase inhibitor developed by Janssen (United States) [72], and as a hydrophobic small molecule drug has the ideal physiochemical property profile for release from an SE-based ring [68
74]. The dapivirine ring is of a matrix construction and contains 25 mg of

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the drug, providing a sustained release to the vagina for up to one month and has demonstrated acceptability, safety, and tolerance [74]. SE-based vaginal rings have been a platform for the simultaneous release of multiple hormones [75
79]. In a similar vein, the material is being used in the fabrication of vaginal rings as combination microbicide products and multipurpose prevention technologies (MPT). Combination microbicide products release antiretroviral combinations of drugs with different mechanisms of action [80], so as to (1) target emerging resistant HIV mutations, (2) provide additive or synergistic effects that lower clinical concentrations for efficacy, and (3) target multiple stages of the HIV virus replication cycle. Combination SE-based vaginal rings are being developed that combines the NNRTI dapivirine with an HIV entry inhibitor (maraviroc) [61,74,81] and also with a protease inhibitor (darunavir) [82,83]. MPT rings are combining dapivirine with contraceptive hormones [84
86]. Unlike the dapivirine-only rings, these MPT rings are of a reservoir construction for the prolonged zero-order release of the microbicide and hormone [86
88]. A variety of hormones were investigated, but levonorgestrel has shown the most potential [86,88,89]. Boyd et al. further reported the simultaneous and continuous in vitro release of dapivirine and levonorgestrel over a period of 60
180 days from an SE-based vaginal ring [90]. Drug loading for dapivirine was at 100
200 mg and 16
32 mg for LNG. Dapivirine release was at 4132
6113 μg on Day 1 and 284
454 μg on Day 60. Levonorgestrel release was 129
684 μg on Day 1 and 2
9 μg on Day 60. McCoy reported the impact of dapivirine polymorphism on the performance of an SE-based vaginal ring via DSC, PXRD, and solubility analysis [91]. Results indicated that dapivirine packing polymorphism had no impact on the in vitro behavior. Murphy et al. investigated the importance of ring size and drug loading for microbicide-loaded SE-based vaginal rings [92]. They studied three different-sized matrix rings of significantly different surface areas containing a combination of dapivirine and darunavir. Although significant differences were observed for the in vitro drug release of the three-sized rings, there was not a significant difference in the serum and vaginal fluid drug concentrations during in vivo implantation. The study highlighted the importance of size matching of vaginal rings for specific species, as the less suitable of the three rings for these macaques was shown to entangle the cervix and initiate vaginal erosions. Another MPT device is the SE-based SILCS contraceptive diaphragm that also releases microbicide from with either a gel formulation placed prior to insertion [93
95] or from the internal thermoplastic spring core [96]. Wang et al. utilized misoprostol and isosorbide mononitrate in an SE-based vaginal ring for cervical ripening [97]. The ring was tailored to release the two Active Pharmaceutical Ingredients (APIs) at different release rates. Both misoprostol and isosorbide mononitrate had a reduced daily release rate over 14 days. PEVA-based vaginal rings have smaller body diameters to SE-based rings due to superior mechanical properties [98]. The first PEVA ring was developed in the 1990s and included the development of the Nuvaring product [99,100]. The contraceptive product is of a reservoir construction containing etonogestrel and ethinylestradiol in the medicated core, coreleasing a daily constant dose of both drugs.

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The ring is manufactured by melt-compounding both drugs with PEVA using hot-melt extrusion into a constant diameter extrudate which is cut to length and formed into torus shape. Each ring is put in place for three weeks followed by a ring-free week. A large user acceptability study (1950 women) found that 81% preferred Nuvaring to oral contraceptives [101]. Additional PEVA-based hormone releasing vaginal rings have been developed. Columbia Laboratories are exploiting a General Hospital Corporation invention of a segmented ring containing progesterone and estradiol in separate compartments [102,103]. Helbling et al. have reported on a PEVA ring produced via hot-melt extrusion with a progesterone in vitro drug release profile similar to Progering [104,105]. Merck Sharp and Dohme developed a hormone (estradiol or nomegestrol acetate) produced via the coextrusion of a trilayer PEVA extrudate of a nonmedicated sheath surrounding a medicated core and separated by an intermediate layer [106]. Kimball et al. reported on the successful in vivo administration (n 5 6) of the gonadotropin-releasing hormone leuprolide from a PEVA-based vaginal ring. The purpose of their study was to demonstrate the effect of nonparenteral absorption of a polar peptide via vaginal drug delivery [106]. Similar to SE-based vaginal rings, PEVA rings are being developed as microbicide products. McConville et al. developed PEVA vaginal ring to release UC781 at much higher levels than an SE-based vaginal ring during in vitro dissolution [107]. The authors argued that the release was greater as the drug was in the amorphous form in the PEVA matrix because of hot-melt extrusion processing, while the drug remained in the crystalline form in silicone. However, an in vivo comparison in rabbits of the two types of rings showed no significant difference in the pharmacokinetics of UC781 [108]. PEVA-based rings were also developed for corelease of this microbicide and levonorgestrel as an MPT product [109]. Other PEVA-based MPT vaginal rings have been described in the literature including a ring for releasing a microbicide (MIV-150), an antiherpes simplex compound (zinc acetate), antihuman papillomavirus (carrageenan), and a contraceptive hormone (levonorgestrel) [110]. An in vivo macaque study of the rings showed that all drugs released at effective levels for up to 28 days. Chemotherapeutic drugs have also been assessed with PEVA-based vaginal rings. Disulfiram, which has shown efficacy against a range of cancer cell lines, was incorporated with SE- and PEVA-based rings in a side-byside in vitro study [111]. Due to the sulfur groups present in the drug, silicone would not fully cure and these rings had to be abandoned during the study. The drug was successfully released from the PEVA rings at levels in excess of the IC50 for HeLa cell line for a 14-day duration. First-line cervical cancer drug cisplatin has also been incorporated into PEVA rings and successfully released [4]. Contraceptive TPU-based vaginal rings date back over 40 years to a patent by the Ortho Pharmaceutical Corporation for the release of the spermicide nonoxynol9 [112]. Due to this established principal, many patents filed later included the material as a possible base material for construction of the ring even though the actual marketed product was made from PEVA or silicone [52,54,109,113]. It was the late 2000s before the first patents were filed that listed TPU as the preferred material of construction. These included a Ferring B.V. patent for a water-swellable

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ring for the release of a range of hydrophilic APIs including polyurethane proteins and benzodiazepines [114]. A team at the University of Utah developed several TPU-based rings for the zero-order release of a range of drug compounds [115,116]. Another patent in 2011 describes the invention of surface-modified vaginal ring fabricated from a water-swellable TPU for the simultaneous release of multiple drugs [117]. A company invented a contraceptive vaginal ring that had a medicated-water-swellable TPU inside a PEVA-based sheath that was designed to reduce swelling, maintain structural features, and provide zero-order release [118]. The preferred hormones for the invention were listed as ethinylestradiol and levonorgestrel. TPU-based vaginal rings have been extensively investigated for the release of HIV microbicides, and these devices are particularly useful for releasing hydrophilic compounds, such as tenofovir, that is unsuitable for SE-based or PEVAbased rings [119
124]. The Patrick Kiser Lab (Northwestern University, United States) has pioneered the use of these rings in microbicide research and has engineered a means of delivering multiple drugs using segmented ring technology. His team welded drug-loaded segments together to form intact torus-shaped rings. Each segment consists of a different TPU chemistry that is matched to the physiochemical properties of individual drugs, for example, hydrophobic TPU to hydrophobic drugs. The multisegment ring is a perfect device for MPT technology that combines different microbicides and contraceptive hormones. A further iteration of the segmented ring includes reservoir segments of either a microbicide-loaded paste core or a hormone-loaded coextruded core [125]. An in vivo sheep study determined that both drugs were released at effective rates for over 90 days. A multiple vaginal simian-HIV challenge study of a reservoir TPU-based vaginal ring containing the prodrug tenofovir disoproxil fumarate showed complete protection [126
128]. The ring was fabricated from polyether TPU tubing containing a liquid-filled core containing drug and osmotic excipients designed to promote vaginal fluid ingress to solubilize the prodrug and drive faster drug diffusion. In 2018, a Phase I study of a tenofovir disoproxil fumarate-loaded TPU-based reservoir ring was terminated early to the development of Grade 1 vaginal and cervical ulceration [129]. Eight of the 12 women in the tenofovir disoproxil fumarate ring arm developed ulcers. None of the five women in the placebo arm developed ulcers. It is suggested that ulceration is due to fumaric acid or other metabolites inducing inflammation and/or interfering with epithelial repair of microabrasions associated with ring use and/or sex. Another in vivo study by the Kiser group investigated the release of a highly potent nonnucleoside reverse-transcriptase inhibitor IQP-0528 from a TPU ring made from a Lubrizol Tecoflex grade resin [130]. The researcher conducted a doseranging study using pigtailed macaques by using drug-loaded segments comprising 25%, 50%, or 100% of the total ring. In vitro dissolution studies showed a directly proportional relationship between the length of drug-loaded segments and drug release over 30 days. However, there was no statistically significant difference in the average daily release in vivo for vaginal fluid or tissue samples. The authors concluded this was because of poor water solubility of IQP-0528, and slow clearance from the cervicovaginal fluid that would create low drug concentration

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gradients slowing drug release. Traore et al. investigated the impact of hydroxychloroquine-loaded TPU-based vaginal rings to treat HIV-1 infection and other sexually transmitted infections [128,131]. The researchers used a combination of hot-melt extrusion and injection molding to manufacture the matrix rings. The rings were capable of providing controlled release of the drug for 24 days during in vitro dissolution, while not negatively impacting protective vaginal microflora or ectocervical epithelial cells. Verstraete et al. reported on the development of TPUbased vaginal rings for treatment of bacterial vaginosis with lactic acid or metronidazole [131]. The research team assessed different grades of TPU for production of the rings via the combination of hot-melt extrusion and injection molding. The rings provided release of lactic acid over a duration of 28-day period and of metronidazole release for up to 7 days. Each drug required a different grade of TPU and drug loading to provide the necessary drug release profile. Kim et al. reported on a segmented reservoir TPU-based vaginal ring with switchable on-demand release of a drug nanoparticle when the pH increased from 4.2 (normal vaginal pH) to 7.0 (in the presence of seminal fluid) [132]. The injection-molded reservoir ring segments were loaded with a pH-sensitive polyurethane bearing dimethylolpropionic acid hydrogel containing the drug nanoparticles. During in vitro dissolution testing, the ring would release around 20% of the nanoparticles during a 1-hour period at pH 7.0. At the lower pH, the ring released less than 2% of the nanoparticles during a 1-hour period.

5.5

Nondegradable ocular implants

Ocular implants are designed to treat vitreoretinal diseases. The Helios ring (Allergan, United States) is a bimatoprost-loaded ocular implant developed by ForSight Vision5 Inc. for the treatment of glaucoma [133]. This 24
29 mm ring consists of a nonmedicated polypropylene support core covered by bimatoprostloaded SE-based sheath. A Phase II clinical trial involving 130 patients demonstrated that this implant reduced intraocular pressure by 6 mmHg over 6 months. Escalon Medical Corp. (United States) patented Ocufit SR, which is a rod-shaped ocular implant made from SE for treating ocular disease or infections [134]. The implant is placed in the lower and upper conjunctival fornix. The Iluvien (Bausch & Lomb, United States) is an FDA approved ocular implant for treating chronic noninfectious posterior uveitis [135]. The implant consists of an SE-based cup that holds a fluocinolone-acetonide-loaded tablet covered by a polyvinyl alcohol membrane [136,137]. The implant is surgically implanted into the posterior of the eye. PEVA-based ocular implants have also been utilized for the treatment of chronic eye disorders including glaucoma and refractory chorioretinal diseases. Ocusert was one of the first ocular implants developed consisting of pilocarpine in alginate gel and PEVA-based rate controlling membrane [138
140]. Ocusert was developed in order to deliver sustained hourly dosages of 20 or 40 μg of pilocarpine for several days. The implant reduces intraocular pressure for glaucoma patients while also

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reducing miosis and myopia side effects due to the controlled constant dosing. Vitrasert (Auritec Pharmaceuticals, United States) is a treatment for cytomegalovirus retinitis in AIDS patients. It is essentially a PEVA-coated compressed tablet containing ganciclovir [138
140].The I-vation implant (Surmodics, United States) is used to treat diabetic macular edema [141]. It consists of a thin cap attached to a PEVA- and polyvinyl-alcohol-coated titanium helical coil. The cap is coated with PVA and EVA layers to control the release of the drug. The helical coil shape maximizes surface area for drug release and also enables sutureless anchoring of the device to the sclera [141].

5.6

Biodegradable implants

Although nondegradable implants are well established for localized drug release, the benefits of biodegradable implants have captivated the majority of future research. When compared to their nondegradable equivalent, biodegradable implants contribute to a more patient-friendly therapy. Where no secondary operation for removal is required, the disintegration of the implant allows for tissue regeneration within the implanted region and implant does not interference with radiological imaging. There is an innovative range of biodegradable materials available including biodegradable polymers, injectable in situ forming implants (ISFIs), bioresorbable ceramics, and biodegradable metal alloys. Varying mechanical properties can be achieved, allowing for optimization of implant for the intended use. Biodegradable materials, usually formed from polymers or metal alloys, are degraded initially by structural chains breaking down into smaller fragments, followed by phagocytosis of small particles by macrophages and/or chemical dissolution. Another form of biodegradable materials, known as bioimplants, are made of resorbable substances, for example, bioceramics and metal alloys [142]. They are used to aid the function of defected biological structures such as orthopedic or dental structures [143
145]. The ideal biodegradable implant should satisfy the following criteria: 1. Biodegradable in a physiological environment without inducing any harmful effect on the body. The rate of implant degradation and debris particle formation shouldn’t exceed what the tissue is able to tolerate. 2. Nontoxic, implant surface should permit surrounding tissue healthy cellular growth. 3. The degradation rate of the implant should match the rate of tissue regeneration.

Implant degradation rate is often influenced by factors such as implant geometry, contact with the body fluid, implant location within the body, temperature, motion, molecular weight of components, the crystallinity of components, and formulation. For biodegradable drug-loaded implants, the degradation rate of the implant must remain at a consistent rate to sustain a stable drug release profile. Two biodegradable drug delivery systems can be considered for drug release: (1) a reservoir system, where drug release occurs prior to implant structure degradation, the drug is released via diffusion through the implant matrix [146]. (2) Surface-eroding system, where system degradation is proportional to drug release [147,148] (Fig. 5.1).

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Figure 5.1 Drug release profile of reservoir and surface-eroding biodegradable implants [147,148].

As in the case for nondegradable implants, drug release profile is also influenced by several design factors including geometry, formulation components and loading ratio, implant microstructure, and the crystallinity of components [147,149
152]. Although nondegradable implants are well established for localized drug release, the advantages of biodegradable implants resulted in a research shift toward the use of biodegradable implants for drug delivery. In this section we describe four new and most common biodegradable implants systems: (1) biodegradable polymers, (2) ISFIs, (3) biodegradable ceramics, and (4) biodegradable metal alloys, and elaborate on their essential function, properties, and therapeutic capacity for drug delivery. In general, the development of biodegradable drug-eluting systems is more complicated than the nondegradable counterpart. The additional consideration for in vivo degradation must be controllable for sustained drug release and safety. Biodegradable implants may also incur additional costs of materials, research and development, and more extensive regulatory procedure. Nevertheless, the range of materials and new formulation techniques emerging provides a more favorable outcome for commercial success (Table 5.2). So far biodegradable drug-releasing implants have seen reasonable commercial success in cancer, infections, trauma, immune therapy, etc. This presents their commercial viability, cost-effectiveness, and clinical acceptance as an alternative route to sustained drug delivery. With the advancement of biodegradable material, no doubt new technologies will emerge and matters such as development cost will fade.

5.7

Biodegradable polymers

Polymers are macromolecules composed of covalently bonded monomers. Repeating monomer units can be assembled from either one molecule (homopolymer) or from multiple (copolymers). For biodegradable polymers, both natural and

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Table 5.2 Summary of biodegradable materials: their properties and functions. Material

Properties

Application examples

Biodegradable polymers

Flexible and can be fabricated into a variety of different shapes and structures. Low elastic modulus and low strength Controlled and tailored properties Easily administered through a small needle Variable in implant shape and structure Bioactive Low tensile strength and high compressibility Stiff High mechanical strength Bioactive in some cases

Slow and fast drug delivery of chemotherapeutics, proteins, antibiotics, and antiinflammatory drugs

In situ forming implants

Bioactive ceramics Biodegradable metal alloys

Dental and cancer drug delivery applications

Applications limited to applications, such as small bone fractures, where loading is relatively low Bone fixation Stents

synthetic varieties are available. Synthetic polymers include starch, chitosan, collagen, fibrin gels, and derivatives of hyaluronic acid [153]. Although synthetic polymers possess great biocompatibility and are inexpensive, they are rarely utilized for drug delivery because they lack in mechanical strength and predictability for interactions when in vivo. Synthetic polymers while expensive are more practical for drug delivery applications. Some of the most commonly used are poly(lactic-coglycolic acid) (PLGA), polyglycolide (PGA), polylactic acid (PLA), polyhydroxybutyrate, and PCL (Table 5.3) [153]. PLGA, Poly(lactic-co-glycolic acid); PGA, polyglycolide; PLA, polylactic acid; PHB, polyhydroxybutyrate; PCL, polycaprolactone. Biodegradable drug-loaded polymer implants are often forged into relatively simple, homogenous, and singular devices. These devices are composed of therapeutic agent incorporated homogenously within either a polymer or polymer plus additives matrix. During the development procedure two main design processes are involved: (1) formulation and (2) manufactory. The design procedure is extremely extensive with considerable knowledge of polymer mechanical properties required. During the formulation stage, a polymer with the most appropriate mechanical properties for intended use is selected. For instance, if long-term drug release is required, then a polymer with slow degradation rate is ideal. Additives such as plasticizers, fillers, stabilizers, and excipients can be selected to enhance implant’s mechanical properties and reduce manufacturing costs [148]. Currently, amongst the many biodegradable polymers that are available: PGA, PLA, and their copolymer PLGA are most utilized [161]. This can be attributed to

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Table 5.3 Summary of common biodegradable polymers. Polymer

Repeating units

Mechanical properties

Applications

PLGA

Copolymer of PLA and PGA

Short-term drug release [154]

PGA

Homopolymer

PLA

Homopolymer

PHB

Homopolymer

PCL

Homopolymer

50:50 Degradation rate 1
2 months [154]. 75:25 Degradation rate 4
5 months [154] 85:15 Degradation rate 5
6 months [154,155] Glass transition temperature (Tg) decrease with lower lactide ratio and with lower molecular weight [153,156] High crystallinity [15] Low solubility [15] Slow degradation rate [16] Tg 5 35
45 C [17,18] PLA can be highly crystalline (PLLA) or amorphous (PDLLA) [157] PLA Brittle [158] Slow degradation rate: up to several years [157] Tg 5 45 C
60 C [159] PLLA 37% crystalline Tg 5 60
65 C [157,160] Slow degradation rate [157] PDLLA Tg 5 50 C
60 C [158,160] High crystallinity Inflexible and brittle Tg 5 5 C
15 C [157,160,161] High flexibility Low mechanical strength Slow degradation rate [157,160
162] Tg 5
60 C [18,19]




Varied release profile

Long-term drug release [153] Long-term drug release [153]

their (1) ease of manufacture by several techniques. Such as extrusion, compression molding, melt casting, solvent casting, and electrospinning. (2) Well-documented record of biocompatibility, degradation pathway, and FDA preapproval within use in pharmaceutical products. This eases the regulatory process for an implantable device to reach the market [163]. During the manufacturing stage, the formulation is shaped into a defined geometric shape, which leads to a predictable and reproducible drug release and degradation profiles. Drug-loaded polymer implants have been utilized for the delivery of chemotherapeutics [164
167], proteins [168,169], antibiotics [170,171], antiinflammatory

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drugs [172] etc. With clinical success reached when the FDA approved the use of biodegradable drug-loaded wafer, Gliadel. The implant is placed during brain resection surgery on tumor surface to deliver controlled-release carmustine dose [173]. With the recent clinical success and abundance of polymer materials with varying mechanical properties. A promising endeavor is emerging for the development of new drug-loaded biodegradable polymer implants.

5.8

Injectable in situ forming implants

In injectable ISFIs, a drug-loaded biodegradable polymer is dissolved in a biocompatible solvent to produce a liquid polymer formulation [148,174,175]. The API can be either dissolved in the liquid matrix as a solution, suspension, or dispersion [34]. Most commonly utilized polymers are polylactides, polyglycolides, caprolactones, polyanhydrides, and polyorthoesters [174,175]. To administer, ISFIs are injected using conventional needles and syringes into the active site. Upon injection, the polymer solution solidifies to form a solid or semisolid in situ implant, and the drug is trapped or encapsulated inside [176
180]. ISFIs phase transition can occur in response to several different mechanisms, such as solvent exchange, temperature transition, pH transition, ionic crosslinking, UV irradiation, and enzymatic activity [175,181
184]. Depending on their mechanism for the phase transition, ISFIs are divided into six categories: thermoplastic pastes, thermally induced gels, in situ precipitation, in situ cross-linked systems, in situ gelling organogels, and hydrophobic fatty-acid-based injectable pastes [148,185] (Table 5.4). PLGA, Poly(lactic-co-glycolic acid). For drug delivery, ISFIs are an attractive option: they are versatile, easy to administer, and relatively inexpensive to manufacture [148,184,205,206]. The wide range of polymers and solvents available enables for property optimization, for example, sustained drug release and degradation rate of the implant can be adjusted from days to months [174,175,184]. The only notable disadvantage for usage of ISFIs is the initial bust in drug release. The bust effect can occur due to a number of reasons associated with ISFIs’ implant formation: Time delay between administration and implant formation, the release of drug adsorbed to the implant surface, and unequal drug distribution [175,187,207
210]. However, with intelligent formulation design, this bust effect can be reduced [175]. Drug-loaded ISFIs have been enormously successful in clinical applications. Products like Eligard (leuprolide acetate for prostate cancer) [211], PerioChip (chlorhexidine digluconate for gingivitis) [212], Atridox (doxycycline hyclate for periodontal disease) [213], Arestin (minocycline hydrochloride for periodontitis) [214], and Zoladex (goserelin acetate for endometriosis and prostate cancer) [215,216] are some of the currently marketed products.

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Table 5.4 Injectable in situ forming implants phase transition categories. Category

Phase transition mechanism

Polymers

Thermoplastic pastes

Cooling down to body temperature

Thermally induced gels

Sol
gel transition

In situ precipitation

Phase inversion

In situ cross-linked systems

Cross-linking through chemical or physical (e.g., heat, photon absorption, ionic mediated reactions) Hydrogen bonding

Polycaprolactone (PCL), poly(D,L-lactide), poly (glycolide) [186
188] PLGA
Polyethylene glycol (PEG)
PLGA, PEG
Poly (p-phenylene oxide) (PPO)
PEG, PEG
PLGA
PEG, chitosan, poly(N-isopropyl acrylamide), PEG [189
194] PLGA, poly(ethylene carbonate) [195,196] PEG, alginate, PCL [197
200]

In situ gelling organogels Hydrophobic fattyacid-based injectable pastes

Upon contact with water

Amphiphilic lipids [201,202] Poly[(sebacic acid)-co(ricinoleic acid)] [202
204]

Biodegradable ISFIs have proven to be versatile for therapeutic drug delivery. The current groundwork research is stupendous. If the rate of development continues, it is likely that many more applications and new ISFI implants will be manifested.

5.9

Bioresorbable ceramics

Bioceramics are generally defined as solid crystalline material consisting of inorganic compounds held together by covalent or ionic bonds [217,218]. Bioresorbable ceramics, such as calcium phosphate ceramics and silica-based glasses, are designed to support, replace, or aid in healthy cellular tissue growth. Orthopedic applications, such as significant bone weakness, are commonly treated with bioceramics [138,209
211]. Amongst the list of available bioresorbable ceramic materials, hydroxyapatite, dicalcium phosphate, dehydrate, and tricalcium phosphate are the most extensively researched [153,219]. Bioceramics generally have low tensile strength and high compressibility [220,221]. Similar to polymers, the mechanical properties and the application of the

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material are dependent on the preparation process. They can be prepared via a variety of different methods, such as high-temperature methods, controlled precipitation, and the sol
gel process, where the final formulation could be in the form of granules, self-hardening cement, coatings, and porous devices [222
225]. With many hard tissue degenerations occurring because of a disease requiring drug therapy, for example, tumors, infections, and osteoporotic fractures. A growing interest to incorporate drugs within or on the fictionalized surfaces of bioceramics is developed [225
232]. Implantation of drug-loaded bioceramics in functions such as orthopedic or dental can provide simultaneous structural support and localized drug release. Compared with polymer-based systems, bioceramics can achieve enhanced structural rigidity and bioactivity. For hard tissue applications, this makes it the more suitable implant candidate. For drug loading, there are effectively two ways bioceramics can incorporate drugs: inside the porous structure of materials and the less utilized method of polymerized drug attachment to a material’s functional groups [226
232], with both techniques capable of attaining controlled drug release [230
232]. Porous bioceramics, also known as mesoporous bioceramics such as calcium phosphates or bioactive glasses can serve as a reservoir for drugs (e.g., antibiotics, anticancer drugs, antiinflammatory drugs) [225
232]. The porosity of the material can aid a favorable microenvironment for tissue growth by facilitating a space for diffusion of nutrients, oxygen, and metabolic products [153,233]. Increased porosity can enhance drug-loading capability, but it can be detrimental for mechanical strength [153,234]. Hence for biomedical applications, an equilibrium between porosity and mechanical strength is necessary. Another method for drug loading is to attach active therapeutic substances to bioceramics-functionalized groups [226
232]. This drug delivery method can be used to avoid rapid drug degradation and systemic clearance [87]. It often also results in a more compact material arrangement with enhanced mechanical strength [231]. For drug-loaded bioceramics, there are several design and characteristic features that have to be considered: 1. Nontoxic within the in vivo environment. 2. Biocompatible: Implant interphase can interact and form an attachment to a tissue. For example, bioactive ceramics can cultivate healthy cellular growth through assisting in the production of proteins, cell adhesion, and stimulating regeneration of tissue [226]. 3. Biodegradable/bioresorbable: A controlled degradation rate that is simultaneous with the replacement and growth of a natural host tissue. During degradation of the implant, loading is redistributed to tissue; this avoids stress-shielding effect [235,236]. 4. Shape. The geometry of implant is often chosen to mimic natural body constituents. 5. Surface modification: Drug or proteins may be loaded onto the surface of bioceramics to enhance delivery and interactions with cells [226]. 6. Mechanical properties: Capable of withstanding mechanical stresses and strains. This is often highly dependent on pore size and shape of material [153,228]. 7. Pore size: The pore size can range from hundreds of micrometers down to the nanometer scale. A minimum of 100 μm pore size is needed for cellular proliferation and diffusion of nutrients, oxygen, and metabolic products [153,228].

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Drug-loaded bioresorbable ceramics are an outstanding alternative to polymer-based implants. With their ability to facilitate bioactivity and contribute to the structural support of biological structures, they contribute as an essential addition to the variety of biomaterial engineering approaches available. Currently, the bioceramics field of applications is limited to the field of low loadbearing bone pathologies and dental implants. Many researchers have assessed bioceramics capability for drug delivery with success. It has proven to provide synergy capability for drug delivery, bioactivity, and biological structure support. However, with a limited range of applications, this hinders the therapeutic development of drug-loaded bioceramics. A recent upsurge in market demand within road accidents and aging population could reorientate research focus for bioresorbable ceramics.

5.10

Biodegradable metal alloys

The concept of employing drug-eluting biodegradable metal alloys is very recent. Zinc, copper, magnesium, and iron have all been investigated to reveal excellent biodegradability and are proposed as better materials for load-bearing application, because of their enhanced strength against polymers and ceramics [237
241]. Additional advantages for the use of copper alloys is its ability to minimize the risk of infections. Copper possesses antibacterial properties along with the ability to play a role in enhancing the immune system and endothelial cellular proliferation [242
245]. Many binary or ternary metal alloys are formed, for example, Mg
Zn, Zn
Cu
Mg to optimize implant properties such as degradation rate, strength, and toxicity. Obtaining a well-controlled degradation rate is particularly crucial for metals as their excessive bloodstream intake could lead to irreversible toxic damage including neurodegenerative diseases such as Alzheimer’s, Menkes, and Wilson [245
247]. Most research for biodegradable metal alloys is focused on replacing current nondegradable equivalents such as the use of metals in stents, bone fixations, plate, and screw, etc. But few drug delivery concepts have developed. A clinical trial assessment by Haude et al. [248] for sirolimus drug delivery using absorbable metal scaffold in patients with de novo coronary lesions, showed a sustained safety profile and drug-eluting capability. This provides evidence for the commercial capability of drug-eluting biodegradable metal alloys. Another commercial development recent stride is the clinical emergence of the first biodegradable metal alloy (magnesium based) product. Magnezix screw, employed for bone fracture fixation [249]. In conclusion, biodegradable metal alloys can serve as drug delivery systems for application requiring enhances structural rigidity.

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1365. Available from: https://doi. org/10.3390/ijerph7041342. [248] M. Haude, H. Ince, A. Abizaid, R. Toelg, P.A. Lemos, C. Von Birgelen, et al., Sustained safety and performance of the second-generation drug-eluting absorbable metal scaffold in patients with de novo coronary lesions: 12-month clinical results and angiographic findings of the BIOSOLVE-II first-in-man trial, Eur. Heart J. 37 (2016) 2701
2709. Available from: https://doi.org/10.1093/eurheartj/ehw196. [249] R. Biber, J. Pauser, M. Geßlein, H.J. Bail, Magnesium-based absorbable metal screws for intra-articular fracture fixation, Case Rep. Orthop. 2016 (2016) 9673174. Available from: https://doi.org/10.1155/2016/9673174.

Further reading J.Z. Yang, R. Sultana, X.Z. Hu, P. Ichim, Novel layered hydroxyapatite/tri-calcium phosphate-zirconia scaffold composite with high bending strength for load-bearing bone implant application, Int. J. Appl. Ceram. Technol. 11 (2014) 22
30. Available from: https://doi.org/10.1111/ijac.12024. Z. Sheikh, S. Najeeb, Z. Khurshid, V. Verma, H. Rashid, M. Glogauer, Biodegradable materials for bone repair and tissue engineering applications, Materials (Basel) 8 (2015) 5744
5794. Available from: https://doi.org/10.3390/ma8095273.

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Lilith Mabel Caballero-Aguilar1,2, Saimon Moraes Silva1,2 and Simon E. Moulton1,2 1 ARC Centre of Excellence for Electromaterials Science, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, VIC, Australia, 2 BioFab3D@ACMD, St Vincent’s Hospital Melbourne, Fitzroy, VIC, Australia

6.1

Introduction

Three-dimensional (3D) printing, which can also be referred as additive layer manufacturing, is a rapid prototyping process, which permits the fabrication of a physical object from a computer-aided design (CAD) software [1]. The 3D printing methodology was first introduced by Charles Hull in 1986 as an attempt to improve the plastic fabrication process. Charles Hull developed the first 3D printer the so-called “stereolithography apparatus” and first commercial printer SLA-250, which was a breakthrough for manufacturing process of different industry sectors [2]. Afterwards, the 3D printing techniques, selective laser sintering and fused deposition modeling (FDM) were developed by Carl Deckard in the late 1980s and Sachs et al. in 1990, respectively [1,3]. Over the preceding two decades, 3D printing technologies have evolved and been applied in varied fields either to enhance functionality of existing systems or as an innovative manufacturing process [4]. Since the first drug-formulation (SPRITAM) manufactured by a 3D printing technology reached the market, there is a revitalized attention in the exciting sector of drugs manufacturing, not only for solid oral dosage form but also to make personalization of DDSs and multifaceted controlled DDSs [5]. Promising results using 3D printing have been reported in implants, tablets of different geometries, and microneedles for transdermal routes [6
8]. The advantages of employing 3D printing for dosage form designs include the capability of precisely controlling the spatial distribution of an active pharmaceutical ingredient (API) within a dosage form, deposit minute quantities of API, can generate complex geometries, diminish the waste and permit for quick production of variable compositions to allow for screening activities on preparation of customized pharmaceutics [9
12]. In 3D printing, a drug formulation is produced layer-by-layer employing a printer. Usually, three stages are involved in manufacturing a drug: (1) the drug formulation is designed in a computer, (2) the 3D design is then translated into several single layers, and (3) finally printing the drug formulation. A CAD software is used Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00006-8 © 2020 Elsevier Ltd. All rights reserved.

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to generate a design of the dosage form [2]. This design is then divided into small fragments to stablish the required processing parameters including printing speed, the platform and extrusion temperature, and infill percentage. After all the parameters have been selected, it can be executed on the 3D printer to fabricate the desired drug formulation. Contingent on the printing technique utilized, inks or filaments can be made using suitable excipients [13]. The quality of the produced drug can be adjusted by altering the fabrication parameters and/or formulation. Usually, the different 3D printing techniques share some common steps as, for example, the initial material processing which frequently involves actual API with the other excipients [2,14]. The next step involves the printing followed by other usual procedures such as polishing and drying. The selection of method is fundamental to attain the optimum desired drug formulation using 3D printing [13]. In this sense, this chapter discusses the main three 3D printing methods (direct write: pressureassisted system, FDM, inkjet printing) and some of the variations frequently reported in the literature for the production of drug delivery systems (DDSs).

6.2

Direct write: pressure-assisted systems

The term “direct write” describes fabrication approach in which an ink is extruded through a nozzle to form a pattern or structure. The deposition can be achieved by creating lines or droplets [15]. A 3D architecture is achieved by subsequent extrusion, after the first layer is deposited the layer must be solid enough to allow the deposition of the next layer on top. As in other additive manufacturing techniques, direct write 3D printers that are bench-based rely on the use of a CAD model to form a predefined object by depositing the ink material. In the case of direct writing, the material properties of the ink as well as the type of printing system play a crucial role in defining the extrusion rate, the complexity, and resolution of the formed structure [16]. The extrusion of materials can be assisted by pressure through air injection in syringes or microtubes at room temperature. The control of the flow rate is usually achieved through pneumatic or mechanical systems based on pressured-air piston (B3
5 bars) or pressure-force extrusion (e.g., 10 N) that dispense a viscous material (Fig. 6.1). The material requires to be viscous enough to be controlled and at the same time needs to pass throughout the nozzle without creating clogs [16]. Extrusion deposition 3D printers using pressure-assisted systems that can deposit hydrogels have been explored for tissue engineering due to the high biocompatibility of hydrogel materials. Hydrogels are formed by natural or synthetic polymers and can retain large amounts of aqueous solutions making them highly biocompatible. Due to these properties, hydrogels have been 3D printed to fabricate scaffolds and bioengineered anatomical shapes. Incorporation of water-soluble drugs into hydrogels facilitates the preservation of the drugs bioactivity. Therefore investigation of these materials into 3D printers have been the focus of recent investigations seeking controlled DDSs [17,18].

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Figure 6.1 Pressure-assisted three-dimensional printer using two printing heads and a movable stage: photograph (A) and scheme (B). The printer allowed the production of bilayer tablets (C). Scale bar is 10 mm. Source: Reproduced with permission from S.A. Khaled, et al., Desktop 3D printing of controlled release pharmaceutical bilayer tablets, Intl. J. Pharm. 61 (1
2) (2014), 105
111, Elsevier.

Recently, the group of Khaled and coworkers developed a bilayer tablet by using hydrogel Methocel E5 and Methocel K100M/Carbopol 974 P NF extruded from a pressure-assisted 3D printer [16]. The printer uses a double-syringe nozzle and a pneumatic system to deposit layers of two different materials loaded with the drug guaifenesin. The aim of this study was to obtain a tunable release profile by changing the polymers’ concentration. The bilayer tablets showed an initial burst release from the Methocel E5 layer followed by a sustained release from the Methocel K100M/Carbopol 974 P NF up to 12 hours.

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Figure 6.2 (A) Gelatin microparticles, scale bar 5 200 μm. (B) Three-dimensional printed alginate scaffolds with gelatin microparticles, scale bar 5 2 mm. Source: Reproduced with permission from M.T. Poldervaart et al., Sustained release of BMP2 in bioprinted alginate for osteogenicity in mice and rats, Plos One 8 (8) (2013).

As 3D printing is rapidly being utilized in the area of tissue engineering, there arises an opportunity to use 3D printing to incorporate drugs into the tissue engineering structure. The combination of scaffolds with specific cytokines and growth factors that can potentially drive specific cell’s fate has been the focus of recent investigations [19]. For example, Poldevaart et al. used gelatin microspheres bone morphogenetic protein 2 loaded with alginate hydrogels to create a 3D printed scaffold for bone regeneration [18] (Fig. 6.2). The results from the in vitro experiment showed a continuous release of bone morphogenetic protein up to 2 and 3 weeks. In vivo testing was performed using subcutaneous implantation of the scaffolds in mice or rats proving that the release of the factor directs osteogenic differentiation. Another application of direct write 3D printing pressure-assisted in tissue engineering is the facilitation of drug-screening. Chang et al. automated a syringepressure-based system to direct write 1 3 106 cells/mL contained in alginate directly into a CaCl2 bath to form ionic cross-linked “microorgans.” The aim of this investigation was to demonstrate that through the use of direct writing it is possible to bio fabricate a 3D cell-encapsulated construct to be used as in vitro pharmacokinetic model [20]. The system was designed to operate at low pressure, patterning multiple cell lines in a predefined architecture and enabled with different nozzle sizes (30
500 mm) to deposit alginate hydrogels with varied viscosity. The printed “microorgan” was coupled to a device to allow continuous flow perfusion of drugs [20]. As stated above the ability to integrate water-soluble molecules directly into hydrogels has given rise to an opportunity to encapsulate a wide range of compounds due to the hydrophilic
hydrophilic interaction. Moreover, using hydrogels for 3D printing drug delivery features good biocompatibility and biodegradability. However, 3D printing hydrogels are not exempt from drawbacks. Hydrogels often have weak mechanical properties that can challenge the formation of 3D structures. Another drawback of hydrogels in delivering biomolecules is the characteristic burst release that these materials present when in contact with an aqueous media. For these reasons, strategies to overcome the drawbacks from hydrogels include the

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use of more than one hydrogel to create a double network [21] as well as the chemical modification of the hydrogel [22]. The entrapment in emulsions and micelles has also been investigated by mixing a nonwater-soluble compound into a hydrophilic polymer [23]. In the same context the use of materials that are not water soluble has been explored. The drug-eluting mechanics can depend on the bulk degradation of the hydrophilic material or diffusion throughout porosities of the formed structures. However, to 3D print polymers that are thermoresponsive different 3D printing methods must be applied.

6.3

Fused deposition modeling

FDM is a process to create 3D models that uses temperature control to melt a filament. A thermoplastic material is supplied in a filament that passes throughout a heated nozzle that will melt the polymer to be extruded at certain rate and pressure. The temperature parameters can be controlled based on the type of polymer to be extruded taking the material melting and glass transition (usually between 60 C and 200 C) into consideration. In general, the melting point will determine the drop pressure needed to extrude the polymer. If the polymer has a high melt viscosity, close to the limit of the printing heating capabilities, the extrusion produces rough surfaces which can limit its printability [24]. In FDM the nozzle tip or head can move in three directions depositing a thin layer onto a platform. The extruded material will lose heat and then solidify, forming a 3D object layer-by-layer. It is also possible that the print head is static, in that case the platform will move allowing the formation of a layered object (Fig. 6.3) [25]. Another possibility of FDM is to have more than one print head in which case each can be controlled independently allowing to extrude different materials

Figure 6.3 Fused deposition modeling process. A filament passes through a heating chamber (I), the filament is melted and extruded while the building platform moves down (II) until the last layer is deposited (III). The product is then ready for postprocessing and finished (IV). Source: Reproduced with permission from R.A. Perez, H.W. Kim, Core-shell designed scaffolds for drug delivery and tissue engineering, Acta Biomater. 21 (2015) 2
19, Elsevier.

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deposited based on the CAD model and the printer parameters [26]. This type of printer is particularly useful to create robust delivery systems in which it is possible to print various filaments containing different drugs or to provide additional support material to the extruded polymer containing the drug. Owing to the physical and chemical properties of the thermoplastic materials used in FDM, most drugs that are integrated in the filaments are nonwater-soluble. A method to introduce the drugs in the filament is by impregnation. In this method, a filament is soaked for a period of minimum 24 hours in a fluid containing the drugs; the filament will be then dried to allow the impregnation of the active drug that will be later released by diffusion mechanism [27]. A disadvantage of this method is that it requires a high concentration of the drug coupled with a significant time to allow the drug to penetrate the filament matrix. Another method consists of having the polymer in its fluid form by melting [hot melt extrusion (HME)] it or dissolving it (e.g., using an organic solvent) to mix it directly with the drug. The next step is to extrude the raw material to form a new filament loaded with the drug that will be directly printed or collected for further printing [28,29]. Owing to its simplicity, HME is a popular method to integrate drugs within the filaments. A drawback of this method is that the melting temperature that the polymers need to reach can sometimes compromise the bioactivity of the substances [30]. An interesting strategy to load and extrude drugs without compromising its bioactivity is to choose a drug with a melting point above the melting temperature of the polymer host. Zhang et al. mixed polylactic acid (PLA) with the drug acetaminophen (melting point is B170 C) and melt extruded the composite at 140 C
160 C using a FDM 3D printer coupled with HME [24]. The release of the drug was achieved upon initial diffusion (short time frame) and bulk degradation (longer time frame) from the printed structure. A similar approach was followed by Shim et al. where tobramycin, a highly thermostable drug, was integrated in its powder form to a liquefied poly-ε-caprolactone
polylactic-co-glycolic acid polymer blend (PCL
PLGA) to be extruded forming 3D printed scaffolds for treatment of chronic osteomyelitis (Fig. 6.4) [31]. The same research group used this approach to load 5-fluorouracil, an anticancer drug, into a PCL
PLGA 3D printed patch to control pancreatic cancer further proving the capability of these loading approach to integrate drugs [32]. FDM has been widely employed to prototype drug delivery devices, such as tablets. For devices that aim to be implanted or used as delivery vehicle within the body, the material choice is crucial. Thermoresponsive polymers that are suitable for biomedical applications are carefully selected by their physical characteristics and their degradation. Polymers that have high biocompatibility, such as PCL and PLGA, have been used in FDM. Melocchi et al. used a swellable/erodible polymer [hydroxypropyl cellulose (HPC)] to fabricate an oral form of medication that contained acetaminophen [25]. The material was employed in conjunction with PLA and PLA
PEG (2%
10% by weight) to fabricate the delivery device [33]. The capsular device was able to rapidly release up to 100% of the total content between 60 and 90 minutes after it was immersed in aqueous solutions. In a later publication the same research group created a second version of this device.

Three-dimensional printed drug delivery systems

PCL and PLGA

Molten PCL/PLG (110°C)

153

Tobramycin (powder)

Blending

Dispensing material

3D Tobramycin-loaded scaffold

Temperature : 120°C~130°C Dispense : pneumatic pressure

Figure 6.4 Polycaprolactone (PCL) and polylactic-co-glycolic acid (PLGA) was melted and mixed with tobramycin. The solution was loaded in a fused deposition modeling 3D printing system and used to construct a drug-loaded scaffold. Source: Reproduced with permission from J.H. Shim, et al., Three-dimensional printing of antibiotics-loaded poly-epsilon-caprolactone/poly(lactic-co-glycolic acid) scaffolds for treatment of chronic osteomyelitis, Tissue Eng. Regen. Med. 12 (5), 2015, 283
293.

The capsular device was divided into compartments to host versatile release profiles. The delivery device was able to yield successive release pulses of a drug tracer before complete degradation. Continuing on from this, the group of Chai et al. created hollow-structured tablets for intragastric drug delivery [34]. Domperidone (DOM) which has an insoluble weak base (pKa 1 5 7.8, pKa 2 5 11.5), was chosen as a model drug. DOM was loaded into HPC filaments using HME. The tablets were formed by an outer shell composed of DOM-loaded HPC filaments while the inner was air-filled to maintain low density (Fig. 6.5). The formulation showed sustained release and floatability for 10 hours in vitro [34]. Devices that aim to be used in drug delivery in an external form of delivery can also be 3D printed using FDM. The advantage of these devices is that the internalization of the printed material and its degradation products is not necessary; therefore the biodegradation of the device is not a requirement. However, biocompatible materials are needed if the device is in direct contact with a body part [35]. A recent example of this type of devices is topical patches. In Goyanes et al., a personalized antiacne drug-loaded patche was fabricated using HME FDM of Flex EcoPLA (FPLA) and PCL filaments that were loaded with salicylic acid [36]. This patch was intended to work as a therapeutical topical treatment that can be removed and replaced daily, therefore the disintegration of the materials was not necessary. Similarly, in Genina et al. the authors investigated ethylene vinyl acetate (EVA)

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Figure 6.5 Three-dimensional printed tables (A, B) and scanning electron microscope (SEM) of their side surfaces (a, b). Drug release profiles of the formulations used in the experiments comparing a commercial fast-release tablets and the 3D printed [domperidone (DOM)-floating sustained release (FSR)] tablets. Source: Reproduced with permission from X.Y. Chai et al., Fused deposition modeling (FDM) 3D printed tablets for intragastric floating delivery of domperidone, Sci. Rep. (2017) 7, (https://creativecommons.org/licenses/by/4.0/).

copolymers to elute a drug model (indomethacin) [37]. The EVA-drug melt polymer was extruded, and 3D printed to form custom-made T-shaped intrauterine systems and subcutaneous rods [37]. These types of devices are intended to be removed before the degradation of the material occurs.

6.4

Inkjet printing

As the pioneering work by Sachs and coworkers reported in 1996, the inkjet technology for 3D printing in pharmaceutics has come a long way [3]. It is the technique used to formulate Spritam, the first 3D printed drug approved by the Food and Drug Administration and released to the marked in 2016 by Aprecia Pharmaceuticals. The inkjet printing can be described as an additive manufacturing methodology, digitally controlled, that uses formation and deposition of fluid droplets with a systematic control over droplet volume, speed and motion of the travel, and the production time [38
40]. There are essentially two inkjet printing approaches, so-called continuous inkjet printing and drop-on-demand inkjet printing [41]. In the continuous inkjet approach a stream of droplet is created through the Rayleigh instability of a liquid column expelled through a nozzle. The nozzle is positioned at a potential relative to the substrate that transfers a small charge onto each drop. Separate drops are directed by applying another potential to the deflector (Fig. 6.6A). Continuous inkjet nozzles generate a continuous stream of droplets and the undesirable droplets are deflected into an ink collector, which can be recycled. One advantage of this approach is the high rate of droplet production per minute

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Figure 6.6 Schematic representation of (A) continuous inkjet printing and drop-on-demand printing with (B) thermal and (C) piezoelectric actuation.

(can be .50 kHz and droplets can expelled at speed .10 m/s) [41]. Despite the great production speed, this technique is limited in terms of placement accuracy. The drop-on-demand approach can be categorized as drop-on-solid and drop-ondrop. In the drop-on-solid approach, the formulated droplets from the printer head are deposited onto solid materials, while in the drop-on-drop approach formulated droplets are deposited onto each other in order to produce a solid layer of the rapid prototyping building material [42]. In both approaches a printer head is present, which can be a piezoelectric or thermal translator. With the piezoelectric head velocity waves are created in a fluid to produce a droplet from a nozzle when required. In other words, the piezoelectric crystal quickly changes leading to a volume change. This change produces an acoustic pulse that exposes the fluid to shear rates of up to 105/second, creating enough pressure to expel the droplet (Fig. 6.6C) [43]. The thermal heads make the use of resistive materials to generate current. In brief, when the current pulses over the electrical resistance in these materials, the temperature increases till around 300 C, which causes the evaporation of a small volume of liquid [41]. The produced bubble expands and transmits the necessary energy to expel a droplet (Fig. 6.6B) [41]. One advantage of piezoelectric translators over thermal inkjet systems is the variety of inks that can be used, while only water can be used as a solvent with thermal heads due to the elevated temperatures generated. Similar sizes (10
50 μm) and volumes (1
70 pL) of droplets can be obtained using both print heads [10].

6.4.1 Drop-on-solid inkjet printing In the drop-on-solid technique, the final solid product is obtained by the layer-bylayer deposition of small droplets of binder liquid onto a layer of powder. The binder droplets facilitate the binding of loose particulates to form solid structures within the layer [10]. After the first layer of droplets is printed onto a powder layer, another thin layer of powder is deposited on top, and the droplet printing process is repeated. The binder liquid can also work as a powder “bed” after its own

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solidification. After fabrication the final product goes through a thermal sintering process to guarantee its perpetuity and remove residual solvents. There are many advantages of using drop-on-powder process for producing pharmaceutical formulations. The well-defined deposition of binder droplets permits precise and accurate dosing of active ingredients, being possible to achieve even microgram quantities. This well-defined deposition of binder also allows the creation of multifaceted internal geometries of defined spatial composition and it has resulted in the achievement of customized drug release profiles. Lastly, drop-onsolid printing permits for rapid prototyping of dosage forms since it eliminates the necessity for any custom tooling for new designs. During the fabrication of pharmaceutical formulations, two main parameters have been shown to affect the performance of this technique, (1) how the material reacts with the binder and (2) the powder topology [44]. In terms of topology, particle size has shown to be the principal characteristic affecting the quality of the final printed structure [45]. The material reactivity can be affected by powder-binder wettability, binder spread behavior (it can affect the nucleation process), and droplet penetration into powder pores [46]. The great potential of using the drop-on-solid approach has been exploited in developing pharmaceutical formulations in different dosage forms such as tablets, implants, TeriFlasht, and devices (no-conventional pharmaceutical form). Controlled release of tablets using conventional pharmaceutical materials and drugs was reported by Katstra et al. [10]. In their formulation both top and bottom of the dosage form were formed by six placebo layers of the ammonio-methacrylic acid copolymer B Eudragit RLPO only. Eudragit RL is water insoluble but permeable polymer, independently of the pH. The middle region was formed by eight layers of Eudragit RLPO infused with chlorpheniramine maleate. Dissolution tests showed that release profiles could be modulated with polymer content [47]. This study confirmed the capability to modulate release properties of dosage forms by tuning the printing parameters, validating the precise dosage control and spatial positioning. However, as the drugs were incorporated into to tablets through binder solutions and the void space in the spread powder layers is limited, the amount of drugs that can be loaded into the tablets through binder solutions is also limited. Yu et al. reported an alternative to overcome this limitation of amount of incorporated drug into to tablets by using acetaminophen, HPMC E50, and PVP K30 as a drug models [46]. The drugs were incorporated inside the powder bed and formed tablets with a binder solution containing release-retardation materials, as an alternative of dispersing the drug binder solution. Both top and bottom of the tablets were made of ethylcellulose to produce inert layers that were impermeable to both water and drug diffusion (Fig. 6.7). The resulting tablets presented zero-order drug release and 98% of the drug could be released from the ethylcellulose gradient tablets in 12 hours [48]. The first 3D printed tablets commercialized was Spritam, which is an antiepileptic drug in a rapidly dispersing dosage form of levetiracetam (it can be totally dispersed in less than 5 seconds) and it was developed by Aprecia Pharmaceuticals using the ZipDose technology. Owing to the fast dispersion, Spritam is highly

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Figure 6.7 A schematic diagram of the tablet with material gradients, (A) barrier layer; (B) drug-containing region; and (C) gradients of release-retardation material. Source: Reproduced with permission from B.K. Lee et al., Fabrication of drug-loaded polymer microparticles with arbitrary geometries using a piezoelectric inkjet printing system. Intl. J. Pharmaceut. 427 (2) (2012) 305
310, Elsevier.

recommended for people with swallowing difficulties [5]. The ZipDose printing approach creates a porous formulation that binds powders without compression enabling the delivery of high drug dose medications of up to 1000 mg in a single dose. In this approach a powder blend is deposited as a single layer and then an aqueous binding fluid is applied, interaction between the powder and liquid binds these materials together. These steps are repeated numerous times to produce the final tablets. Monkhouse and coworkers reported the development of subdermally implantable device using drop-on-solid printing for controlled delivery of the estrogen medication ethinyl estradiol [47]. In their study, three different implant prototypes were fabricated and PCL was used as the polymer matrix [49]. In the first prototype, a single channel of ethinyl estradiol was embedded in a matrix of PCL (Fig. 6.8). In the second prototype, ethinyl estradiol was homogenously distributed in the PCL matrix and in the third prototype ethinyl estradiol was positioned in a concentration gradient in the polymer matrix. It was shown that the implants in a rabbit model did not cause any inflammation response, infection, or irritation after 13 weeks. Ethinyl estradiol could be released from all three implants, in different patterns, achieving the desired therapeutic effects [49]. The examples highlighted in this section showed the great potential of using the drop-on-solid printing approach for DDSs. Advantages of modular, easy, and repeatable process are offered by this methodology. One limitation related to the drop-on-solid methodology is the production of only macrostructures.

6.4.2 Drop-on-drop inkjet printing Drop-on-drop deposition is a direct-writing printing methodology that allows fabrication of drug carriers with different geometries, in microscopic dimensions, and high drug loadings [50]. Drop-on-drop inkjet printers usually consists of a cartridge, a printer head with multinozzles, a print head driver, and a stage plate with X
Y

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Figure 6.8 Schematic illustration of the single-channel design prototype ethinyl estradiol implant, in which the drug is incorporated in a single-channel drug loading surrounded by layers of poly-ε-caprolactone. Source: Reproduced with permission from M. Kuang, L. Wang, Y. Song, Controllable printing droplets for high-resolution patterns, Adv. Mater. 26 (40) (2014) 6950
6958, Springer Nature.

coordination, all coupled to a computer (Fig. 6.9) [41,48]. In this configuration the piezoelectric heads are commonly used, which allows the employment of different solvents. Hence, the nozzles are activated by the velocity waves to produce a droplet. The produced droplets are in the range of picolitres and size between 18 and 50 μm [49]. Parameters such as polymer inkjet physical properties and concentration are crucial to ensure systematic spreading of droplets. The choice of the solvent also has impact on the vapor pressure of the system and the solubility of the incorporated material [51]. Lee et al. demonstrated the potential to use drop-on-drop inkjet printing for the fabrication of paclitaxel-loaded polymer particles in micrometer dimensions with well-defined and controlled shapes [48]. After the optimization of physical properties of the ink (solubility, volatility, viscosity, and surface tension), paclitaxelloaded PLGA microparticles were obtained in different geometries, including grids, circles, rings, and honeycombs. A biphasic release profile was obtained, with an initial burst due to diffusion in the first day and the drug continued to be slowly released during six days due to degradation of the PLGA matrix. The drug release rate presented to be dependent on the microparticle geometry, principally the

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Figure 6.9 Schematic representation of a drop-on-drop inkjet printer. Source: Reproduced with permission from B.K. Lee et al., Fabrication of drug-loaded polymer microparticles with arbitrary geometries using a piezoelectric inkjet printing system, Intl. J. Pharmaceut. 427 (2) (2012) 305
310, Elsevier.

surface area. In addition, the released paclitaxel presented inhibition of the growth of HeLa cells in in vitro tests [48]. Although well-defined shapes of different sizes of drug-loaded structures can be achieved by using the drop-on-drop methodology, the related therapeutic and physicochemical properties are still to be investigated.

6.5

Summary and future perspectives

This chapter highlighted power and versatility of 3D printing for developing DDSs using the following techniques: direct writing, FDM, and inkjet printing. In many instances, harnessing this versatility has meant development of DDSs with improved performance with respect to critical parameters such as therapeutic dosing and delivery kinetics. The literature for direct writing using pressure-assisted systems shows the great potential to deposit hydrogels for tissue engineering, which presents high biocompatibility. FDM showed to be a potent tool for DDSs that needs to be implanted, and one of the main strategies is to incorporate thermoresponsive polymers that are suitable for biomedical applications. Inkjet printing is so far from the most successful case of a 3D printing technique used for DDSs, being the technique used to produce Spritam. It is clear that 3D printing as a manufacturing tool is playing a key role in pharmaceutical industry in providing new routes to the development of improved formulations, drug combinations, drug stability, and modes of delivery. In fact, drug delivery 3D printing systems provide the pharmaceutical industry the ability to develop new formulations that may give new life to drugs that have proven difficult to administer through convention systemic delivery routes. The diverse modes of 3D printing described in this chapter, in addition to

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the rapid improvements continually being reported, provide a vast toolbox of fabrication approaches to ensure that current limitation of drug delivery may overcome in the near future.

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202.

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Sepehr Talebian and Javad Foroughi Intelligent Polymer Research Institute, University of Wollongong, Wollongong, NSW, Australia

7.1

Introduction

Controlled drug delivery systems (DDSs), which are intended to deliver drugs at predetermined rates for predefined periods of time, have been used to overcome the shortcomings of conventional drug formulations. In fact, it would be most desirable if the drugs could be administered in a manner that precisely matches physiological needs at proper times (temporal modulation) and/or at the proper site (site-specific targeting). It would be highly beneficial if the active agents were delivered by a system that sensed the signal caused by disease, judged the magnitude of signal, and then acted to release the right amount of drug in response. Such a system would require coupling of the drug delivery rate with the physiological need by means of some feedback mechanism [1]. Notably, pharmacologically active agents are not inherently effective; their benefit is directly coupled to the manner by which they are administered into human. Administration route affects drug bioavailability, absorption, distribution, duration of therapeutic effect, and toxicity [2]. On the basis of such notion, drugs can either be administered systemically (orally or by intravenous injections) or they can be applied locally directly to the diseased site. While systemic administration of drugs has been the gold standard in treating numerous diseases, it still imposes certain limitations in treating complex diseases such as cancer. These limitations include low efficiency of delivering the drugs to the diseased site at therapeutic concentrations, and the ensuing toxicity to healthy tissues [3]. Of note, although intravenously administered nanoparticles (NPs) showed potential in treating certain diseases, however, they represent their own sets of drawbacks including poor loading efficiency, rapid clearance from blood, and inevitable cytotoxicity caused by nonspecific accumulation in nontarget tissues [4]. To overcome the limitations associated with systemic administration, the concept of local drug delivery has recently gained much attention and it is expected to enhance drug uptake and efficacy (Fig. 7.1) using the following tactics: (1) they are delivered locally (directly at the site of the disease) and therefore offer strategic and precise spatial control that significantly reduces the required dose and often the side/ off-target effects; (2) mitigates toxicity to healthy cells; (3) present temporal control over release profile of therapeutic agents (i.e., sequential release or pulsatile presentation) and maintaining the therapeutic concentrations over a longer duration of time; Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00007-X © 2020 Elsevier Ltd. All rights reserved.

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Figure 7.1 Schematic of systemic drug delivery versus local drug delivery for treatment of malignant tumors, including various types of biopolymeric implantable systems used for local drug delivery and different anticancer therapeutic modalities employed based on these implantable systems. 3D, Three-dimensional; DDS, drug delivery system. Reproduced with permission from S. Talebian, J. Foroughi, S.J. Wade, K.L. Vine, A. Dolatshahi-Pirouz, M. Mehrali, et al., Adv. Mater. 30 (2018), 1706665. Copyright 2018, Wiley-VCH.

and (4) can protect the loaded drugs from degradation or clearance until they are released [5
8]. To this end, implantable DDSs have emerged as a promising approach to effectively treat disease [9]. Implantable DDSs are denoted as implantable systems containing one or more therapeutic agents placed around or inside the diseased site to facilitate local delivery of these agents [10]. On that note, owing to astonishing properties of biopolymers, including biocompatibility, biodegradability, and flexibility in design and fabrication, they have entered the fray in instituting new generation of implantable DDSs that do not demand a surgical extraction after treatment and [11] even some of them including Gliadel wafers, oncogels, and LODER

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(local drug eluter) made it to clinical trials [12]. Biopolymers offer dazzling properties such as high porosity with controllable porous size, biodegradability, biocompatibility, versatility in chemistry as well as great design flexibility that make them befitting for various biomedical applications [13]. Indeed, biopolymers can be more specifically categorized into hydrogels (hydrophilic gels) and nonhydrogels (namely hydrophobic polymers). Hydrogels are being recognized as porous, soft materials possessing three-dimensional (3D) network structure composed of crosslinked hydrophilic polymer chains [14,15] and owing to their unique features such as water absorption/retention competency, tunable physical and chemical properties, and biocompatibility they have been established as promising carriers in drug delivery as well as scaffolds in tissue engineering [8,16,17]. On the other hand, hydrophobic biopolymers are categorized as synthetic biodegradable polymers, such as saturated aliphatic polyester (or any of their copolymers) and they all share the common characteristic of water repulsion. Additionally, owing to their water-repellent attribute they do not require subsequent crosslinking to be utilized in aqueous and psychological environments indispensable for medical applications [18]. These polymers offer certain advantages over hydrogels including flexibility in properties (including molecular weight, crystallinity, degradation rate, etc.) and befitting for most engineering fabrication methods (such as 3D printing and electrospinning) [19]. On account of such competencies, these hydrophobic biopolymers were also extensively exploited in the field of drug delivery [17,20,21]. These biopolymers not only act as a carrier for the drug but they could also exert control over its temporal presentation, for instance, smart polymers including pHand thermoresponsive polymers showed different drug release profile in different environmental conditions [22]. In addition, through employment of composite approach, nanomaterials responsive to light or magnetic field (Fig. 7.2) can be incorporated into these biopolymers to provide on-demand drug release (also known as pulsatile release) [23]. Nevertheless, biopolymeric implantable DDSs are commonly classified under two distinctly different categories, either being preformed ahead of implantation or formed in situ upon injection into the tumor site. Accordingly, each category requires certain class of polymers as well as fabrication methods to bring about the implant. For instance, preformed implants are often made from commercially available biopolymers such as FDA-approved thermoplastic polymers [including polycaprolactone (PCL), polylactic acid (PLA), and poly (lactic-co-glycolic) (PLGA)] as well as hydrogels (including alginate, gelatin, silk, and dextran) and they are most frequently fabricated through well-established fabrication methods such as casting, electrospinning, and 3D printing [24,25
28]. In contrast, in situ forming implants are directly injected into the tumor site and they are frequently made from two types of hydrogels, either thermo-/pH-sensitive biopolymers or in situ cross-linkable biopolymers [29
31]. In this chapter we will comprehensively study the implantable biopolymeric DDSs for drug delivery applications. We will start by defining the concept of active and passive delivery while providing some relevant examples. Subsequently, we will focus on application of hydrogels and hydrophobic polymers on delivery of bioactive molecules. Next, we will explore the application of novel

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Figure 7.2 Schematic showing different mechanisms of drug release from biopolymeric implantable drug delivery systems (DDSs). Reproduced with permission from S. Talebian, J. Foroughi, S.J. Wade, K.L. Vine, A. Dolatshahi-Pirouz, M. Mehrali, et al., Adv. Mater. 30 (2018), 1706665. Copyright 2018, Wiley-VCH.

implantable biopolymeric DDSs, including microdevices and transdermal patches, for delivery of biomolecules to target sites. Eventually, we will discuss some novel therapeutic modalities, based on implantable biopolymeric DDSs, which have revolutionized the field of biomedical engineering.

7.1.1 Active and passive drug delivery So far a variety of biopolymers have been harnessed to institute DDSs, yet, based on their release mechanisms we can entirely classify them into two categories: (1) stimuli-responsive and (2) nonstimuli-responsive systems. Stimuli-responsive (so called smart) systems show a rapid change in dimension or physical properties, instigated by either external or internal stimuli, which ultimately leads to convection-mediated delivery of the drug into the target site [32]. Of note, this rapid transformation is ensued from either a change in solubility, alteration of hydrophilic/hydrophobic state, change in degree of intermolecular associations or a reversible sol-to-gel transition [33].

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The externally stimulated systems rely upon externally exerted stimuli that are supplied by a stimuli-generating source, which ultimately endows pulsatile drug delivery [34]. Examples of such stimuli includes near infrared (NIR) light, magnetic fields, and ultrasound [35]. These systems are normally fabricated by incorporation of functional nanomaterials into thermosensitive biopolymers and their mechanism of action follows the pursuing sequences: (1) the nanomaterial starts to generate heat upon introduction of the external stimuli, (2) the generated heat will cause the biopolymeric matrix to undergo physical transformation (either expansion or contraction), (3) this will subsequently lead to rearrangement of pore structure within the biopolymeric matrix and resulting in release of the bioactive agent [5]. The type of nanomaterial used can determine the type of stimuli that will initiate the release, for instance, gold NPs show great sensitivity to NIR light while iron oxide NPs possess sensitivity toward electromagnetic waves. For instance, a thermosensitive hybrid hydrogel comprised methoxylpoly(ethylene glycol)-poly(ε-caprolactone)acryloyl, glycidylmethacrylated chitooligosaccharide (COS-GMA), N-isopropylacrylamide (NIPAm), and acrylamide (AAm), was synthesized (abbreviated as PCNA) and subsequently loaded with anticancer drug doxorubicin (DOX) and gold nanorods (GNRs) [36]. A NIR laser was used to initiate the release of DOX by utilizing photothermal effect of GNRs which caused contraction of the thermosensitive hydrogel. In a similar manner, a thermosensitive hydrogel made of hydrazide-functionalized poly(N-isopropylacrylamide) (PNIPAAm) and aldehydefunctionalized dextran loaded with superparamagnetic iron oxide NPs (SPION) showed to release the model drug (fluorescein-labeled dextran) in a pulsatile manner after exposure to alternating magnetic field [37]. Indeed, both mentioned systems rely on the generated heat (as a consequence of exposure to external stimuli) to bring about changes to the internal structure of polymer career, facilitating the drug release. Interestingly, in the context of cancer drug delivery this generated heat can also be used to shrink the tumors or prevent their growth, a phenomenon commonly known as photothermal therapy [3]. As it was mentioned at the beginning of this section, ultrasound can also be used to trigger the release of drugs from DDSs. For instance, Huebsch et al. fabricated a self-healing alginate hydrogel that was crosslinked with calcium sulfate [38]. The polysaccharide was crosslinked by divalent cations (Ca21) and it was observed that ultrasound was capable of disrupting ionic crosslinking to accelerate mitoxantrone release, while Ca21 ions in physiological fluids allowed crosslinks to reform after removal of the ultrasound. Interestingly, in vitro studies using MDA-MB-231 and MCF7 breast cancer cells revealed that ultrasound-mediated pulsatile release of mitoxantrone had significantly higher antitumor activity when compared with sustained release of drug from the same hydrogel. In contrast to externally stimulated systems, the internally stimulated systems function based on a trigger that is generated within the body, such as pH change or temperature fluctuation [39]. More often, the sensitivity of these systems to internal stimuli arises from susceptibility of biopolymer carrier to environmental factors such as soluble-to-nonsoluble transition upon heating above low critical solution temperature, or dissociation of specific bonds of polymeric network in acidic pH

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(e.g., acylhydrazone bonds, imine bonds) [40]. Owing to nature of such DDSs, they are not capable of pulsatile release of the drugs as factors such as pH or temperature do not fluctuate that often within human body. Consequently, the release profile from these systems includes an initial burst release of the drug molecules at the beginning followed by a gradual release in the following hours [41]. Thermoresponsive DDSs are normally made from thermosensitive polymers such PNIPAAm, poly(organophosphazene) (PPZ), and pluronic F127 [3]. The thermosensitive aspect of these polymers allows them to solidify upon implantation in living bodies which makes them ideal as injectable DDSs. For instance, in a noteworthy study, Ouyang et al. made an injectable thermosensitive hydrogel, from copolymer of aminated hyaluronic acid (HA) and PNIPAAm, which was loaded with small molecule gefitinib (an epidermal growth factor receptor inhibitor) [42]. In vitro [in Phosphate-buffered saline (PB)S at 37 C] this hydrogel was capable of controlled release of gefitinib over a span of 10 days. What is more, injection of gefitinib-encapsulated thermosensitive hydrogel into a puncture-induced rat intervertebral disk (IVD) degeneration model indicate that controlled release of gefitinib protects IVD from degeneration by suppressing cartilage matrix degradation while boosting type II collagen synthesis. pH-sensitive polymers are not as prevalent as thermosensitive polymers for fabrication of implantable DDSs as they need a secondary mechanism to ensure their in vivo stability. Besides, pH-sensitive polymers normally respond to lower pH values (acidic pH) to initiate the drug release and such conditions only exist in the gastrointestinal system which often demands a systemic approach of drug delivery [43]. In other words, these polymers can be used to institute implantable DDSs; however, their sensitivity toward pH-value does not instigate the drug release in vivo and that is why we will not study such systems in this chapter. Lastly, in nonstimuli-responsive systems the release is passive and it is solely controlled by diffusion/drug-carrier affinity/degradation of polymer or any combination of these [5,44]. Consequently, in these systems factors such as tortuosity of pores, steric interactions between drug and matrix, reversible chemical interaction of drug and matrix, and molecular weight of polymeric matrix are crucial in determining the release profile of therapeutic agent [5,41]. Consequently, these DDSs have been made from a variety of polymeric materials and owing to their simplistic mechanism of action they have been graciously embraced by the clinics [5]. Salient examples of these DDSs include (1) Taxus (Boston Scientific), a paclitaxel (PTX)eluting stent, which uses poly(styrene-b-isobutylene-b-styrene) coating to control drug release by retarding drug diffusion [45] and (2) INFUSE, a bone graft, consists of a rhBMP-2-soaked collagen sponge that stimulates bone formation in spinal reconstruction procedures, affinity, and diffusion retard the release of rhBMP-2 from the sponge [46].

7.1.2 Hydrogel drug delivery systems Of the many biopolymers currently utilized for drug delivery applications, hydrogels are among the most promising ones. Hydrogels are composed of polymeric

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networks that are capable of absorbing and retaining high amount of water [15,47]. Hydrogels are also tunable (both physically and chemically), are injectable, and have been used over the years for various drug delivery applications [23,48]. In fact, a variety of synthetic or natural biopolymers can be used to form hydrogels. This includes (1) synthetic polymers such as poly(ethylene glycol) (PEG), poloxamers or block copolymers of poly(ethylene oxide) and poly(propylene oxide), poly (acrylic acid) (PAA), PNIPAAm, and polypeptides; and (2) natural polymers such as chitosan, alginate, HA, collagen, gelatin, silk, and dextran [3]. Naturally derived polymers often possess adequate biocompatibility, while synthetic polymers may elicit negative reactions from the body [49]. Regardless, the gelling mechanism of hydrogels is either based on physical or chemical interactions [50]. Physically crosslinked hydrogels usually involve employment of either secondary bonds (such as ionic bonding, hydrogen bonding, and van der Waals interactions) or inherent phase transition behaviors [51]. On the other hand, chemically crosslinked hydrogels use covalent bonds or chelation or dynamic covalent bonds, to facilitate the gelation [52]. Nevertheless, factors such as hydrogel network’s mesh size, swelling ratio, and degradation rate will determine the release rate of drugs from these DDSs [53]. What is more, most of these features can be modified by adjusting the polymer concentration and degree of crosslinking [54]. As a consequence, hydrogels can be specifically engineered to be befitting for delivery of a variety of drugs and bioactive molecules. For instance, a hydrogel system composed of agar and gelatin crosslinked with genipin was used for delivery of vascular endothelial growth factor A165 (VEGF-A165) [55]. The results showed that VEGF-A165 loaded hydrogel (200 ngr/mL) was capable of sustained release of growth factor over a period of 65 days. In addition, it was shown that the released growth factor from the hydrogel was still bioactive as it caused phosphorylation of VEGFR-2, Erk-1/2, and Akt pathways in human umbilical vein endothelial cells. What is more, the released growth factor from the hydrogel displayed increased neurite out growth in dorsal root ganglia explants. In another study, a heparin functionalized four-arm PEG hydrogel was synthesized which was further processed postsynthesis into injectable microparticle aggregates [31]. Owing to presence of heparin in these hydrogel microparticles, they have shown high affinity toward DOX especially at higher concentrations, which lead to slow in vitro release of this drug in PBS. Moreover, injection of the DOX-loaded hydrogel microparticles close to breast cancer tumor in mice model has shown significant reduction in tumor burden as well as breast cancer metastasis when compared with bolus-treated animals. Hydrogels can also be designed to be responsive to environmental changes or external stimuli. For instance, triblock copolymers of PLGA acid and PEG (PLGAPEG-PLGA) are specifically attractive thermoresponsive systems on account of their biodegradability and acceptable safety profile and they showed to undergo sol
gel transition at physiological temperature (37 C) [56]. For instance, Yu et al. showed that incorporation of DOX into this thermogel did not interfere with its sol
gel transition (except at high concentration of DOX; 4 mg/mL) and after injection of this drug-loaded hydrogel in the vicinity of a sarcoma-tumor in mice, significant suppression of tumor growth as well as strong apoptosis of tumor cells were

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achieved [57]. Another example is PPZ, which is a hydrogel that undergoes a sol
gel transition in psychological temperature that makes it a suitable candidate for establishing DDSs. Al-Abd et al. loaded the PPZ thermogel with DOX for local delivery of this drug to human gastric tumor xenografts in mice [58]. Interestingly, it was observed that DOX-loaded gel (POL) had 40% and 90% sustained drug release over 5 weeks in vitro and in vivo, respectively. What is more, in vivo results suggested that POL had similar efficacy in suppressing tumor growth as DOXloaded solution (SOL); however, POL showed dramatically decreased systemic toxicity compared with SOL leading to higher survival rates of animals after 28 days of treatment. In a different study, Zhang et al. synthesized PPZ nanocapsules loaded with SPION that transformed to a hydrogel (SPION-NHs) upon exposure to body temperature [30]. The same group loaded the SPION-NHs hydrogel with tumor necrosis factor
related apoptosis-inducing ligand (TRAIL) (T/S-NHs) to synergistically abrogate tumors by multiple magnetic hyperthermia (MHT) and anticancer potential of TRAIL [59]. Accordingly, it was shown that chemical structure of PPZ had a direct effect on in vitro degradation and subsequent release of TRAIL and SPIONs from the hydrogel, as nonionic side group on PPZ (indicated as T/S-NHs1), leads to fastest degradation rate as well as fastest SPIONs and TRAIL release, while hydrogel made from PPZ with ionic side groups (indicated as T/S-NHs-2 and T/S-NHs-3) showed much slower degradation rate and consequently slower TRAIL and SPIONs release. Additionally, intratumoral injection of T/S-NHs in glioblastoma tumor xenograft in mice revealed that tumor growth was significantly inhibited in animals that received T/S-NHs-2 with two cycles of MHT as a result of combinational therapy. As it was mentioned early on in this chapter, hydrogels are interesting platforms as they can be engineered in a variety of ways to bring about peculiar properties. Self-healing hydrogels (capable of mending themselves upon damage) and thermodegradable hydrogels (will undergo degradation and/or chain dissociation after being exposed to heat) are among recent advances in the field of material science [5,60]. Of note, healing procedure in self-healable hydrogels is facilitated through implementation of either dynamic covalent bonds or noncovalent bonds, which allows reversible formation of hydrogel network [61]. Self-healing hydrogels have recently gained a foot hold in the field of drug delivery as stability of implants after implantation is of great importance particularly considering its direct effect on corresponding drug release profile. For example, Wang et al. developed an injectable self-healing hydrogel from a mixture of glycol chitosan, telechelic difunctional poly(ethylene glycol) (DF-PEG), and saline ions, where the selfhealing capability stemmed from Schiff base bond between amino groups of chitosan and benzaldehyde groups of DF-PEG [62]. Yang et al. used this hydrogel loaded with PTX and intratumoral injection of this DDS in liver cancer tumor xenograft (in mice) showed significant inhibition of tumor growth when compared with that in mice that received pluronic F127 hydrogel loaded with PTX [63]. This observation was assumed to be the result of unique self-healing ability of the hydrogel allowing it to rebuild as a whole after injection which prevented the leak of PTX and subsequently resulted in longer term antitumor efficacy. As it was

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mentioned above thermodegradable hydrogels are another class of newly established hydrogels. This feature can be further used to allow on-demand delivery of therapeutic agents via administration of heat. For instance, Hu et al. designed a thermoresponsive injectable hydrogel by crosslinking four-arm amine-terminated PEG with an azo-containing linker and the hydrogel showed rapid degradation above 44 C due to breakage of azo bonds [64]. Embedding this hydrogel with photothermal NPs (dendritic platinum
copper NPs) enabled on-demand and dosetunable release of entrapped DOX in vitro upon irradiation with NIR light. In vivo injection of this hydrogel under the dorsal skin of nude mice followed by NIR irradiation (once per day for 10 minutes) degraded the majority of hydrogel mass after 3 days of treatment. Finally, intratumoral injection of DOX-loaded gel in breast cancer tumor xenograft (in mice) followed by NIR irradiation led to significant regression of tumor volume as a result of fast release of DOX in the tumor site.

7.1.3 Thermoplastic drug delivery systems Biodegradable thermoplastic polymers offer a number of advantages over other biopolymers for developing DDSs. The key advantages include the ability to tailor mechanical properties and degradation kinetics to suit various applications. Synthetic polymers are also attractive because they can be fabricated into various shapes with desired pore sizes using a range of fabrication methods including casting, electrospinning, and 3D printing [3,65]. Some of the most common thermoplastic polymers used to institute DDSs are PLA, PLGA, and PCL [21]. All these polymers share the common characteristic of being hydrophobic and they have often been used to deliver hydrophobic drugs. Besides these polymers do not demand any crosslinking step to ensure their stability in biological environment. For instance, Wang et al. incorporated NPs of The tripeptide Arg-Gly-Asp (RGD)modified PEGlated polyamidoamine (PAMAM) dendrimer loaded with DOX (RGD-PPCD) into PLGA/PLA solution containing PEG (as drug release modifier) and then casted them into a cylindrical die [66]. In vitro release studies revealed that the ratio of PLGA/PLA had an impact on the release of DOX, with an increase in PLA content decreasing the drug release rate. It was also observed that different amounts of PEG can influence the DOX release, as an increase in PEG lead to significant increase in release rate of DOX. Furthermore, in vivo results indicated that Poly(lactic-co-glycolic acid) (PLG)/PLA scaffolds containing RGD-PPCD NPs could significantly reduce glioma tumor size in mice model as opposed to their counterparts including PPCD implant, DOX implants, and blank implants. These findings were assumed to be result of better penetration of RGD-PPCD NPs into the tumor which highlighted beneficial role of RGD sequence in tumor retention of NPs. PLGA was also utilized to fabricate a biopolymeric cylindrical implant, also known as local drug eluter/LODER, that contained an siG12D [an siRNA against the mutated V-Ki-ras2 Kirsten rat sarcoma viral oncogene homolog (KRAS) oncogene] for growth inhibition of pancreatic tumors in a mouse model. [67]. It was found that encapsulated siG12D in LODER was active and stable for 155 days

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in vivo. Further, it was shown that these implants were capable of impeding the tumor growth and prolonging mouse survival time. As it was highlighted earlier thermoplastic polymers can be applied to different fabrication techniques to make DDSs. For instance, Ding et al. loaded the poly-D, L-lactide nanofibers with docetaxel (DTX) to prevent breast cancer reoccurrence in mouse model [28]. It was observed that nanofibrous membranes were capable of sustained release of DTX in vitro (in PBS) over a period of 24 days. What is more, it was shown that animals treated with drug-loaded membranes had significant decrease in locoregional reoccurrence after primary tumor resection (16.7%) when compared with systemic administered DTX (75.0%) or local administered DTX (77.8%). Additionally, these drug-loaded membranes showed minimal signs of inflammation in the surrounding tissue indicating a high biocompatibility. In a different study, 5-fluorouracil (5-FU)-loaded Poly-L-lactic Acid (PLLA) nanofibrous membranes were developed for suppressing colorectal cancer in xenografted mice [68]. In vivo it was shown that these membranes were more capable of suppressing tumor growth than an intraperitoneal injection of 5-FU [at median lethal dose (LD50) concentration] due to prolonged and continuous release of 5-FU from the membranes. 3D printing is another fabrication method that has recently been applied to thermoplastic biopolymers to establish implantable DDSs. For example, most recently extrusion printing was used to fabricate a 3D patch made from a mixture of PLGA (lactide:glycolide 5 85:15) and PCL loaded with 5-FU for growth suppression of pancreatic cancer [25]. The patches were printed with three different pore shapes and geometries (latticed, slant, and triangular) in different thicknesses and it was found that these features can greatly affect the drug release profile by altering the surface area:volume ratio of the structure. Implantation of drug-loaded patches (P100;100 mg 5-FU and P150;150 mg 5-FU) to the bottom of pancreatic cancer tumor in mice, showed substantial decrease in tumor size when compared with nondrug loaded patch (P0) groups. In another study, Sun and his colleagues fabricated a PCL scaffold using a Fused Deposition Modeling (FDM) printer and coated the structure with a mixture of chitosan, chitosan-modified montmorillonite clay, and β-tricalcium phosphate which was subsequently coated with DOX solution (DESCLAYMR_DOX) to inhibit growth of breast cancer tumor in mice model [26]. These drug-eluting implants showed a significant burst release of DOX in the first 24 hours that was followed by 4 weeks of sustained release. Further, subcutaneous implantation of DESCLAYMR_DOX in mice showed prolonged presence of DOX to a much larger extent at the treatment site which resulted in higher tumor growth inhibition when compared to subcutaneous injection of DOX (INJECTION_DOX). Also compared with INJECTION_DOX, DESCLAYMR_DOX showed decreased multiorgan metastasis as well as cardiotoxicity due to local delivery of DOX. Unlike hydrogels, thermoplastic biopolymers cannot be responsive to environmental changes (such as pH and temperature); however, they can still bestowed responsivity to external stimuli through incorporation of functional nanomaterials such as gold, iron oxide, and up-conversion particles. For instance, Chen and his coworkers made nanofibers from combination of PCL and gelatin (PG) that

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contained core-shell NPs of Cu9S5-mesoporous SiO2 (Cu9S5@mSiO2) loaded with DOX and used them for synergistic chemo- and photothermal therapy of hepatoma tumors in mice [27]. In vitro DOX release in PBS (phosphate-based buffer) revealed that these fibrous membranes possessed pH-responsive release owing to intrinsic properties of Cu9S5@mSiO2 NPs, also it was shown that after 5 minutes laser irradiation the temperature of Cu9S5@mSiO2 PG fibers dramatically increased from 21 C to 54.6 C. Additionally, in vivo results showed that DOX-loaded Cu9S5@mSiO2 PG composite fibers under laser irradiation had a more efficient tumor suppression effect once compared against single photothermal therapy of tumors by Cu9S5@mSiO2 PG fibers or with single chemotherapy by DOX-loaded Cu9S5@mSiO2 PG fibers. In a different study, Chen et al. incorporated the iron(III) oxide (Fe3O4) NPs into a PLGA solution for MHT regression of tumors [69]. Intratumoral injection of this solution in breast cancer tumor xenograft in mice followed by only single magnetic ablation showed complete disappearance of tumor after 3 days that did not show reoccurrence even after 1 month. Same group also developed another injectable DDS based on PLGA, which MoS2 nanosheets and DOX were incorporated into the implant (PMD) for synergistic photothermal and chemotherapy of tumors [70]. Therefore this implant showed controlled release of DOX in pH- and NIR-responsive manner. Furthermore, intratumoral injection of this composite in breast cancer tumor xenograft in mice followed by only one time NIR irradiation (5 minutes) led to disappearance of tumor after 7 days and no tumor recurrence was observed within 2 months from the treatment. Lastly, mice treated with PMD 1 NIR had 100% of survival rate over a period of 50 days, which indicated high efficiency of this local synergistic treatment.

7.1.4 Microdevices delivery systems Recent effort to institute miniaturized drug delivery devices has ultimately led to development of integrated systems that incorporated device technology with therapeutic molecules to create implantable devices capable of disease treatment in situ [71,72]. On the basis of the design of these microdevices, they could present different options such as allowing continuous or intermittent delivery, operating for short or long periods [71]. Similar to traditional DDSs, these devices can either exert passive delivery of a drug or they can permit real-time control of drug dosage by means of external stimulus [73]. In the following section, some of the most salient examples of polymeric microdevices for cancer therapy are reviewed. Jonas et al. recently engineered a small cylindrical device (via micromachining from a medical grade acetyl resin blocks) designed to release a wide range of anticancer drugs from 16 discrete reservoirs on its perimeter [74]. The device was directly implanted into tumors of different types using a biopsy procedure and left in there for 24 hours for quick, parallel investigation of drug sensitivity in in vivo tumors. Next, the device and a small region of tissue were removed using a coring biopsy needle, to determine the antineoplastic effect of each treatment. In order to modify the release rate of each drug in the reservoir, three different techniques were used: (1) altering the size of reservoir opening, (2) embedding drugs into

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polymeric matrix (PEG), and (3) using a hydrophilic expansive polymer that would expand upon contact with fluid in the tumor in order to deliver hydrophobic drugs. Consequently, It was shown that the device was capable of delivering individual or combination (codelivered from one reservoir or delivered separately from two adjacent reservoirs) of drugs into confined regions of tumor tissue. The authors further evaluated the pharmacokinetics of DOX release from the device in human breast carcinoma tumor xenografts and it was shown that the drug released from device can diffuse 200
300 μm into tumor tissue. In addition, this device was found to be useful in combatting tumor heterogeneity (a major challenge in all diagnostics) by implanting a device, with 16 identically loaded reservoirs (with DOX) on different locations of the device, in the nonnecrotic periphery of skin cancer tumor xenografts. Dosage of drug and frequency of its administration play a crucial role in the treatment of chronic diseases such cancer [75]. Consequently, advanced techniques that allow on-demand, precise, and controlled local delivery of drugs are highly needed for individualized disease treatment [76]. Owing to digital capabilities of microelectromechanical systems (MEMS) they allow for complex temporal profiles, while the manufacturing techniques used in microelectronic industry can lead to greater device uniformity and reproducibility than currently available to pharmaceutical and biomedical industries [77,78]. On account of such desirable features, MEMS devices have been recently explored for anticancer drug delivery purposes and have shown some promising results [78
80]. However, the major issues with MEMS devices are related to biocompatibility of materials used to institute the device as well as constant need for a sustained power supply. Consequently, in a notable effort Chin et al. used photolithography to develop a biocompatible and biodegradable MEMS microdevice from PEG diacrylate, which contained moving parts that can be controlled by an external magnet [80]. Specifically, two different designs of this device, single-gear and geneva drive, were tested in vivo on a mouse model. Single-gear device contained a multireservoir gear, where one of them was loaded with iron oxide NPs and the others with drugs. On the other hand, geneva drive had two engaged gears, a driving gear that was doped with iron oxide NPs and a driven gear with six drug-loaded reservoirs. The concept of release for both designs were the same, where movement of an external magnet caused motion in the gears (owing to presence of NPs) and the drug in each reservoir was released once the aperture on top it aligned with another aperture on topmost layer of the assembled device. Subcutaneous implantation of the geneva drive device (in a mouse model) loaded with two different fluorescent dyes showed successful movement of gears in vivo which led to controlled release of the dyes. Furthermore, the DOX-loaded single-gear device (loaded with 10% of a single standard chemotherapy dose) was implanted adjacent to a human osteosarcoma xenograft (in a mouse model), which was actuated once every 2 days, over a period of 10 days. Consequently, this showed the greatest tumor volume reductions with the presence of tumor necrosis, and also caused the lowest levels of cardiotoxicity compared with all other treatment groups (including systemic administration of various concentrations of DOX with different frequencies). Lastly, analysis of the excised

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device and surrounding tissue showed no signs of chronic inflammation and revealed normal wound healing indicating good biocompatibility of this device.

7.1.5 Transdermal patches delivery systems As it was discussed in previous chapters local delivery of drugs holds a great advantage over systemic delivery, and various device technologies were implemented to fabricate the corresponding DDSs. One of the well-recognized devices is transdermal patches that allow local delivery of drugs through skin tissue. Specifically, local delivery of drugs via skin could be ideal for treating diseases that happen in the skin tissue, including melanoma (a form of skin cancer). In spite of that, transdermal delivery is extremely hindered by the inability of the large portion of drugs to cross skin at therapeutic rates owing to the immense barrier imposed by skin’s outer stratum corneum layer [81]. To increase the skin permeability, a range of methods have been utilized, this includes application of chemical/lipid enhancers [82], iontophoresis and electroporation [83], and ultrasound or photoacoustic effects to generate pressure waves [84]. While all the abovementioned methods are capable of disrupting the stratum corneum structure, they can only create nanometric holes that are presumed to be large enough to allow for transport of small drug molecules and perhaps macromolecules. An alternative approach involves creating larger transport pathways of microns dimensions using arrays of microscopic needles. These generated pathways are normally orders of magnitude larger than the drug molecule dimensions and therefore they can easily allow transport of macromolecules as well as supramolecular complexes and microparticles [85]. Worth noting that these generated holes in the skin are likely to be safe, given that normal injuries to the skin tissue or hypodermic needles create larger holes in the skin [86]. Regardless, the microscopic needles can be made from a variety of materials including polymer, metals, and ceramics; however, herein we will solely focus on biocompatible and biodegradable polymeric microneedles as they are gaining more attention in this field [87]. Generally, these microneedles can be used for a number of applications including but not limited to insulin delivery for diabetic patients, antineoplastic agent delivery for cancer treatment (or in combination with other cancer therapeutic modalities including photothermal therapy and photodynamic therapy), and gene delivery for various disease treatments [81]. For instance, in a noteworthy study, Di and his colleagues developed a stretch-triggered drug delivery platform from microneedles assembled on top of an elastomer film containing therapeutic depots [88]. The release of drug from the microdepot is promoted by applying a tensile strain to the elastomer film; this is due to the enlarged surface area for diffusion and Poisson’s ratio-induced compression on the microdepot. The microneedles were made from a mixture of methacrylate-modified HA and a crosslinking agent (N,N-methylenebis(acrylamide)), and they were specifically designed to allow transcutaneous delivery of drugs encapsulated in the microdepots. This drug-loaded wearable device was shown capable of delivering different drugs (including DOX, ciprofloxacin, and insulin) on-demand by simply stretching the device (that would

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initiate the drug release) and specifically insulin-loaded devices were successful in maintaining blood glucose level in mouse animal models. Similar to other DDSs, microneedles can take on active delivery of drugs (through employment of functional nanomaterials) as well as other therapeutic modalities to treat diseases [89]. In a recent study, Chen et al. fabricated a microneedle system where the supporting array patch was made of a mixture of polyvinyl alcohol and polyvinylpyrrolidone and the microneedle tips were made of PCL that was loaded with the anticancer drug (DOX) and photosensitive nanomaterial (lanthanum hexaboride/LaB6) [90]. The supporting array patch offers mechanical strength for completely inserting the microneedles into the skin and it dissolves upon penetrating into the skin. Presence of LaB6 in the microneedle tips allowed photothermal therapy using NIR irradiation which also caused melting of the PCL matrix that further facilitated on-demand release of encapsulated DOX. What is more, single application of microneedles and three cycles of laser treatment led to complete eradication of breast cancer tumors (4T1 cells) in mouse models, due to synergistic application of photothermal therapy and chemotherapeutic drug. The microneedles patches were also shown to be capable of delivering specific genes to target sites to stop the progression of various diseases [91]. As an example, Wang et al. fabricated microneedle patch from acrylate-modified HA, where the microneedles were loaded with dextran NPs encapsulated with anti-PD-1 antibodies (aPD1), glucose oxidase (GOx), and catalase (CAT) [92]. The aPD1 that target the inhibitory receptors have demonstrated significant antitumor activity in phase II and III clinical trials of advanced melanoma. GOx (with the assistance of CAT) is used to convert blood glucose to gluconic acid in the presence of oxygen (O2). The generated gluconic acid promoted the gradual self-dissociation of dextran NPs and subsequently led to sustained release of aPD1 over a 3-day period. Furthermore, single administration of microneedle patches induced robust immune responses in a B16F10 mouse melanoma model, which further led to remarkable antitumor efficacy and it was even observed that 40% of mice still survived 40 days after treatment compared with all other control groups where none of the mice survived. This observation was thought to be a result of sustained release of aPD1 from the microneedles. At last but not least, the microneedle patches can be combined with device technologies to institute sensing devices that are also capable of transdermal drug delivery [93]. Most recently in a notable study, Lee et al. developed a wearable sweat-based glucose monitoring system that was equipped with microneedle-based transdermal drug delivery module [94]. The diabetes drugs (metformin or chlorpropamide) were loaded into two different temperature-responsive phase change NPs, and they were subsequently embedded in HA hydrogel microneedles that were additionally coated with phase change materials. The device would allow multimodal glucose sensing (based on collection of sweat on the skin) and if the glucose level was higher than the standard value then it would initiate heat generation via three channels of heaters that would eventually lead to melting of phase change materials and release of drugs through microneedle patches. Interestingly, each drug was encapsulated in a different phase change material NP with varying

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melting points (38 C and 43 C) to allow release of certain drug just by controlling the temperature. What is more, animal studies using diabetic mice, type 2 diabetes mellitus model, showed that this device was capable of closely monitoring sweat glucose level (that was shown to have a direct correlation with blood sugar level) and concomitantly controlling the blood glucose level via programmed release of drugs based on recorded sweat glucose levels.

7.1.6 Emerging therapeutic methods based on implantable drug delivery systems Considering the versatility of implantable biopolymeric DDSs, it allows them to be loaded with novel anticancer therapeutic agents to deliver high concentrations of them to the disease site locally. Therefore in this section we will briefly discuss a recently emerging therapeutic modality to abrogate primary tumors and subsequently bring examples of implantable biopolymer-based DDSs that were used to deliver these agents to the tumor microenvironment.

7.1.6.1 Cancer immunotherapy Cancer immunotherapy involves any intervention that aids the immune system to eradicate a malignancy [95]. Immunotherapy modalities can function via two different mechanisms, either by stimulation of the intrinsic immune system or by unleashing extrinsic immune system components such as specific proteins to be recognized and targeted by immune cells [7]. However, multiple factors have hampered the success of these therapies in clinical cases including associated toxicity, difficulties involved with defining the optimal dosage and schedule of immunotherapy, and at last lack of efficiency in patients with a large tumor burden [7,96]. Biomaterial-based systems, however, are capable of addressing these issues owing to their intriguing properties such as protecting the bioactive agents or cells, exerting control over their spatiotemporal release and competency to deliver more than one agent from a single platform [97]. Consequently, a wide range of biomaterials have been used for delivery of immunomodulatory factors and here we will only highlight the most recently developed implantable biopolymeric DDSs used for this purpose. In the context of cancer immunotherapy, these implantable DDSs can be used in two well-distinguished ways, either for delivering immunomodulatory agents that override checkpoints in the cancer
immunity cycle or to deliver cells for adoptive cell transfer (ACT) into cancerous tissue to enhance their survival and proliferation [98].

7.1.6.2 Delivery of immunomodulatory factors Many types of immunomodulatory agents, antibodies, antigens, adjuvants, cytokines, and checkpoint inhibitors have been loaded into implantable biopolymeric DDSs to provide immune-protection against cancer [99]. For example, Bencherif and his colleagues developed a methacrylated alginate sponge-like cryogel

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containing covalently conjugated RGD peptide which was encapsulated with granulocyte-macrophage colony-stimulating factor, cytosine-phosphodiester guanine oligodeoxynucleotide, and irradiated tumor cells (B16-F10 mouse skin melanoma cells) [100]. Subcutaneous injection of this cryogel vaccine into a mouse model brought about a durable and specific antitumor T-cell response which led to 80% survival rate in a B16-F10 tumor model. Moreover, it was shown that cryogel vaccines bestow long-lasting protective immunity as 100% of vaccinated mice that survived the initial tumor challenge survived the subsequent second challenge.

7.1.6.3 Delivery of cells for adoptive cell transfer ACT offers great advantages over other methods of cancer immunotherapies, which heavily rely on in vivo recruitment of a sufficient number of antitumor T cells with required functions to instigate cancer regression [101]. Usually in ACT a large pool of antitumor lymphocyte, capable of recognizing tumor cells, are grown and proliferated in vitro and subsequently administered to the patient with the aim of providing a tumor microenvironment that supports antitumor immunity [102]. Nevertheless, recent years have witnessed emergence of biopolymeric scaffolds as platforms for improving function of these adoptive cells and enabling their local delivery to the tumor site [95,103]. For example, Stephan et al. developed a bioactive polymeric carrier designed for expanding and delivering tumor-reactive T cells for treating inoperable or incompletely removed tumors in a mouse breast cancer resection model [104]. The microporous scaffolds were made of alginate functionalized with collagen-mimetic peptide (used to facilitate migration of T cells through the scaffolds by means of α2β1 collagen receptor) using ionic crosslinking. Furthermore, porous silica microparticles containing anti-CD3, anti-CD28, and antiCD137 antibodies were incorporated into the scaffold to create stimulatory microenvironment for cytotoxic T-lymphocyte proliferation. For in vivo studies breast cancer tumor xenografts (in mouse model) were intentionally partially resected to mimic reoccurrence postresection. Ultimately, implantation of these scaffolds at the resection site led to higher T-cell proliferation and consequently none of the treated mice showed tumor reoccurrence after a period of 80 days, in contradiction to the control group which all died due to cancer relapse.

7.1.7 Summary Biopolymeric DDSs have shown immense potential in treating various diseases by allowing spatiotemporal control over release of bioactive agents. These DDSs can be made from a wide range of biopolymers and they can endow passive or active delivery of drugs. What is more, these systems do not necessarily have to be implanted at the diseased site using invasive surgery, for instance, that is the case with injectable DDSs where they could easily be injected in the target site to avoid further complications that arises from surgeries. These systems even showed competency in pairing up with device technologies to bring about precise control over temporal presentation of drugs. What is more interesting is that these systems do

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not always have to be implanted deep within the human body, but they could also be applied to treat diseases through transdermal drug delivery as it is the case for microneedle patches. Owing to favorable properties of biopolymer DDSs, they have also shown success in delivering new therapeutic modalities to diseased sites (such as immune cells) to specifically target the impaired tissues and avoid unnecessary damage to healthy tissues. In spite of all great attributes of these biopolymer DDSs, they still fail in completely treating such complicated diseases such cancer and that is why these systems have to be administered along with nanoparticulate systems to treat cancer both locally and systemically. Nevertheless, in order for biopolymeric implantable DDSs to reach a broad clinical implementation, scientists have to discover simple polymer chemistries that are easy to follow while offering sufficient safety. Lastly, continual communication between scientists and clinicians is crucial for establishing new treatment standards.

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Polymers and hydrogels to deter drug abuse

8

Samaneh Alaei, Niloofar Babanejad, Rand Ahmad and Hamid Omidian Department of Pharmaceutical Sciences, College of Pharmacy, Nova Southeastern University, Fort Lauderdale, FL, United States

8.1

Introduction

8.1.1 Prescription drug abuse The abuse of prescription drugs has turned into a crisis over the past decades. There are many drugs prone to abuse and the common ground for all is the fact that they can alter the abuser’s thinking and judgment. According to the National Institute on Drug Abuse (NIDA), besides opioids, many other drug classes such as stimulants, sedatives, and hypnotics like benzodiazepines and other central nervous system (CNS) suppressants are also subject to abuse [1,2]. According to the National Survey on Drug Use and Health, approximately 6 million out of 30.5 million Americans aged 12 years or older who were illicit drug users, have abused a prescription medicine in 1 month in 2017. The prescription pain relievers (mostly opioids) are the second most common abused drugs after marijuana, with 3.2 million people have misused a prescription pain reliever in only 1 month. This is higher than what has been reported for cocaine abuse. Furthermore, with 1.8 and 1.7 million Americans, respectively, misused the stimulants and hypnotics in 2017, these two classes of drugs rank fourth and fifth following cocaine abuse [3]. Besides the significant socioeconomic impact of drug abuse (e.g., risk of hospitalization, overdose, and addiction) for opioids and CNS suppressants [4], the abuse of the CNS stimulants such as Ritalin is also surging [5]. Fig. 8.1 shows the number of illicit drug users in the Unites States in 2017 just in 1 month. The NIDA has identified three classes of prescription medicines that have a high potential for abuse [6]. These include opioids, CNS depressants (sedatives and hypnotics), and stimulants. Fig. 8.2 shows the prescription drugs with high potential for abuse that have been approved in the United States.

8.1.2 Routes of abuse The activation of brain reward pathway is responsible for the euphoric effect of the prescription drugs [7]. The intense euphoric effect or the rapid “high” is achieved when the amount of drugs such as opioids in the blood reaches to a very high concentration in a short period of time. To achieve this, the abusers manipulate the Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00008-1 © 2020 Elsevier Ltd. All rights reserved.

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Figure 8.1 Illicit drug users in the United States in 2017.

Prescription opioids

CNS depressants

Oral formulations

Hydromorphone, morphine sulfate, hydrocodone, oxycodone, tapentadol, tramadol, methadone, buprenorphine, codein, meperidine, levorphanol

Parenteral formulations

Meperidine, morphine, buprenorphine, butorphanol, nalbuphine, sufentanil, pentazocin, remifentanil

Transdermal formulations

Fentanyl, buprenorphine

Barbiturates

Mephobarbital, phenobarbital, pentobarbital sodium

Benzodiazepines

Diazepam, clonazepam, alprazolam, triazolam, estazolam

Non benzodiazepine sedative hypnotics

Zolpidem, eszopiclone, zaleplon

Common drugs of abuse Dextroamphetamine

Dextroamphetamine/amphetamine combination products Prescription stimulants Methylphenidate

Pseudoephedrine

Over-the-counter medications

Cough & cold medications, loperamide

Figure 8.2 Common prescription drugs susceptible to abuse. CNS, central nervous system.

Polymers and hydrogels to deter drug abuse

187

Figure 8.3 Common methods of prescription drug abuse.

drug, or they overdose the original opioid medication. The most common routes of drug abuse include oral route (ingesting or chewing the medication usually in doses higher than prescribed), parenteral route [intravenous (IV), where the abusers introduce the drug directly into their bloodstream], and inhalation route (smoking and snorting the substance) [8]. To achieve a faster drug release and greater absorption, abusers manipulate or modify the original formulation to change the intended release profile or the delivery route. This is called tampering and include crushing or grinding the dosage form into powders or fine particles suitable for nasal insufflation, dissolving the active pharmaceutical ingredient in various solvents and extracting it for injection, coadministration of the drug product with alcohol or other psychoactive substances with the similar effect, and simple oral administration of multiple units of the drug product (Fig. 8.3) [8]. Among the various routes of abuse, oral ingestion of the prescription drugs is the most common route for opioids and the CNS suppressants [9,10] followed by inhalation and injection of the drug products. Although the pattern can be different for other classes, the IV injection remains the most dangerous route of abuse. Nevertheless, the pattern of drug abuse can be different for immediate and extended-release drug products [10].

8.2

Abuse deterrent formulations

To combat prescription opioid epidemic, along with regulatory measures, abusedeterrent formulations (ADFs) have been developed that utilize various technologies to resist manipulation and subsequent abuse by alternate routes of administration. Meanwhile, the ADF is expected to provide a safe and effective treatment when the

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Figure 8.4 Approaches to develop abuse-deterrent formulations.

drug is taken as prescribed. An ADF should ideally withstand most of the abuse methods such as particle size reduction, alcohol coingestion, and drug extraction. To design an ADF, it is also important to understand the abuse potential of the targeted drug and its common methods of abuse. For instance, if a drug is commonly abused by nasal insufflation, the primary focus should then be on improving the mechanical and crush resistance properties to impede particle size reduction rather than focusing on alcohol coingestion or IV abuse. As shown in Fig. 8.4, the ADFs are classified into several categories based on the mechanisms by which they deter the drug abuse. The most common approaches used in designing ADF products include the physical barriers that help to resist mechanical manipulations (e.g., crushing, chewing, cutting, or grinding), chemical barriers (e.g., instant gelling of drug solution intended for IV abuse) that provide resistance to extraction, and combining agonist and antagonist in a single formulation. Other strategies include aversive agents (e.g., nasal irritants or bittering agents) and delivery systems capable of maintaining the therapeutic drug release even after the product is manipulated. Currently there are 17 ADF products approved by the FDA that utilize the aforementioned technologies to resist manipulation and abuse. Among all, the physical and chemical barriers are the most widely used, as shown in Table 8.1. A simple patent survey clearly shows that polymers have played a major role in developing ADFs, and this is due to the fact that polymers can provide unique properties that other materials simply cannot. These include viscosity, hardness, toughness, water

Table 8.1 Deterrent technologies utilized in the FDA-approved abuse-deterrent formulations. Deterring class

Deterring technology

Mechanism

Manufacturing process

Main deterring agent(s)

FDA-approved product name

Dosage form

Dose strength

Manufacturer

Physical and Chemical Barrier

DETERx

Crush and extraction resistant controlledrelease microspheres

Hot melt extrusion

Xtampza ER (oxycodone)

Extended-release capsules

9
36 mg

Collegium Pharm Inc

INTAC

Crush and extraction resistant extendedrelease matrix

Hot melt extrusion

Beeswax, carnauba wax, myristic acid PEO

Extended-release tablets Extended-release tablets

50
250 mg 5
40 mg

Collegium Pharm Inc Endo Pharms

Crush and extraction resistant extendedrelease matrix Crush and extraction resistant controlledrelease matrix A crush resistant semipermeable membrane and an extraction resistant core Crush and extraction resistant extendedrelease tablet

Injection molding

PEO

Nucynta ER (tapentadol HCl) Opana ER (oxymorphone HCl) Arymo ER (morphine sulfate)

Extended-release tablets

15
60 mg

Egalet

Granules coating and compression Wet granulation and coating

Glyceryl behenate, HPMC, EC Cellulose acetate, PEO

Vantrela ER (hydrocodone bitartrate) Exalgo (hydromorphone HCl)

Extended-release tablets

15
90 mg

Teva

Extended-release tablets

8
32 mg

SpecGx LLC (Mallinckrodt)

Thermal compression

PEO

Hysingla ER (hydrocodone bitartrate) OxyContin (oxycodone HCl) Morphabond ER (morphine sulfate) Roxybond (oxycodone HCl)

Extended-release tablets

20
120 mg

Purdue Pharma

Extended-release tablets Extended-release tablets Immediate-release Tablets

10
80 mg

Guardian

OraGuard

OROS

RESISTEC

SentryBond

Crush and extraction resistant formulation

Direct compression and coating

Acrylate copolymer, alginic acid

15
100 mg

Daiichi Sankyo Inc

5
7.5 mg

(Continued)

Table 8.1 (Continued) Deterring class

Deterring technology

Mechanism

Manufacturing process

Main deterring agent(s)

FDA-approved product name

Dosage form

Dose strength

Manufacturer

Chemical barrier

BeadTek

Extraction resistant matrix

PEO

Zohydro ER (hydrocodone bitartrate)

Extended-release capsules

10
50 mg

Pernix

Depomed’s Acuform and Mallinckrodt’s technology Deactacore technology

Extraction resistant formulation

Multilayer coating of sugar spheres Wet granulation and blending

PEO

Extended-release tablets

7.5/325 mg

Mallinckrodt

Multilayer pellets comprising a sequestered antagonist

Multilayer coating of sugar spheres

Naltrexone HCl

Extended-release capsules

20/0.8
100 mg/ 4 mg

Pfizer

Drug-antagonist matrix




Naloxone HCl

Drug-antagonist matrix




Naloxone HCl

Extraction resistant and nasal irritating formulation

Direct compression

PEO, SLS

Xartemis XR (oxycodone HCl and acetaminophen) Embeda (morphine sulfate and naltrexone HCl) Troxyca ER (oxycodone HCl and naltrexone HCl) Suboxone (buprenorphine and naloxone) Targiniq ER (oxycodone HCl and naloxone HCl) Oxaydo (oxycodone HCl)

Agonist/ antagonist combination

Agonist/antagonist

Aversion

Acura’s Aversion

EC, Ethyl cellulose; HPMC, hydroxypropyl methyl cellulose; PEO, poly(ethylene oxide); SLS, sodium lauryl sulfate.

10 mg/ 1.2
80 mg/ 90.6 mg Sublingual film

2 mg/0.5
12 mg/ 3 mg

Indivior Inc

Extended-release tablets

10 mg/5
40 mg/ 20 mg

Purdue Pharma

Immediate-release Tablets

5
7.5 mg

Egalet

Polymers and hydrogels to deter drug abuse

191

Figure 8.5 Major polymer properties utilized in the manufacturing of the abuse-deterrent formulations. CMC, Carboxymethyl cellulose; HPMC, hydroxypropyl methyl cellulose; PEO, poly(ethylene oxide).

solubility, organic solubility, molecular weight, melting temperature (Tm), glass transition temperature (Tg), binding, adhesion, and swelling.

8.3

Polymer properties in the abuse-deterrent products

Shown in Table 8.1, most FDA-approved abuse-deterrent (AD) products utilize various properties of polymers to deter abuse by insufflation and injection. These products are manufactured via either cold (conventional tableting) or hot processing (heated compression, fluidized bed, injection molding, and hot melt extrusion) techniques. In most of these products, a powerful viscosifying agent is an integral part of the formulation, while others benefit from the low molecular weight polymers and waxes. Fig. 8.5 shows major polymer properties that have been utilized in the manufacturing of the ADF products with abuse-deterrent labeling and features. These include thermal, rheological, swelling, and mechanical properties.

8.3.1 Thermal Thermal properties of polymers are best characterized by their glass transition and melting temperatures. These two features have successfully been used in the preparation of crush resistant ADFs. Crystalline polymers with low Tm and amorphous polymers with low Tg have been reported as successful candidates for such application. Examples include high molecular weight poly(ethylene oxide) (PEO)and vinyl acetate vinylpyrrolidone polymer blends (Kollidon SR), respectively. Thermoplastic or thermosoftening polymers are low-to-high molecular weight polymers that undergo solid to rubber or liquid transition at above their glass transition or melting temperatures, respectively. Due to enhanced flow properties of the polymer at above its Tg or Tm, the final drug product containing such polymers possesses enhanced mechanical properties such as greater hardness, toughness, and less friability. A low Tg or low Tm enables the drug product containing such

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polymers to be processed at reasonably low temperatures that is desirable to avoid degradation and thermal oxidation of the drug as well as the excipients [11]. McGinity et al. disclosed pharmaceutical formulations containing a hot melt extrudable mixture of a drug and a high molecular weight PEO [12]. PEO is a synthetic thermoplastic water-soluble polymer available in a wide range of molecular weights (200,000
7,000,000 Da). PEOs with molecular weights below 25,000 Da are waxy solids or viscous liquid, and are commonly referred to as poly(ethylene glycols) [13,14]. PEOs at higher molecular weights are commonly used in the preparation of controlled-release drug products (e.g., osmotic tablets, gastroretentive dosage forms, and hydrophilic matrices). Although different grades of the PEO polymer surprisingly display melting points in a narrow range of 65 C
70 C, their flow properties are dramatically different. Muppalaneni et al. evaluated the thermal and flow properties of two different PEOs (100,000 and 7,000,000 Da) and found out adhesive property of the low molecular weight PEO is superior to that of high molecular weight counterpart. Greater flow and adhesion of the low molecular weight PEO polymer equipped the tablet with greater mechanical properties and hence crush resistance [15,16]. There are currently several patented ADF technologies such as INTAC, Guardian, and RESISTEC (Table 8.1) that utilize thermal processing of PEO to achieve crush resistance properties. PEO has been incorporated in many FDA-approved ADFs [e.g., Arymo ER, Hysingla ER, Nucynta ER, reformulated Oxycontin, reformulated Opana ER (now off the market), and Exalgo] (Table 8.1). These tablets are manufactured either by a thermal process (Nucynta ER [17], Opana ER [18], Arymo ER [19], Oxycontin and Hysingla [20]), or wet granulation (Exalgo [21]). Given its very hydrophilic nature and solubility in water, the crush resistance feature of the ADFs containing PEO may severely damage if the product is manipulated under wet conditions. One way to overcome this issue is to use low Tg waterinsoluble polymers in the formulation. Omidian et al. disclosed the use of Kollidon SR in the preparation of ADFs [16,22]. Kollidon SR is a polymer blend of vinyl acetate (PVA) and vinylpyrrolidone (PVP), which is primarily used in the preparation of pH-independent sustained-release tablets by direct compression or hot melt extrusion. The product contains 80% PVA, 19% PVP, and about 1% of sodium lauryl sulfate and silica [23]. Owing to low Tg of the polymer, the products containing Kollidon SR can be compressed into a tablet at a very low compression force. Muppalaneni et al. have shown that tablets composed of Kollidon SR, when processed at low and high temperatures, not only provide abuse-deterrent crush resistance property but also provide two different immediate and sustained drug release profiles. It was shown that the heat treatment of the tablets at above the glass transition temperature of the PVA component of the Kollidon SR could improve the tablet mechanical properties as evidenced by the larger proportion of coarse particles obtained upon crushing. Tablets containing a drug model, Kollidon SR, and some other excipients were thermally processed and evaluated for their crush resistance using different manual and powered crushing tools. For the nontreated tablet, 88% of the crushed particles were smaller than 250 μm, while about 90% of the crushed particles of the heat-treated tablet were larger than 850 μm making them unsuitable for abuse via a nasal route [15,16,24].

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8.3.2 Rheological Rheological properties of polymers are best characterized by their radius of gyration in different solvents; the longer the radius of gyration, the greater the solution viscosity. With a single pair of a polymer-solvent system at a given polymer molecular weight, the radius of gyration is primarily determined by the type of solvent, pH, temperature, and ionic strength. In other words, a given polymer at certain molecular weight can provide various radii of gyration depending on the service conditions. This situation is far more complicated in drug abuse as abusers freely use as many solvents as possible to maximize the amount of extractable drug. According to the FDA guideline, the most commonly used solvents include water, hydroalcoholic solutions (at different alcohol content), buffers, vinegar, and water containing baking soda [25]. Therefore the preferred viscosifying agent is the one that can provide the desirable viscosity in almost all solutions, and this is far from reality. Despite the challenge, ethylene oxide polymers have widely been used for this purpose [18,19,26]. Other polymers including xanthan gum [27], methyl cellulose, hydroxypropyl methyl cellulose [28], and carboxymethyl cellulose [29] have also been used in the preparation of ADFs. Supplied up to very high molecular weights (e.g., 7,000,000 Da), PEO is a nonionic polymer that has minimum (if any) interaction with various drugs and excipients [30]. The amount of PEO that can offer deterrence to the tablet can vary in the range 40
90 wt.% per tablet weight [18,20,26,31]. Given that the abusers normally use up to 10 mL of a solvent to extract the drug, the PEO concentration of 2
3 wt. % would be more than sufficient to build significant viscosity in aqueous media, preventing drug extraction and hence abuse by injection. PEO is widely used in the ADF products manufactured via cold processing. Products including Oxaydo immediate-release tablets (manufactured by simple blending and direct compression [32]), Xartemis extended-release tablets (manufactured by wet granulation [26]), and Zohydro extended-release capsules (manufactured by filling the capsules with indistinguishable beads of PEO and opioid drug) [33] are the examples of such preparation that are expected to resist drug extraction and deter abuse by injection. On the other hand, several other products containing PEO such as reformulated Opana ER, reformulated OxyContin, and Nucynta ER are manufactured via hot processing [18,20]. These products are expected to provide both extraction resistance and crush resistance (as explained in Section 8.3.1). Therefore compared with other conventional highly functional viscosifying agents such as hydroxypropyl methyl cellulose used in Vantrela ER extended-release tablets [28], the high molecular weight PEO polymers can potentially offer deterrence to abuse by insufflation and injection. Despite their benefits in both solid and solution states, the PEO polymers can lose their unique deterrence properties under variety of abuse conditions. The PEO undergoes thermal oxidation [34] at around 170 C and loses its large radius of gyration when solution is near boiling and is concentrated in salts [35] and alcohol as well as agitated at high shear rates [36]. Joshi et al. studied the PEO polymers under various extraction conditions by increasing the solution temperature or adding

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various percentages of alcohol to the solution. Their results showed that the deterrence properties of PEO can easily be defeated when the drug product containing PEO is manipulated using conventional materials and tools such as heating by a lighter or a candle or by adding alcohol to the solution [37,38]. Additionally, it was demonstrated that the rate at which the PEO increases the viscosity of the extraction medium is far slower than the rate at which drug is dissolved in the same medium. Therefore abusers would have enough time to extract the drug and abuse it by injection. Furthermore, the use of high amounts of PEO in oral tablets has been associated with choking and other dysphagia-like symptoms [39]. Other adverse effects such as thrombotic microangiopathy, visual disturbances, and kidney dysfunction have also been reported with products containing PEO that were abused via the IV route [40].

8.3.3 Swelling Polymers that can swell in aqueous media can alternatively be used to prevent drug extraction. Swellable polymers or hydrogels are best characterized by their crosslink densities; the higher the crosslink density, the less swellable the polymer will be in a specific solvent. Swellable polymers can range from lightly to very highly crosslinked structures, and both have found applications in the manufacturing of the ADFs. The lightly crosslinked swellable polymers are called superabsorbent, and very highly crosslinked swellable polymers are currently being used as superdisintegrant in the preparation of pharmaceutical dosage forms [41]. High-swelling superabsorbents are generally based on hydrophilic monomers including acrylic acid, acrylamide, and sulfopropyl acrylate [42]. Although low-swelling polymers can also be made using similar monomers, those commonly used in pharmaceutical preparations are made of vinylpyrrolidone (crospovidone; Kollidon CL), carboxymethyl cellulose (croscarmellose sodium; Ac-Di-Sol), and carboxymethyl starch (sodium starch glycolate, Explotab) [43]. As the term implies, high-swelling superabsorbent polymers can swell up to hundreds to thousand times of their own weight in solutions [44]. Given the fact that abusers use up to 10 mL of a solvent for drug extraction and injection, even minute amounts of a high-swelling polymer would be more than sufficient to absorb the solution and entrap the drug within its structure. High-swelling superabsorbents such as those based on anionic or cationic monomers are sensitive to solvent composition, salts, and alcohol [45
48], and their swelling and entrapment efficiency can largely be affected under harsh conditions. On the other hand, those based on nonionic monomers such as acrylamide have shown to be more resistant to solvent composition and display greater swelling capacity in hydroalcoholic solutions [22,49,50]. Mastropietro et al. [39] have shown that superabsorbents have little-tono impact on the tableting process, and the tableting process has little-to-no effect on the swelling properties of superabsorbents. They also studied the effect of superabsorbent on the drug release profile in a single layer and bilayer tablets, and found out that superabsorbents, even when used at high concentrations, did not affect the drug release profile [39,51].

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Figure 8.6 Abuse deterrence using low-swelling polymers.

As opposed to the tablets containing PEO that suffer from a slow hydration kinetics, those comprising superabsorbents can absorb the extraction medium in about few minutes and can effectively prevent drug extraction and hence abuse by injection. Although high-swelling polymers help to deter abuse by solvent absorption, the low-swelling polymers function during the later steps of the extraction process. Even with their low-swelling capacity, the size of the swollen polymer is still large enough to block the filter pores and clog the syringe needle that are further used in the process of abuse by injection (Fig. 8.6).

8.3.4 Mechanical As far as their mechanical properties are concerned, pharmaceutical polymers range from those providing high modulus to those possessing great toughness. Drug products containing high modulus polymers can resist high crushing forces; however, they fail under force via brittle fracture mechanism. On the other hand, drug products containing tough polymers resist crushing force and may break via a ductile fracture mechanism [52]. Both features have been used in the preparation of the ADFs. Mastropietro et al. have shown that heated tablets containing a drug model, superabsorbent polymers and Kollidon SR could resist high crushing forces and generate large granules after crushing [22,39]. Similarly, Muppalaneni et al. have shown that thermally processed tablets containing a drug model, crosslinked carboxymethyl cellulose, and low molecular weight PEO could resist high crushing force and generate very large pieces after crushing [15,16]. Alternatively, Omidian

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et al. disclosed inherently soft polymer matrices based on poly(vinyl alcohol) cryogels that can provide soft and rubbery drug formulations possessing very high crush resistance feature [16]. A cryogel is a product of a freeze-thaw process on certain polymer solutions. An aqueous solution of Poly(vinyl alcohol) (PVOH) can be transformed into a solid rubber-like material via a freezing-thawing treatment. Upon freezing a PVOH solution, hydroxyl groups (2OHs) of the adjacent polymer chains interact via intra- and intermolecular forces, creating an ordered water-insoluble structure. The process is favored when the PVOH material is highly deacetylated ( . 99% hydrolyzed), the aqueous PVOH solution is concentrated (up to 20 wt.%), and the PVOH molecular weight is in the range 50,000
130,000 Da. The PVOH in its solid form has been used in the pharmaceutical products including tablets, ophthalmic, implants, transdermal patches, and topical creams due to its unique film-forming, adhesion, and other desirable properties [53,54]. PVOH at certain molecular weights and degrees of hydrolysis has been used in ADF formulations. The polymer can be dissolved at very low concentrations (3
10 wt.%/vol.%) in water containing drug and other deterrent agents. The solution is then freeze-thawed under a programmed freezingthawing cycle until a very tough and soft rubbery material (cryogel) is obtained. A PVOH cryogel possesses desirable mechanical strength and viscoelastic properties that help its dosage form to resist crushing. Muppalaneni et al. have shown that a successful use of a PVOH cryogel in the preparation of ADFs requires a fine balance between elastic (hardness) and viscous (gumminess) properties of the PVOH matrix. Although processing factors, PVOH molecular weight, degree of hydrolysis, and freezing and thawing temperatures can significantly affect the viscoelastic properties of the PVOH cryogels, the most important factor is the PVOH concentration in the cryogel solution. At concentrations as high as 15
20 wt.%, an elastic cryogel is prepared that has successfully been used for biomedical applications where polymeric scaffolds with high mechanical properties are desirable. Omidian et al. have disclosed that PVOH cryogels prepared at concentrations as low as 3
5 wt.% display very desirable viscous properties that can be utilized in the preparation of the crush resistant ADFs. The cryogel tablets can build up adhesive (viscous or gumminess) properties when manipulated using crushing tools as shown in Fig. 8.7.

Figure 8.7 A cryogel tablet before crushing (left), after crushing (middle), and its adhesive property (right).

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The adhesive data measured by a texture analyzer showed that the change in cryogel adhesive properties is primarily a function of polymer concentration in solution. For instance, at 3 wt.% PVOH solution concentration, the cryogel adhesiveness was found around 0.16 mJ before and after crushing. However, as the PVOH concentration increased to 6 wt.%, the tablet adhesiveness increased from 0.03 to 0.19 mJ when crushed [16]. Nevertheless, despite their unique abuse-deterrent crush resistant properties, the PVOH cryogels are sensitive to temperature due to the fact that their gelation is only physical.

8.3.5 Binding Most drugs of abuse, in particular opioids, are supplied in the form of weak bases or salts of the weak bases, for instance, oxycodone and oxycodone hydrochloride. As such, polymers carrying carboxyl functional groups can potentially form electrostatic forces with positively charged opioids. Such interactions can interfere with the opioid extraction process in aqueous solvents. ADFs relying on the binding property of such polymers can be prepared by simple physical blending of the drug and the polymer. Tampering with such dosage forms triggers the binding in solution that reduces the amount of the free drug available for extraction. As an alternative to the drug polymer physical blend, a drug polymer complex can be formed first and then used in the formulation [15,55,56]. Examples of polymers that have successfully been used for this purpose include linear and crosslinked sodium carboxymethyl cellulose (croscarmellose sodium) and crosslinked carboxymethyl starch [22]. Although such polymers can provide a stable binding with the weak base in most aqueous solutions (Fig. 8.8), the drug can easily be released from the complex in gastric medium, providing a therapeutic level of the drug when used as intended. Compared to high molecular weight PEOs that offer deterrence to abuse by injection only by increasing the viscosity of the drug extraction solutions, functional viscosifying agents such as linear carboxymethyl cellulose [57] contribute to abuse deterrence via two mechanisms, enhancing solution viscosity and entrapping the drug through forming a complex in solution. Owing to their carboxyl functionality, the crosslinked polymers of carboxymethyl cellulose (XCMC) (such as Ac-Di-Sol) and carboxymethyl starch (XCMS) (such as Explotab) have also been extensively studied for their binding ability with the weak base drugs in various extraction solutions [16,22]. The XCMC and XCMS polymers are available as internally crosslinked polyacids that possess different functionality and crosslink densities. The amounts of drug that can be entrapped via complexation with these polymers is dependent on the available functional groups and the crosslink density of the polymer [55,58]. Since carboxyl groups are used up during the preparation of the crosslinked polymer, a polymer with more functional groups and a lower crosslink density can potentially entrap more drug via binding, offering a greater deterrence. Omidian et al. reported that the XCMC could bind reasonably well with tramadol (an opioid model drug) in aqueous solutions, and the drug binding was reversible at low pH (e.g., 0.1 N HCl). This feature allows a successful use of these polymers for oral delivery while providing a great deal of deterrence when the drug product is

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Figure 8.8 Deterring drug abuse via binding.

manipulated. Moreover, their kinetic data also showed that the binding of the polymer with the drug in solution is instant, a very desirable feature for deterring drug abuse by injection [39]. Owing to its powerful binding with weak bases, the XCMC has been included in the PVOH cryogel platform and the Kollidon SR formulations to enhance their extraction resistant capability. It has been shown that the PVOH cryogels prepared at as low as 3
5 wt.% PVOH concentration were able to accommodate more than 10 wt.% of XCMC, and offered very desirable extraction resistance properties [15]. Although highly crosslinked low-swelling polymers with anionic functionality can generally provide an effective deterrence to drug extraction and injection, their deterrence capability will be damaged in solutions of high ionic strength and alcohol content.

8.3.6 Film-forming The thermal, rheological, swelling, and mechanical properties of polymers have been directly utilized as physical and chemical barriers in the preparation of crush resistant and extraction resistant abuse-deterrent compositions. Owing to their desirable adhesion and film-forming properties, polymers have also been used in the preparation of agonist/antagonist ADFs. Embeda is an encapsulated multilayercoated pellets of morphine sulfate (an opioid agonist) and naltrexone (an opioid receptor antagonist). In this product, agonist and antagonist have been enclosed in

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Figure 8.9 Schematic structure of Embeda’s pellets incorporating several layers of functional coating using different grades of poly(meth)acrylates copolymers.

different layers separated by polymeric coatings. The main excipients in the coating layers are various poly(meth)acrylates. As shown in Fig. 8.9, the antagonist coated on the sugar spheres is covered with a sequestering copolymer of ethyl acrylate and methyl methacrylate with a low content of methacrylic acid ester having quaternary ammonium groups. This coating (seal coat) is insoluble in water and has low permeability, preventing the antagonist from being released when the drug is taken as intended. The seal coat is covered by the drug layer, which is in turn covered by another grade of poly(meth)acrylates. This grade of the polymer is water permeable and is dissolved at the duodenum pH, resulting in the therapeutic release of the drug. When the pellets crushed for abuse by injection, the sequestering effect of the seal coat is compromised. As such, the agonist bioavailability is reduced since both morphine and naltrexone are released from the dosage form [59].

8.4

Conclusion

Abuse-deterrent formulations (ADFs) have been developed to alleviate the epidemic nature of the prescription drug abuse. Although less than 20 ADF products are currently available in the US market, most of these products offer deterrence to abuse using a limited number of polymers. Thermal, rheological, swelling, mechanical, binding, and adhesion properties offered by these polymers have enabled the pharmaceutical manufacturers to develop therapeutic opioid medications that can resist crushing and drug extraction and therefore deter drug abuse by insufflation and injection. Numerous patents in the area of abuse and tamper-resistant formulations reflects the fact that polymers can still play a very major role in developing more effective and safer ADF products; however, the net effectiveness and safety of any of these polymers or technologies can only be proven following a thorough postmarketing studies.

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Glucose-sensitive materials for delivery of antidiabetic drugs

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Maria Saeed and Amr Elshaer Drug Discovery, Delivery and Patient Care (DDDPC), School of Life Sciences, Pharmacy and Chemistry, Kingston University, Kingston upon Thames, London, United Kingdom

9.1

Introduction

Diabetes mellitus is one of the oldest healthcare conditions to be reported in history. A description of the disease goes back to a 3500-year-old Egyptian manuscript. The papyrus Ebers is an ancient Egyptian book of medical herbal knowledge dating back to 1550 BC, found in 1862, which describes “polyuric syndrome” and its remedies. In modern times, this syndrome is known as diabetes mellitus [1]. People diagnosed with diabetes are increasing constantly and numbers will continue to grow for the foreseeable future. According to a global report on diabetes by the World Health Organization (WHO), the number of adults living with diabetes in 1980 rose from 108 million to 422 million in 2014. The report indicates a rise from 4.7% to 8.5% in adult population globally, over 18 years. In 2015, around 1.5 million deaths were reported due to diabetes. It is also evident that the incidence of diabetes is more prevalent in low- to middle-income countries, compared with high-income countries as shown in Fig. 9.1 [2]. Diabetes has become a major cause of blindness, kidney failure, myocardial infarction, stroke, and lower limb amputation. Nearly half of all deaths due to high blood glucose happen before the age of 70 years. The WHO considers diabetes to become the seventh leading cause of death by 2030. It is also reported that healthy diet, consistent physical activity, keeping a normal body weight, and avoiding tobacco use can prevent or delay type 2 diabetes’ onset. Diabetes can be managed and its consequences avoided or delayed with diet, physical activity, medicines and regular screening, and by treating complications [2].

9.1.1 Types and causes of diabetes Diabetes is a lifelong condition and can be divided into two main types: type 1 and type 2 diabetes. Type 1 or “immune-mediated diabetes” usually leads to absolute deficiency of insulin and is an autoimmune disorder, characterized by nonfunctional pancreatic β-cells. It was previously known as insulin-dependent diabetes or juvenile-onset diabetes and represents 5%
10% of diabetes. Glucose and hemoglobin A1C Engineering Drug Delivery Systems. DOI: https://doi.org/10.1016/B978-0-08-102548-2.00009-3 © 2020 Elsevier Ltd. All rights reserved.

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Figure 9.1 Prevalence of diabetes among WHO Regions in 1980 (A) and 2014 (B) [2].

(HbA1c) levels increase early before the clinical onset of the disease. Type 1 diabetes studies established that hyperglycemia is due to persistent presence of specific autoantibodies in such patients. However, autoimmune destruction of β-cells has many genetic susceptibilities, besides environmental factors can also be the cause for type 1, but family history plays a significant role. Type 2 diabetes is defined by insulin resistance caused by progressive loss of β-cells and comprises 90%
95% of diabetes. The cause of type 2 diabetes is poorly defined but it is commonly associated with age, obesity, and sedentary lifestyle. The other risk factors include hypertension, dyslipidemia, family history, and ethnicity such as Asians and Americans ethnic groups. Type 2 diabetes can also be associated with other diseases or infections or with use of some drugs (glucocorticoids and thiazide diuretics etc.) but usually remain asymptomatic during the early stages of the disease until blood glucose levels rise gradually and finally lead to micro and macro vascular complications. Type 1 and type 2 diabetes can occur in both children and adults and are not age restricted as described in the past [3].

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Gestational diabetes mellitus (GDM) is a third type of diabetes which occurs in females only during the second and third trimester of their pregnancies. One possible cause of GDM is claimed to be production of several hormones in the placenta leading to impairment of insulin action during pregnancy [4]. Women having GDM should be screened annually for diabetes and should adopt lifestyle interventions to prevent diabetes because of high risk of developing type 2 diabetes after delivery. Other types cover idiopathic, immunogenic, and monogenic diabetic syndromes which accounts for less than 5% of diabetes. It includes neonatal diabetes that occurs under 6 months of age, caused by a single gene alteration affecting insulin production and maturity-onset diabetes of young which occurs before 25 years of age, pancreatic disorders such as cystic fibrosis and pancreatitis, drug- or chemicalinduced diabetes, and diabetes after organ transplantation mainly caused by immunosuppressant therapy. All of these types lead to insulin insufficiency and greater mortality rates [5].

9.1.2 Diabetes management There are several antidiabetic medications present to normalize hyperglycemia. The treatment strategies are based upon lifestyle intervention alone and in combination with antihyperglycemic drugs as shown in Fig. 9.2. As the disease progresses and reaches chronic stages, the glycemic levels become difficult to control. However,

Figure 9.2 Flow chart diagram of diabetes and its management.

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Table 9.1 Types of diabetes and recommended products for its management [8]. Antidiabetic drugs

Generics

Brands

Diabetes

Insulin

Lispro Aspart Glargine Detemir Pramlintide

Humalog Novorapid Lantus Levemir Symlin

Type 1, type 2, CFRD, neonatal, MODY and GDM

Type 1 and type 2

Liraglutide

Victoza

Type 2

Metformin Glibenclamide Glipizide Pioglitazone Sitagliptin Acarbose

Glucient Gliken Minodiab Actos Januvia Glucobay

Type 2 and GDM Type 2, neonatal, and GDM Type 2 and MODY Type 2 Type 2 Type 2

Amylin analogues GLP-1 receptor agonists Oral hypoglycemics

GDM, Gestational diabetes mellitus; MODY, maturity-onset diabetes of young; CFRD, cystic fibrosis-related diabetes.

research is in progress to adopt novel treatments based on individual needs such as age, comorbidities, duration of disease, and life expectancy. According to WHO Diabetes Report, type 2 diabetes can be delayed or prevented in people with impaired glucose tolerance and obesity. A systematic literature review has also revealed that education combined with exercise and diet modifications might result in long-lasting effects in diabetes control [6,7]. In addition, pharmacological interventions have been widely used over the years (Table 9.1).

9.2

Need to redesign insulin delivery systems

It is obvious that exogenous insulin plays a major role in the management of diabetes and is usually self-administered. Type 1 diabetes is managed exclusively with insulin, while type 2 diabetes is initially managed by oral antidiabetic molecules combined with lifestyle modifications [9], yet usually progresses to insulin deficiency depending upon prognosis of disease and age, therefore, requires an insulin regimen as well. Moreover, patients with diabetes are bound to follow a strict insulin administration schedule which could not be controlled with oral hypoglycemics, in order to achieve the normal level of blood glucose [10]. Although insulin therapy is well tolerated by most patients, complications have been reported as a result of poor compliance and inadequate glycemic control [11]. These complications range from acute hypoglycemia [12], nonhealing wounds, cardiovascular mortality to coma, and death due to lingering variability in blood glucose levels [13,14].

Glucose-sensitive materials for delivery of antidiabetic drugs

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Apart from the importance of subcutaneous insulin injection, daily insulin injections may not be capable of normalizing blood glucose levels and delaying longterm complications of diabetes [15]. This is due to numerous reasons including improper dose calculation, pain and infection at the site of injection which often lead to noncompliance [16]. Nevertheless, insulin formulations have been improved significantly as compared to the past and are still under pharmacokinetic adjustments, Nonetheless, there is a need to redesign these formulations to achieve better glycemic control and to circumvent both acute and long-term complications associated with diabetes treatment [17,18]).

9.3

Glucose-sensitive materials

So far, the most cutting-edge approach toward closed loop therapy for diabetes is to develop smart material-based system or formulation that can sense glucose in the surrounding tissues. These formulations are designed with matrices of glucoseresponsive material in which the insulin is embedded. The matrix undergoes structural modifications upon changes in glucose concentration, followed by the insulin release. Glucose-sensing systems are generally composed of one of three different mechanisms: (1) enzyme-catalyzed pH changes, such as using glucose oxidase (GOx); (2) multivalent glucose-binding proteins, for instance, concanavalin A (Con A); or (3) molecular recognition by diol-binding chemical moieties such as phenylboronic acid (PBA) [13].

9.3.1 Glucose oxidase An approach to prepare a glucose-sensitive insulin delivery system is based on using the enzyme GOx which is widely used in glucose-sensing technologies. GOx can be embedded into pH-sensitive matrices such as hydrogels [19,20], liposomes [21], and microcapsules [22]. These types of glucose-sensitive systems can be prepared by using immobilized GOx in pH-sensitive polymers that can control insulin release. The GOx enzyme generally works by production of gluconic acid (Fig. 9.3) which lowers the pH and causes the swelling of the hydrogels. The latter helps in disassembling the liposomal components or microcapsules which leads to the release of the entrapped insulin [23]. Glucose sensitivity of such systems is oxygen dependant of their surroundings. As GOx is a flavoprotein and requires an oxidizing agent to function. Oxygen reoxidizes the flavin adenine dinucleotide from its reduced state gained after glucose consumption [24,25]. As the reaction proceeds, gluconic acid accumulates and subsequently decrease the pH which protonates the polyelectrolytes present in the drug delivery matrix. The gain and loss of protons by these polyelectrolytes will cause structural changes

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Engineering Drug Delivery Systems

OH

HOOC O

OH

OH

+ O2 HO

OH

HO OH

+

H2 O 2

HO

OH D-glucose

Glucose oxidase

Oxygen

HO D-gluconic acid

Hydrogen peroxide

Figure 9.3 Chemical reaction representing conversion of glucose into gluconic acid by glucose oxidase enzyme and in presence of oxygen.

Figure 9.4 Glucose-sensitive formulation based on glucose oxidase as glucose-sensing moiety, indicating physical changes in polymeric matrix of upon contact with glucose and insulin release.

such as swelling or shrinking to the drug delivery matrix which subsequently leads to insulin release (Fig. 9.4). For instance, at low pH, the glucose-sensitive systems based on polyanionic electrolytes will form hydrogen bonding within the copolymerized matrix and this hydrogen bonding will cause the matrix to collapse due to raised hydrophobicity in the polymer network. Eventually, this leads to squeezing of the matrix which in turn releases insulin [15]. While cationic polymers will swell up to release the insulin [15]. Many recent studies have utilized GOx to design glucose-sensitive drug delivery systems and devices (Table 9.2).

Table 9.2 Glucose oxidase-based drug delivery systems. Polymeric system used GOx-catalase (CAT)based peptidehydrogel

Summary G

G

G

G

GOx-modified erythrocyte as insulin carriers

G

G

G

Oleic acid-graftedaminated beta cyclodextrin (OAg-ACD) copolymer, coated with GOx and CAT

G

G

G

Mechanism of release

In vivo studies

In vitro insulin release

Reference

pH-sensitive insulin delivery Injectable Sustained drug release Biocompatible

Glucose
GOx interaction lowering pH, causing peptide disassembly and insulin release

Progressive insulin release up to 45% and a four-fold rise over 21 hours in hyperglycemic solution (400 mg/ dL)

[26]

Pulsatile insulin release Ultrafast glucose sensitivity Injectable

Hydrogen peroxide formation alongside gluconic acid ruptures erythrocytes to release insulin

25 μg/mL insulin release in 25 mmol/L glucose within 2 minutes

[27]

Intravenous insulin delivery Hydrophobic copolymer used to increase encapsulation capacity Strict insulin control

Drop in pH around outer coating causes system to swell and release of drug

Blood glucose level (BGL) dropped sharply after first hour in streptozotocin (STZ)-induced diabetic mice and remained low for 9 hours Controls hyperglycemia within 1 hour in STZ-induced diabetic rats and maintain normoglycemia for 9 days NA

Initial release in 30 minutes, reaching 78% in 4 hours when exposed to 400 mg/ dL glucose

[28]

(Continued)

Table 9.2 (Continued) Polymeric system used Chitosan-alginatecoated dextran nanoparticles crosslinked by GOx and CAT network containing insulin

Polyethylene glycol (PEG)
poly(serKetal and PEG polyserine nanovesicles containing insulin, GOx and CAT β-cells encapsulated microneedles patch with glucose signal ampifiers: containing GOx, α-amylase (AM) and glucoamylase (GA)

Summary G

G

G

G

G

G

G

G

G

G

G

G

Injectable and long acting formulation pH-sensitive Biodegradable Prepared by double emulsionbased solvent evaporation Biocompatible Long acting Noncytotoxic against HeLa cells (MTT assay) pH and glucose sensitive Minimum chance of infection Nonhypoglycemic in healthy mice Tight BG control for prolonged time System contain first synthetic amplifier for quick glucose detection

Mechanism of release

In vivo studies

In vitro insulin release

Reference

Polymer dissociation in hyperglycemic solution and structural deformation carried out by gluconic acid formation releasing insulin

BGL maintained normal up to 3 weeks in STZinduced diabetic mice

1400 μg/mL insulin released in presence of 400 mg/dL glucose over 8 hours and pH decreased from 7.4 to 4.2

[29]

Passive diffusion of glucose within bilayers of system lowering pH and hydrolyze nanovesicle assembly GOx, AM, and GA synergistic effects on gluconic acid formation, triggering β-cells to release more insulin from MN

Single (SC) dose kept BGL under 200 mg/ dL for up to 5 days in STZ-induced diabetic mice

Rapid insulin release (1100 μg/mL) in 400 mg/dL glucose solution within 4 hours

[30]

With one patch BGL dropped nearly 200 mg/dL for 2 hours in chemically induced diabetic mice

More than 15 μg/mL insulin release recorded in an hour when subjected to 400 mg/dL glucose concentration

[31]

Microneedle-array patch with nitroimadazole conjugated hylauronic acid containing insulin and GOx

G

G

G

G

Insulin- and GOxloaded microneedle patch with diblock copolymer of PEG and polyserine conjugates

G

G

G

Hylauronic acid microgels with ketal-modified dextran nanoparticles containing GOx

GOx, Glucose oxidase.

G

G

G

Painless, biocompatible and easy to administer Hypoxia-sensitive Better skin penetration properties Insulin burst release avoided Hypoxia and H2O2 dual sensitive system Painless microneedle-array patch H2O2 consumption controlled free radical-induced skin damage Injectable Hylauronic acid prevented insulin burst release and loss of GOx Nanoparticles crosslinked into microgels by emulsion method

High BGL stimulated gluconic acid production, creating hypoxia within system leading dissociation of vesicles and insulin release

BGL in mice reduced to about 200 mg/dL within 0.5 hours and takes up to 4 hours before gradual rise again

Approximately 250 μg/mL insulin released in 400 mg/ dL glucose levels in 4
6 hours and pH reduced progressively in 4 hours

[32]

Hypoxia induced by glucose
GOx reaction. And H2O2, production caused system dissociation and insulin release

Rapid decline in BGL in 1 hour. Normal glucose levels recorded up to 6 hours in STZinduced adult diabetic mice

90 μg/mL insulin released over 12 hours in typical hyperglycemic level (400 mg/dL)

[33]

Hyperglycemia leads to gluconic acid formation and nanoparticles dissolve to release insulin

A maximum of 1000 μg/mL insulin released over 6 hours in hyperglycemic conditions

Single SC injection in chemically induced diabetic mice maintained normoglycemia (200 mg/dL) for 1 week

[34]

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Advantages and limitations of glucose oxidase-based systems

GOx-based systems have excellent glucose-sensitivity and are commonly employed in blood glucose measuring devices [25]. Besides, GOx is highly selective to glucose and more stable than other enzymes. GOx can be used in erodible polymers, for instance, polyesters to accelerate matrix degradation at elevated glucose levels [13]. The GOx-based systems are pH and glucose dual responsive drug carriers which make them potential candidates for self-regulated insulin release systems [35]. One of the limitations of enzyme-based glucose-sensitive systems—involving GOx—is that their activity is lost quickly when subjected to increased temperature and high or low pH. pH values less than 2 or more than 8 and temperatures over 40 C may damage the system completely and irreversibly [25]. Similarly, GOx enzymes can be seriously affected by the use of surfactants such as sodium-dodecyl sulfate when pH becomes low and hexadecyl trimethyl ammonium bromide at elevated pH. Exposure to humidity can also affect the GOx-containing glucose sensors and the enzyme’s activity is lost under such conditions [25]. Another shortcoming is related to the presence of oxygen to convert glucose into gluconic acid. As the enzyme requires free oxygen in its surrounding and if oxygen concentrations are low, this will limit the enzyme efficiency [36]. Nonetheless, GOx-based systems are associated with enzyme denaturing by formation of covalent bonds with polymeric matrix which can significantly affect the efficiency of this system and reduce the shelf-life of such insulin delivery systems [37]. Moreover, there are limited clinical trial data available due to the immunogenicity and slow structural changes (swelling and disassociation of matrix) of GOxcontaining drug delivery systems [38].

9.4.1 Concanavalin A Lectins are multivalent proteins with binding abilities to carbohydrates, glycoproteins, and glycolipids. Lectins were discovered about 100 years ago and are found abundantly in nature. Interestingly, lectins are well known for their highly specific sugar-binding characteristics. Among these, Con A is one of the most studied lectins for its glucose sensitivity [15]. Con A is obtained from the plant “Canavalia ensiformis” or “Jack Bean” [39]. It is the first commercially available lectin used frequently for purification of various glucose-containing molecules and cellular structures [40]. Con A is composed of 237 amino acids and contains four binding sites that can bind to mannose and glucose. Each binding site has approximately 26,000 Da of molecular weight [41]. Con A has been reported in various studies as a glucosesensitive material and has been used in insulin delivery systems with promising results [35].

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213

Con A can form a complex with a chemically modified form of insulin known as glycosylated insulin [42]. The glycosylated insulin is formed by the addition of glucose to insulin which then attaches to the carbohydrate-specific binding subunits of Con A and can be incorporated into a hydrogel. As glycosylated insulin is specifically designed, it can be competitively released from Con A in presence of free glucose. Thus, complementary binding between free glucose and Con A is achieved on its binding subunits [15,43]. This phenomenon can be exploited to prepare a glucose-regulated insulin release system such as hydrogels and can be used for insulin delivery to treat hyperglycemia [44]. A drug delivery model based on Con A
glycosylated insulin complex can be further crosslinked or coated with a polymer to form a hydrogel. This type of system has the capability to accumulate insulin. The polymer acts as a membrane allowing the diffusion of insulin and glucose in and out of the membrane, respectively [15]. As the free glucose can enter into the system via a polymeric membrane, it competes for its binding sites present on Con A (Fig. 9.5). Thus, the binding preference of Con A for free glucose causes insulin detachment from Con A [45]. Therefore, Con A acts as glucose sensor and when the blood glucose level rises, insulin is released from the hydrogel system into the blood [45]. This type of insulin delivery model is composed of phase-reversible hydrogels which change from gel into solution phase when free glucose enters the system. This causes the hydrogel system to swell and increase in size which ultimately allows the discharge of insulin. The system shrinks back and starts to gel upon depletion of the glucose from the surrounding microenvironment [46]. Con A has been widely used in a number of self-regulated insulin drug delivery systems owing to its inbuilt glucose sensitivity and some of the latest studies are reported in Table 9.3.

Figure 9.5 Polymer-coating concanavalin A as glucose-sensitive material, free glucose enters the system and replaces insulin attached to concanavalin A assembly, triggering insulin release.

Table 9.3 Concanavalin A-based drug delivery systems. Polymeric system used Insulin-loaded Con Aglucosyloxyethyl methacrylate (GEMA) and N-(2(dimethylamino) ethyl)methacrylamide (DMAEMA) microhydrogels Insulin-loaded dextran-G and Con A microhydrogels

Summary of the study outcomes G

G

G

G

G

Insulin-loaded glucosyloxyethyl acrylated chitosan and Con A microgels Con A-amylopectin nanoparticles for insulin delivery

G

G

G

G

G

Mechanism of release

In vivo studies

In vitro studies

Reference

Glucose- and pHsensitive insulin delivery Apparent diffusion coefficient determined by Fick’s law

Free glucose-driven swelling of system to decrease crosslinking density and insulin release

NA

[47]

Injectable Reversible insulin release in response to variable glucose concentration Noncytotoxic (via MTT assay) Noncytotoxic Pulsatile insulin release influenced by varying glucose concentrations Spherical diameter range 5 100
300 nm EE 5 69.73% Insulin release at physiological pH (7.5)

Discosication of the Dex-G and Con A complex due to free glucose causing insulin release

NA

0.4 mg insulin released by stepwise increase in glucose concentration ranging from 0 to 18 mg/dL in 60 minutes 0.20 mg insulin release by stepwise glucose addition from 0 to 10 mg/dL

Swelling of microgel in presence of free glucose resulted in insulin release Glucose-induced disintegration of nanoparticles releasing insulin

NA

Around 37 μg insulin release in 4 mg/dL glucose

[49]

NA

Cumulative insulin release of 65% after 48 hours with initial burst release

[50]

[48]

Dextran
methacrylate and Con A methacrylamide crosslinked gels loading insulin

G

G

Con A-based hydrogel glucose sensor

G

G

G

Con A and insulin loaded network of poly(Nisopropylacrylamide) (poly(NIPAM)) nanogels.

Insulin containing Con Amodified PLGA nanoparticles.

G

G

G

G

G

G

Effect of varying acrylic substitution on dextran and Con A is investigated Insulin influx affected by degree of substitution on dextran and Con A Ultrafast glucose sensing due to ultrathin structure Glucose detection limit 5 0
70 mM Functional at physiological pH, temperature, and ionic strength Glucose dependant pulsatile release. Good stability. High insulin loading capacity.

Slow release, stable and good absorption profile In vitro, ex vivo and in vivo studies performed Higher insulinloading capacity.

Hyperglycemia caused dissembling of crosslinked matrix facilitating insulin flow

NA

In 1 hour, maximum insulin flow observed in 5.8% dex-methacrylate containing gel

[51]

Permeability is enhanced by glucose causing swelling of hydrogel and drug release

NA

NA

[52]

A reversible and rapid volume phase transition in response to glucose concentration releasing insulin. Biphasic release mechanism: initially diffusion through polymer matrix followed by matrix dissociation.

NA

Pulsatile release observed every 30 minutes in 15 mM glucose.

[53]

BG level dropped to 196 mg/dl after 3 hours of oral dose in Wistar rats and further reduced to 87.50 mg/dL after 24 hours. (ex vivo: 77% insulin absorption in 1 hour across intestinal mucosa.)

Glucose-dependant insulin release, from 9.3% in first 2 hours to 58.1% in 24 hours

[54]

216

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Engineering Drug Delivery Systems

Advantages and limitations of concanavalin A-based systems

Con A can be used as a nonenzymatic glucose biosensor with high specificity and excellent sensitivity [55]. The main advantage of incorporating Con A in selfregulated insulin delivery systems is its strong glucose sensitivity among other glucose-sensitive materials including PBA and GOx-based insulin carriers [56]. Con A is available commercially and is purified from plant sources. There are several different methods to extract Con A from jack beans [44]. Con A is a natural biodegradable compound thus avoids the unwanted side effects such as blood clearance. Besides, Con A is chemically stable at body temperature (37 C) [57]. A major drawback of Con A-based hydrogel system is the progressive leakage of Con A through polymer membranes leading to irreversible swelling and significant loss of therapeutic activity [15]. Moreover, this might be associated with unwanted immunogenic reactions. Another issue of this model is chemical modification of insulin is mandatory to enable binding with Con A. Therefore, every single insulin molecule should be glycosylated and this is a tedious and complex process. In addition, glycosylated insulin is regarded as a new chemical entity which requires approvals of regulatory bodies and may affect the therapeutic efficacy of insulin [15]. The binding constant between Con A and glycosylated insulin must be higher than glucose itself or there is a risk of developing hypoglycemia. Con A-based hydrogels showed considerable leakage of Con A, due to viscosity changes triggered by free glucose [15] which results in dose fluctuation and might be associated with episodes of hyperglycemia and hypoglycemia. Owing to limited binding sites and large size of Con A, it is not well suited as a drug carrier for long-term use as compared with other new glucose-sensitive materials that are actively being researched [15]. Finally, there are limited clinical studies to show effects of Con A-based preparations on glucose levels in animals.

9.5.1 Phenylboronic acid PBA and its derivatives have shown some potential applications for self-modulated insulin delivery. PBA and other compounds containing boronic acid have been widely used in preparation of glucose-responsive systems such as microgels, micelles, and nanoparticles. These compounds can form reversible cyclic boronic esters with polyols such as cis-diol compounds like glucose when dissolved in water [58,59]. PBA and its derivatives are reported as a Lewis acid having pKa value within 8.2
8.86 [60]. PBA in aqueous solution exists in two forms, that is, a neutral or uncharged trigonal-planar and a negatively charged tetrahedral boronate form in equilibrium as shown in Fig. 9.6 [61]. The uncharged form of PBA moiety is hydrophobic while the negatively charged tetrahedral form is hydrophilic in nature. Both charged and uncharged forms can form a reversible complex with cis-diol

OH HO HO

B

OH

HO

OH B OH

HO

OH R Hydrophobic form of PBA

OH OH

HO

O

O

OH R Hydrophillic form of PBA

CH2OH

O B

O OH

R Glucose attach to PBA

Figure 9.6 Equilibrium between charged and uncharged forms of phenylboronic acid and glucose attachment to negatively charged phenylboronic acid.

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compounds such as glucose. However, the complex formed between neutral form of PBA and 1
2 cis diols of glucose is unstable and highly prone to hydrolysis. Thus, the negatively charged PBA moiety can reversibly combine with glucose and give rise to a hydrophilic product which shifts the equilibrium in direction of forming more negatively charged form and decreasing the uncharged hydrophobic form [59]. It has been also reported that by increasing the chemically stable negatively charged form of PBA, the degree of swelling of crosslinked nanogels or micelles also increases due to repulsion between similar charges. As a result of swelling of the matrix, any entrapped drug molecules, is released out of the system (Fig. 9.7). This mechanism has been employed in construction of smart glucose-sensitive carriers for insulin delivery systems [61]. The PBA moieties can also be crosslinked with polymers such as polyvinyl acid (PVA) and exist in the form of gels. The glucose competes for the hydroxyl ends of PVA and PBA is attached with these glucose molecules at the other end. Thus, increase in substitution of PVA by free glucose causes polymeric matrix to swell and insulin is released from matrix (Fig. 9.7). While in the absence of free glucose, it may reform the gel and decrease the release of insulin [15].

Figure 9.7 Nanostructures containing phenylboronic acid grafted onto polymer chain and entrapped insulin. Entering of free glucose into nanostructure causes the attachment of glucose to phenylboronic acid and replacement of insulin, causing the nanostructure to swell. Nanostructure resume back into original size upon removal of glucose.

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PBA is one of the most commonly researched glucose responsive materials used in designing smart insulin delivery systems due to its unique characteristics. Several studies have been performed using PBA and are summarized in Table 9.4.

9.6

Advantages and limitations of phenylboronic acid

PBA and its derivatives are synthetic compounds, therefore the risk of developing immunogenicity and biotoxicity is very low as compared with other glucosesensitive protein-based compounds [71,72]. Besides, functionalized PBA possess better stability and long-term storability in comparison to GOx and Con A [35]. PBA and its derivates have been extensively studied and used in creating several types of drug vehicles including nanogels, vesicles, micelles, and nanoparticles with improved ionization at physiological pH and better structural integrity through crosslinking techniques [35]. In contrast, PBA systems are associated with many limitations. PBA derivatives have high pKa which limits their ionization at physiological pH resulting in poor water solubility and lower glucose-binding affinity. However, introducing electronwithdrawing groups at metaposition can decrease apparent pKa of PBA [59,73]. Also there is a lack of strong binding between insulin and PBA-based systems, affecting encapsulation efficiency but the issue can be resolved by using polymeric nanocapsules and self-assembly of glycosylated insulin and PBA [74]. Another vital issue of PBA containing systems is the self-regulated quantitative release of insulin. Until now, only a few models were successful and there is a lot of room for research into this specific area [75]. Poly(acrylic)-based PBA polymeric systems being synthetic are nondegradable and therefore their blood clearance and renal excretion is a challenge [63,64]. Moreover, another problem is preparing PBA-based oral nanocarriers that can be effectively absorbed in intestine. However, some chitosan-based oral systems have been reported [59].

Table 9.4 Phenylboronic acid-based drug delivery systems. PBA derivative used

Summary of the study outcomes

PBA-modified poly(D, L-lactideco-glycolide) (PLGA) microparticles and dopaminemodified hyaluronic acid hydrogel with insulin

G G

G

Insulin-loaded micells with monomethoxy poly(ethylene glycol)-b-poly(L-glutamic acid-co-N-3-Lglutamylamidophenylboronic acid) (mPEG-b-P(GA-coGPBA)) copolymers Poly 3-acrylamidophenylboronic acid-b-6-O
vinylazeloyl-dgalactose nanoparticles enclosing insulin

G G

G G G

G

Poly(ethylene glycol)-blockpoly [(2-phenylboronic esters-1, 3-dioxane-5-ethyl) methylacrylate] nanoparticles encapsulating insulin

G G

G G

Mechanism of release

In vivo studies

In vitro studies

Reference

Implantable insulin Biodegradable and highly glucose sensitive at physiological pH (7.5) Simple formulation based on mixing and self-assembling Biocompatible Noncytotoxic as confirmed by MTT and hemolysis assay

Glucose breaks DOP
PBA linkage to remove PBA from hydrogel, releasing insulin

BGL lowered to normal range and maintained in STZ-induced diabetic mice for 2 weeks

55%
77% insulin release in 500 to 1000 mg/dL glucose solution

[62]

Glucose diffusion into micelles caused formation of hydrophilic complexes and swelling of core to release insulin

NA

Rapid release (32.8%) within first 3 hours followed by 80% cumulative release in 15 hours

[63,64]

EE 5 63.8% Biocompatible Slight in vitro toxicity but no in vivo toxicity recorded 50
250 nm spherical EE 5 37% Functional at pH 7.5 and 37 C Low toxicity Size 5 50 nm in aqueous solution

Hydrophilic
hydrophobic interaction between copolymer and insulin is interrupted upon glucose detection, hence release insulin

Lowering of BG levels to normal range for 96 hours after oral dose in diabetic mice

Cumulative insulin release from NPs recorded as 80% in 3 mg/mL glucose

[65]

Dissociation of the selfassembly releasing insulin completely

NA

Total drug release in 8 mg/mL glucose solution for about 60 hours without any burst release

[66]

Insulin-loaded poly(3acrylamidophenylboronic acid-ran-N-maleated glucosamine) nanoparticles

G G

G

G

L-valine

and PBA-modified chitosan nanocarriers for insulin

G G

G

G

Poly N-vinylcaprolactam-coacrylamidophenylboronic acid nanoparticles containing insulin

G G

G G

Phenformin-loaded silica-PBAderived low molecular weight gel nanoparticles

G G G G

EE 5 80.15% Muco-adhesive and cytocompatible (via MTT assay) Nasal drug delivery 120
180 nm spherical EE 5 67% Low cytotoxicity against HT-29 cells. Stable in pepsin and pancreatic secretions. Glucose and pH dependent oral insulin delivery EE 5 65% No animal toxicity and cytotoxicity observed Spherical in shape Particle size is increased by high temperature EE 5 91.4% Biocompatible Noncytotoxic Glucose and pH responsive system for potential drug delivery

Biphasic release pattern: initial burst release followed by slow release in presence of hyperglycemic levels

Plasma glucose levels dropped to minimum in 9.5 hours in diabetic rats

41% insulin release reported after 72 hours in 5 mL PBS

[67]

Nanocarriers’ swell to release insulin in high pH and glucose concentration

BGL reduced in STZ-induced diabetic rats to 60% of the initial value for up to 8.4 hours

92% insulin released in 20 mM glucose solution in 24 hours

[68]

Nanoparticles swell up at increased temperature and glucose levels to release insulin

BGL reduced to 6 mmol in diabetic mice for at least 48 hours, before rising again

After rapid release in 1 hour, more than 85% insulin released in 30 hours in 3 mg/ mL glucose

[69]

Apart from glucose triggered drug release, hydrogen bonds are cleaved at high pH due to PBA ionization, leading to gel-sol transition and drug release

NA

Showing glucosedependant drug release, maximum of 71.4% phenformin release is reported in 30 mg/dL glucose solution

[70]

DOP, Dopamine; EE, Entrapemenrt Efficiency; PBA, Phenylboronic acid; BGL, blood glucose level; STZ, streptozotocin; NA, not applicable.

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Conclusion

In this chapter, three glucose-sensitive materials GOx, Con A, and PBA have been described to design glucose-responsive systems for insulin delivery. These glucosesensitive materials involve native mechanisms such as glucose binding using natural proteins and enzymes and synthetic molecular recognition. The systems formulated using these glucose-sensitive materials have been employed in the formation of hydrogels, nanoparticles, and micelles that facilitate rapid glucose diffusion as reported in several studies described above. These glucose-sensitive materials usually undergo physical and chemical variations, for example, the ability of hydrogels to swell, shrink, and degrade in response to glucose. Each approach has its own limitations, ranging from immunogenicity risks of the protein components such as in case of GOx or Con A to reduced glucose selectivity and restricted role in physiologic conditions for PBA-based systems. The development of “smart drug delivery” approaches for diabetes treatment seems hopeful, provided that there is a need to adjust the dosage depending upon actual disease status of the patient. The absolute target is to create a formulation mimicking the human pancreas, which can detect blood glucose variations and respond accordingly by releasing as much insulin as needed [76]. However, as the natural glycemic control mechanism in the body responses effectively to both high or low blood glucose concentrations, designing a similar but a synthetic system based upon glucose sensing is a real challenge. The stimulus for this specific therapy, glucose, is an abundant, small molecule present in both healthy and diseased states, although in different concentrations. Similarly, the insulin also has a limited stability and, if administered improperly, it can be lethal [77]. An ideal insulin delivery system would be able to rapidly react to the raised levels of glucose in blood and should promptly stopover to prevent overdosing of insulin. It should also avoid burst release or dose dumping of the drug which may potentially lead to unwanted side effects and subsequent hypoglycemia soon after administration. As diabetes is a lifelong condition and its treatment is continued over the life of the patients, the drug delivery system based on glucose-sensing elements must be able to administer in large population and in those with comorbidities as well. The system should obey predictable and practical dosing schedules, possess a suitable degradation pattern and clearance kinetics. Similarly, it should be noninvasive and noninflammatory, nonimmunogenic, and should maintain its activity over time. These are among the many challenges which are required to be addressed in the development of a fully functional glucose-responsive insulin delivery system in order to achieve a better control over diabetes.

References [1] A.M. Ahmad, History of diabetes mellitus, Saudi. Med. J. 23 (2002) 373
378. [2] WHO, Gloabl Report on Diabetes, WHO Press, France, 2016.

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3121.

Index

Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively. A Abuse-deterrent formulations (ADFs), 187
191 FDA-approved, 189t products, polymer properties, 191
199 binding, 197
198 film-forming, 198
199 mechanical, 195
197 rheological, 193
194 swelling, 194
195 thermal, 191
192 Acrylamide, 194 Acrylic acid, 194 ACT. See Adoptive cell transfer (ACT) Active and passive drug delivery, 166
168 Active cancer targeting, 37 Active Pharmaceutical Ingredients (APIs), 118, 147 Active targeting, 2
3, 4t Acute hypoglycemia, 206 Acute toxicity, 58 Additive fabrication, 147
148, 154
155 ADFs. See Abuse-deterrent formulations (ADFs) Adipoyl chloride, 28 Administration route, 18
20, 163
164 implants, 20 nasal route, 19 parenteral route, 18
19 peroral route, 18 pulmonary route, 19 transdermal route, 19
20 Adoptive cell transfer (ACT), 177 delivery of cells for, 178 Alginate hydrogels, 150 Alginic acid, 23 Alkali acid heat surface treatment, 28
29 AmBisome, 89

Amphiphilic copolymer polystyrene-altmaleic anhydride, 8 Amphiphilic dendrimers, 6 Amphotec/Abelcet/Ambisome, 6t Amphotericin B, 6t Amylin analogues, 206t Antibody-responsive hydrogels, 35 Anticancer drugs, 10
11, 36 Antidiabetic drugs delivery, 207
211 concanavalin A-based systems, 216
219 diabetes flow chart diagram, 205f management, 205
206 types and causes, 203
205, 206t glucose oxidase, 207
211 glucose oxidase-based systems, 212
215 insulin delivery systems, 206
207 phenylboronic acid, 219
221 Antidiabetic medications, 205
206 Antigen and antibody-responsive polymers, 35 Anti-HER2 monoclonal antibodies, 4t Antipsychotic medication, 18
19 Aptamers, 3 Arestin (minocycline hydrochloride for periodontitis), 126 Aromatic-based TPUs, 113 Asialoglycoprotein receptors, 4t Atridox (doxycycline hyclate for periodontal disease), 126 Autoimmune destruction, 203
204 B Betamethasone, 45 BIND-014, 96
97 Bioactive glasses, 128 Bioavailability, 73 Bioceramics, 127

230

Bioceramics-functionalized groups, 128 Biocompatibility, 7, 18
19, 75, 164
165 Biodegradability, 7, 18
19, 164
165 Biodegradable drug-loaded polymer implants, 124 Biodegradable drug-loaded wafer, 125
126 Biodegradable implants, 122
123 Biodegradable ISFIs, 127 Biodegradable materials, 53 Biodegradable metal alloys, 129 Biodegradable polymers, 123
126 natural polymers, 22
23 synthetic polymers, 23
24 Biopharmaceutical profiling, 60 Bioplotting, 51 Biopolymeric implantable DDSs, 165 Biopolymers, 164
165 Bioprinting technology, 49 Bioresorbable ceramics, 127
129 Bio-responsive polymers, 33
34 antigen and antibody-responsive polymers, 35 enzyme-responsive polymers, 34 glucose-responsive polymers, 34 Biotherapeutic agents, 5 Biotin, 4t Biotin-based NPs, 4t Biotin receptors, 4t Bittering agents, 188 Blood 2 brain barrier (BBB), 35 Bluestar Silicones, 114
115 Brain targeting, 35 Breast cancer tumor xenograft, 172
173 C CAD software, 147
148 Calcium phosphates, 128 Cancer immunotherapy, 177 therapy, 92, 96 tissues targeting, 36
37 Cancerous tissue, 2
3 Cardiovascular mortality, 206 Casting, 165 Celanese, 111
113 Cells for adoptive cell transfer, 178 Cellular targeting, 7
8 Characterization, DDS, 51
53 chemical properties, 53

Index

morphology, 51
52 physical properties, 52
53 Chiral dendrimers, 6 Chitosan, 23, 26
27 Chitosan microsphere, 10 Chronic inflammatory bowel disease, 35
36 Chronic toxicity tests, 58 Cisplatin, 10
11, 119 Coflowing rate, 10 Colloidal nanosystem, 2 Colonic systemic absorption, 35
36 Colon targeting, 35
36 Compudose, 115 Computer-aided design (CAD) software, 48
49 Computer-assisted printing, 49 Con A-based hydrogels, 216 Concanavalin A, 212
215 advantages and limitations, 216
219 composition, 212 phenylboronic acid (PBA), 216
219, 217f, 218f polymer-coating, 213f Contraceptive TPU-based vaginal rings, 119
120 Controlled DDSs, 19 Convection, 2 Conventional delivery methods, 47
48 Cosmetic emulsions, 45 CriPec (Cristal Therapeutics), 96
97 Cross-flow membrane emulsification, 10 Crosslinked polymers of carboxymethyl cellulose (XCMC), 197
198 Crosslinked polymers of carboxymethyl starch (XCMS), 197
198 Crushing, 185
187 Cryogel, 196
197 vaccine, 177
178 Curcumin-loaded polymeric micelles, 96 Cytarabine, 6t D Dapivirine, 117
118 Daunorubicin, 6t DaunoXome, 6t DDS. See Drug delivery system (DDS) Degradation/erosion process, 29

Index

Degradation mechanisms, 24
25 enzymatic degradation, 24
25 hydrolytic degradation, 24 Dendrimers, 6, 82
84 DepoCyt, 6t DepoDur, 6t Dexamethasone-loaded liposomes, 90
91 Dextran, 23 Diabetes flow chart diagram, 205f management, 205
206 types, 206t types and causes, 203
205 Diffusion-controlled drug, 29 Diffusion matrix/reservoir, 9 Difunctional poly(ethylene glycol) (DFPEG), 170
171 Diol-binding chemical moieties, 207 Direct dissolution, 87 Direct write, fabrication approach, 148
151 Dissolution matrix/reservoir, 9 Disulfiram, 119 Docetaxel (DTX), 10
11, 172 Double-emulsification, 82 Dow Corning, 114
117 Doxil/Lipodox/Myocet, 6t DOX-loaded gel (POL), 169
170 DOX-loaded hydrogel microparticles, 169 Doxorubicin, 6t, 97 Doxorubicin hydrochloride, 10
11 3D printing, 147
148, 150, 159
160, 165, 172 Drop-on-drop inkjet printing, 157
159 Drop-on-solid inkjet printing, 155
157 Drug bioavailability, 5
6 Drug carriers lipid-based nanocarriers, biomedical applications, 89
92 lipids based, 75
81 lipid nanocapsules (LNCs), 80
81 liposomes, 75
78 solid lipid nanoparticles (SLNs), 78
80 micro-/nanoparticles, schematic representation, 72f micro versus nanoparticles, 73
75 polymer-based nanocarriers, biomedical applications, 92
98 polymers and lipids as drug delivery systems, 72f

231

polymers based dendrimers, 82
84 nanogels, 84
85 nanospheres and nanocapsules, 81
82 polymeric micelles and polymersomes, 86
89 stimuli-responsive nanocarriers, 99 types, 75
89 Drug delivery system (DDS), 17. See also Implantable drug delivery systems correlation levels, in vivo and in vitro studies, 60
61 drug carriers preparation, 43
46 emulsion, 44
46 pharmaceutical foams, 44 pharmaceutical suspensions, 43
44 microfluidic technologies, 47
48 pharmaceuticals, microfabrication and molding, 46
47 injection molding technology, application, 47 manufacturing parameters, 47 materials used for injection molding, 47 physiochemical characterization of, 51
53 chemical properties, 53 morphology, 51
52 physical properties, 52
53 solid freeform fabrication (SFF), 48
51 in vitro, physiological environment for, 53
55 in vitro toxicity studies, 55
58 in vivo performance evaluations, 58
60 in vivo toxicity studies, 58 Drug dissolution, 54 Drug-loaded bioceramics, 128
129 Drug-loaded bioresorbable ceramics, 129 Drug-loaded hydrogel, 169
170 Drug-loaded ISFIs, 126 Drug-loaded polymer implants, 125
126 Drug-loading capability, 128 Drug 2 polymer affinity, 52 Drug release, defined, 54 Drug-screening, 150 Drug targeting brain targeting, 35 cancer tissues targeting, 36
37 colon targeting, 35
36 Drug transport via passive diffusion, 2

232

Dry foams, 44 DSM Biomedical, 113
114 Dual responsive nanogel, 96 DuPont, 111
113 Dyslipidemia, 204 Dysphagia-like symptoms, 193
194 E EGF, 4t EGFR, 4t Electric field-responsive polymers, 33
34 Electrospinning, 31, 165 Electrospun fibers, 31 Eligard (leuprolide acetate for prostate cancer), 126 Elvax, 20 Embeda, 198
199, 199f Emulsification, 30 Emulsification-evaporation, 82 Emulsion, 44
46 Emulsion-coacervation, 82 Emulsion-diffusion, 82 Encapsulation efficiency, 31, 46 Endocytosis, 3 Energy-releasing process, 53 Enhanced permeability and retention (EPR) effect, 2, 9 Enzymatic degradation, 24
25 Enzyme-catalyzed pH changes, 207 Enzyme-responsive polymers, 34 Estrogen-releasing implants, 115 Ethanol, 4
5 Ethinyl estradiol, 157 Ethylene monomers, 111
113 Ethylene vinyl acetate (EVA), 20, 153
154 monomers, 20 Eudragit RL, 156 Explotab, 197
198 Extravasation, 2 Extrusion, 148 F FDA-approved liposome-based pharmaceutical products, 6t FDM. See Fused deposition modeling (FDM) Femring, 116
117 Field-responsive polymers, 33
34 bio-responsive polymers, 33
34

Index

electric field-responsive polymers, 33
34 light-responsive polymers, 33 Flavin adenine dinucleotide, 207 Flexibility, design and fabrication, 164
165 Floating systems, 9 Flupentixol, 18
19 FMS-like tyrosine kinase-3 (FLT3), 96 Folate-based NPs, 4t Folate receptors, 4t Folic acid, 4t Folic acid 2,4-diaminobutyric acid (FADABA-SMA), 8, 8f synthesis and clinical applications, 8
9 Formulation design in drug delivery, 17
18 administration route, 18
20 implants, 20 nasal route, 19 parenteral route, 18
19 peroral route, 18 pulmonary route, 19 transdermal route, 19
20 biodegradable materials, 22
25 material shape, 31 material size, 30
31 material surface properties, 25
29 adjustment, 27
29 mechanical effects, 27 physiological factors, 25
26 surface and interfacial energies, 26 surface charge, 26
27 surface morphology, 25 nonbiodegradable materials, 20
22 smart materials, 31
35 field-responsive polymers, 33
34 pH-responsive polymers, 33 temperature-responsive polymers, 32
33 structural and bulk properties, 29 target cites, 35
37 Fourier-transform infrared spectroscopy (FTIR), 53 Freeze-drying, 82, 87 FTIR. See Fourier-transform infrared spectroscopy (FTIR) Fused deposition modeling (FDM), 50, 147, 151
154 printer, 172 Fused-filament fabrication 3D printing, 115
116

Index

G Galactose, 4t Galen Pharmaceuticals, 116
117 GDM. See Gestational diabetes mellitus (GDM) Gefitinib, 167
168 Gelatin-based materials, 49
50 Gelatin dip-molded capsules, 47 Gelatine polymeric microspheres, 18 Gelatin microparticles, 150f Gene delivery and transfection, 6 “Generation,” dendrimers, 82
83 Gestational diabetes mellitus (GDM), 205 Gliadel wafers, 164
165 GLP-1 receptor agonists, 206t Glucose oxidase, 207
211 glucose-sensitive formulation based on, 208f Glucose oxidase-based drug delivery systems, 209t advantages and limitations, 212
215 concanavalin A, 212
215 Glucose-responsive polymers, 34 Glucose-sensitive formulation, 208f Glucose-sensitive materials, 207
211 concanavalin A-based systems, 216
219 diabetes flow chart diagram, 205f management, 205
206 types and causes, 203
205, 206t glucose oxidase, 207
211 glucose oxidase-based systems, 212
215 insulin delivery systems, 206
207 phenylboronic acid, 219
221 Glucose sensitivity, 207 Glyceryl monostearate, 91
92 Glycoprotein CD44 receptor, 4t Grade 1 vaginal and cervical ulceration, 120 Grinding, 185
187 H HER2 mABs, 4t Hexamethylene diisocyanate (HMDI), 28 High-swelling superabsorbents, 194 Hormone-loaded coextruded core, 120 Hormone replacement therapy (HRT), 116
117 Hot-homogenization technique, 79

233

Hot melt extrusion (HME), 115
116, 118
119, 152 H2009.1 peptide, 4t Hyaluronic acid, 4t Hydrocortisone, 45 Hydrogel drug delivery systems, 168
171 Hydrogel Methocel E5, 149 Hydrogels, 148, 164
165, 169
170 Hydrolytic degradation, 24 Hydrophilic bioactives, 46 Hydrophilic/lipophilic balance (HLB) system, 44
45 Hydrophilic shell, 5
6 2-hydroxyethyl methacrylate (HEMA), 21
22 Hypertension, 204 Hypodermic injections, 47
48 Hypoglycemia, 206 I IL-13 peptide, 4t IL-13R α2 receptor, 4t Immune-mediated diabetes, 203
204 Immune-protection against cancer, 177
178 Immunogenicity, 3 Immunoliposomes, 5, 9 Immunomodulatory factors, 177
178 Immunotherapy modalities, 177 Implanon, 115
116 Implantable drug delivery systems, 164
165 biodegradable implants, 122
123 biodegradable metal alloys, 129 biodegradable polymers, 123
126 bioresorbable ceramics, 127
129 injectable in situ forming implants, 126
127 nondegradable ocular implants, 121
122 nondegradable polymers, 111
115, 112t nondegradable subdermal implants, 115
116 nondegradable vaginal rings, 116
121 Implants, 20 degradation rate, 122 matrices, 47 INFUSE, bone graft, 168 Injectable in situ forming implants, 126
127, 127t Injection molding (IM) technology, 46
47 Inkjet printing, 154
159

234

Inkjet printing (Continued) drop-on-drop inkjet printing, 157
159 drop-on-solid inkjet printing, 155
157 In situ forming implants (ISFIs), 122
123, 126
127 Insoluble divalent metal salts, 24 Insulin, 206t delivery systems, 206
207 -dependent diabetes, 203
204 therapy, 206 Integrin αvβ6, 4t Intelligent drug delivery systems active and passive drug delivery, 166
168 hydrogel drug delivery systems, 168
171 mechanisms, 166f microdevices delivery systems, 173
175 schematic, 164f therapeutic methods, 177
178 cancer immunotherapy, 177 cells for adoptive cell transfer, 178 immunomodulatory factors, 177
178 thermoplastic drug delivery systems, 171
173 transdermal patches delivery systems, 175
177 Inter-vertebral disk (IVD) degeneration model, 167
168 Intrauterine devices (IUDs), 21 In vitro dissolution testing, 120
121 In vitro 2 in vivo correlation (IVIVC), 60
61 Ionic bonding, 28 Ionic or nonionic surfactant, 44
45 J Jacalin, 4t Jet breaking, 30 Juvenile-onset diabetes, 203
204 K Kirsten rat sarcoma viral oncogene homolog (KRAS) oncogene, 171
172 L Lab-on-a-chip (LOC), 47
48 Lactate-based hydrophobic block, 96
97 Layer-by-layer manufacturing method, 48
49

Index

LCST. See Lower critical solution temperature (LCST) Lectin-based NPs, 4t Lectins, 212 Levetiracetam, 156
157 Levonorgestrel, 21, 115, 118 Ligand 2 receptor attachment, 2
3 Light-responsive polymers, 33 Linker-free covalent immobilization, 28 Lipid-based nanocarriers, biomedical applications, 89
92 Lipid-based nanoparticles, 19 Lipid film hydration method, 77
78 Lipids based drug carriers, 75
81, 80f lipid nanocapsules (LNCs), 80
81 liposomes, 75
78 solid lipid nanoparticles (SLNs), 78
80 Lipoid S75, 91
92 Lipophilic and hydrophilic surfactants, 45
46 Liposomal vesicular shape, 4
5 Liposome-based pharmaceutical products, 5 Liposome-manufacturing techniques, 5 Liposomes, 4
5, 5f, 75
78 classification, 77 conventional preparation techniques, 77
78 drug delivery systems, 76f fabrication, 10 Liquid foams, 44 LODER (local drug eluter), 164
165 Lower critical solution temperature (LCST), 32 Low-swelling polymers, 194 Lung cancer-specific CSNIDARAC (CC9), 97, 98f M Macrophage engulfment, 8 Marqibo, 6t Material surface properties, 25
29 adjustment, 27
29 mechanical effects, 27 physiological factors, 25
26 surface and interfacial energies, 26 surface charge, 26
27 surface morphology, 25 Matrix metalloproteinases (MMPs), 37 Mechanical strength, 18
19

Index

Mechanical stretching, 31 Mesoporous bioceramics, 128 Methacrylated alginate sponge-like cryogel, 177
178 Methocel K100M/Carbopol 974 P NF, 149 Microbicide-loaded paste core, 120 Microcapillaries, 47
48 Microchannels, 47
48 Microdevices delivery systems, 173
175 Microemulsions, 45 Microfluidics (MF), 9
11, 31, 47
48 devices, 47
48 mixing techniques, 77
78 Micrometer-scale disks, 31 Microstereolithography (MSTL), 49 Micro vs. nanoparticles, 73
75 Monoclonal antibody (mAB)-based NPs, 4t Monodisperse liposomes, 77
78 Mononuclear phagocyte system (MPS), 30 Morphone sulfate, 6t MPS. See Mononuclear phagocyte system (MPS) Multifunctional SMART nanotechnology, 11 Multifunction polymer, 8 Multiple emulsions, 45
46 Multipurpose prevention technologies (MPT), 118 Multivalent glucose-binding proteins, 207 Multiwalled carbon nanotube (MCNT), 21 N Naltrexone hydrochloride, 46 Nanocapsules, 81
82 Nanocarriers, 19 fabrication, 9
11 Nanoemulsions, 45 Nanogels, 84
85, 96 Nanomedicine, 72
73 Nanoparticles (NPs), 7
9. See also Drug carriers solid lipid nanoparticles (SLNs), 78
80, 78f Nanoprecipitation, 82 Nanospheres, 25
26, 81
82 Nanovesicular systems, 82 Nasal and pulmonary routes, 19 Nasal irritants, 188 National Institute on Drug Abuse (NIDA), 185

235

Natural polymers, 22
23 NC-6004, 96
97 NDDS. See Novel drug delivery system (NDDS) Nonbiodegradable polymers, 21
22, 28 Nondegradable ocular implants, 121
122 Nondegradable polymers, 20
21, 111
115, 112t Nondegradable subdermal implants, 115
116 Nondegradable vaginal rings, 116
121 Nonhealing wounds, 206 Nonionic surfactant, 44
45 Nonnucleoside reverse-transcriptase inhibitor IQP-0528, 120
121 Norplant system, 21, 115 Novel drug delivery system (NDDS), 1 nanocarrier fabrication, 9
11 SMART extended release drug delivery system, 9 SMART nanocarrier-based drug delivery systems, 2
9 NuSil (United States), 114
115 O Ocular drug delivery, 45 Ocusert, 20
21, 121
122 Oil-in-water emulsion, 79 Oncogels, 164
165 Oral antidiabetic molecules, 206 Oral drug absorption, 60 Oral hypoglycemics, 206t Ortho Pharmaceutical Corporation, 119
120 Osmotic-controlled drug delivery, 35
36 Osmotic systems, 9 OX26 antibody, 92 Oxaydo immediate-release tablets, 193 Oxycodone, 197 Oxygen reoxidizes, 207 P Paclical, 96
97 Paclitaxel, 10
11, 158
159, 168 Particle size, effect of, 30 Passive cancer targeting, 36
37 Passive targeting, 2 PBA. See Phenylboronic acid (PBA) PEGylated immunoliposomes, 9 PEGylated liposomes, 5

236

PEGylation, 9 Pellet-type implants, 115
116 Peptide-based NPs, 4t Peptide dendrimers, 6 PerioChip (chlorhexidine digluconate for gingivitis), 126 Permeability peptide drugs, 29 PEVA. See Polyethylene vinyl acetate (PEVA) Pharmaceutical foams, 44 Pharmaceutical suspensions, 43
44 Phase-change jet printing, 51 Phenylboronic acid (PBA), 207, 216
219, 217f, 218f advantages and limitations, 219
221 drug delivery systems, 220t nanostructures containing, 218f Phosphate buffer solution, 55 Phospholipids, 4
5, 77 Photopolymerization, 49
50 Photosensitive nanomaterial, 176 Photothermal therapy, 167, 172
173, 176 pH-responsive polymers, 7
8, 33 pH-sensitive polymers, 33, 167
168 Physical stimuli, 32 Physiologically based pharmacokinetic (PBPK) models, 59
60 Physiological stimuli, 2 Phytosomes, 4
5 Piezoelectric head velocity waves, 155 PK-Sim, 59
60 Plasma treatment, 28 Plastinating injecting unit (PIU), 46 Poly(acrylic)-based PBA polymeric systems, 219 Polyanionic electrolytes, 208 Polyanions, insoluble divalent metal salts, 24 Polycaprolactone (PCL), 113, 165, 176 Polycarbonate-based TPUs, 113 Polydimethylsiloxane (PDMS), 21 Polyelectrolytes, 207
208 Polyester-based TPUs, 113
114 Polyether-based TPUs, 113 Polyethylene glycol (PEG), 75 chain, 5 Polyethylene vinyl acetate (PEVA), 111
113 -based subdermal implants, 115
116 -based vaginal rings, 118
119

Index

Poly-HEMA-based hydrogel particles, 21
22 Polylactic acid (PLA), 10, 165 Poly (lactic-co-glycolic) (PLGA), 7, 126, 165 Polymer-based nanocarriers, biomedical applications, 92
98 Polymer-coating, 82 Polymeric micelles, 5
6 and polymersomes, 86
89 Polymeric nanosphere, 81
82, 81f Polymeric NPs, 7, 81f, 92
96 Polymer polydimethylsiloxane, 114
115 Polymers, 17 based drug carriers dendrimers, 82
84 nanogels, 84
85 nanospheres and nanocapsules, 81
82 polymeric micelles and polymersomes, 86
89 shape, 31 shrinkage, 47 size, 30
31 Polymersomes, 86
89 Polyols, 113 Polypropylene, 20 fumarate-based materials, 49
50 Polysaccharide coatings, 35
36 Polytetrafluoroethylene (PTFE)-like structure, 28 Polyvinyl acid (PVA), 218 Polyvinyl alcohol (PVA), 20 Poly(vinyl alcohol) (PVOH), 196
197 Polyvinylpyrrolidone, 20 Porous bioceramics, 128 Precirol-based SLN, 91
92 Prednisolone, 45 Prescription drug abuse, 185 methods, 187f Prescription drugs susceptible to abuse, 186f Pressure-assisted three-dimensional printer, 148, 149f Pressure-controlled colon delivery capsules, 35
36 Pressure-force extrusion, 148 Printed microorgan, 150 Progering, 116
117 Progestasert, 21 Prostate-specific membrane antigen, 4t Pulmonary DDSs, 19

Index

R Rate-limiting step, 29 Receptor-mediated endocytosis, 31 Recrystallization, 53 Reservoir-type nanodevices, 82 Reservoir vaginal ring, 116
117 Responsive to stimuli targeting, 4 Reticuloendothelial system (RES), 5 Reverse phase evaporation method, 77
78 Rifampicin-loaded PEG 2 PCL polymersomes, 97 Routes of abuse, 185
187 S Saline ions, 170
171 Scanning electron microscopy (SEM), 52 SE-based subdermal implant, 115 Selective laser sintering (SLS) technology, 50 Self-assembly, 30
31, 86 Self-healing hydrogels, 170
171 Self-regulated insulin release systems, 212 SFF. See Solid freeform fabrication (SFF) Shih-Etsu Silicones (Japan), 114
115 Silicone elastomer (SE), 111
113 Silicones, 20, 114 Simcyp, 59
60 Site selectivity, 2 Site-specific proteins and antibodies, 47
48 SL. See Stereolithography (SL) SLNs. See Solid lipid nanoparticles (SLNs) SLS technology. See Selective laser sintering (SLS) technology “SMART” dendrimers, 6 SMART extended release drug delivery system, 9 Smart materials, formulation design in drug delivery, 31
35 field-responsive polymers, 33
34 pH-responsive polymers, 33 temperature-responsive polymers, 32
33 SMART nanocarrier-based drug delivery systems transport mechanisms, 2
4 active targeting, 2
3 passive targeting, 2 responsive to stimuli targeting, 4 types, 4
9 dendrimers, 6

237

liposomes, 4
5 nanoparticles, 7
9 polymeric micelles, 5
6 “Smart” polymeric hydrogel systems, 18 Smart polymers, 31
32 Sodium alginate, 18 Soft lithography, 31 Solid foams, 44 Solid freeform fabrication (SFF), 48
49 bioplotting, 51 fused deposition modeling (FDM), 50 phase-change jet printing, 51 selective laser sintering (SLS) technology, 50 stereolithography (SL), 49
50 three-dimensional printing (3DP), 49 Solid lipid nanoparticles (SLNs), 78
80, 78f Soya lecithin-based SLNs, 91
92 Spritam, 156
157, 159
160 Stability and biodegradability, 53 Stealth effect, 5
6, 9 Stearic acid, 91
92 Stereolithography (SL), 49
50 apparatus, 147 Stimuli-responsive nanocarriers, 99 Stimuli-sensitive polymers, 32 Subcutaneous insulin injection, 207 Subdermal implants, 111, 115
116 Sulfopropyl acrylate, 194 Superabsorbent, 194 Surface degradation/erosion, 29 Surface-eroding system, 122 Suspensions, 43
44 Sustained release, 115
118 Swellable polymers, 194 Synthetic polymers, 23
24, 123
124, 168
169 T Tampering, 185
187 Target cites, in drug delivery brain targeting, 35 cancer tissues targeting, 36
37 colon targeting, 35
36 Taxus, 168 Tecto dendrimers, 6 Temperature-responsive polymers, 32
33 Temperature-sensitive drug molecules, 79 Template-assisted self-assembly, 31

238

Testosterone, 45 Testosterone therapy for premenopause, 116
117 Thermal oxidation, 28
29 Thermoplastic, 191
192 Thermoplastic biopolymers, 172
173 Thermoplastic polyurethane (TPU), 111
113 Thermoresponsive DDSs, 167
168 Thermoresponsive polymers, 165 Thermosensitive polymers, 32
33, 167
168 Thermosoftening polymers, 191
192 Thomsen 2 Friedenreich carbohydrate antigen, 4t Three-dimensional (3D) printing, 49 fused deposition modeling, 151
154 future perspectives, 159
160 inkjet printing, 154
159 drop-on-drop inkjet printing, 157
159 drop-on-solid inkjet printing, 155
157 pressure-assisted systems, 148
151 Thrombotic microangiopathy, 193
194 T-junction, 10 TPU. See Thermoplastic polyurethane (TPU) Transdermal DDSs, 19
20 Transferomes, 4
5 Transferrin, 4t -based NPs, 4t receptors, 4t Transmission electron microscopy (TEM), 52 Trimethylene carbonate-based materials, 49
50

Index

Tumor necrosis factor 2 related apoptosisinducing ligand (TRAIL), 169
170 Two-dimensional (2D) models, 55
57 U Ultracentrifugation, 82 Ultraviolet ozone (UVO), 27
28 Ultraviolet (UV) laser beam, 49 United States Pharmacopeia (USP), 54
55 V Vaginal microbicides, 117
118 Verteporphin, 6t Vincristine, 6t Vinyl acetate (VA) monomers, 111
113 Visudyne, 6t W Wacker (Germany) resins, 114
115 Water-insoluble drugs, 5
6 Water-insoluble polymers, 24 Water-repellent, 164
165 X Xartemis extended-release tablets, 193 X-ray powder diffraction, 52
53 Z ZipDose technology, 156
157 Zohydro extended-release capsules, 193 Zoladex (goserelin acetate for endometriosis and prostate cancer), 126