Biopolymer-Based Composites. Drug Delivery and Biomedical Applications [1st Edition] 9780081019153, 9780081019146

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Biopolymer-Based Composites. Drug Delivery and Biomedical Applications [1st Edition]
 9780081019153, 9780081019146

Table of contents :
Content:
Related titles,Front Matter,Copyright,List of contributors,Editors’ biographiesEntitled to full text1 - Biocomposites in therapeutic application: Current status and future, Pages 1-29, Sabyasachi Maiti, Sougata Jana, Subrata Jana
2 - Redox-responsive hydrogels, Pages 31-60, Weiren Cheng, Ye Liu
3 - Stimuli-responsive guar gum composites for colon-specific drug delivery, Pages 61-79, Sougata Jana, Sabyasachi Maiti, Subrata Jana
4 - Biopolymer-based nanocomposites for transdermal drug delivery, Pages 81-106, Rakesh K. Tekade, Rahul Maheshwari, Muktika Tekade
5 - Composites of hydrogels and nanoparticles: A potential solution to current challenges in buccal drug delivery, Pages 107-138, Sandra J. Morantes, Diana M. Buitrago, José F. Ibla, Yenny M. García, Gloria I. Lafaurie, Jenny E. Parraga
6 - Biocomposites in ocular drug delivery, Pages 139-168, Sabyasachi Maiti, Sougata Jana
7 - Dendrimers: Smart nanoengineered polymers for bioinspired applications in drug delivery, Pages 169-220, Keerti Jain
8 - Nanoparticles for tumor targeting, Pages 221-267, Ting Jiang, Kai Jin, Xianpping Liu, Zhiqing Pang
9 - Bioinspired nanotheranostics for cancer management, Pages 269-288, Rajeev Sharma, Nishi Mody, Suresh P. Vyas
10 - Biopolymers for gene delivery applications, Pages 289-323, Lucimara G. de La Torre, Caroline C. Sipoli, Aline F. Oliveira, Ismail Eş, Amanda C.S.N. Pessoa, Micaela T. Vitor, Franciele F. Vit, Thays F. Naves
11 - Biomedical and drug delivery applications of functionalized inorganic nanomaterials, Pages 325-379, Ayan K. Barui, Rajesh Kotcherlakota, Vishnu S. Bollu, Susheel K. Nethi, Chitta R. Patra
12 - Chitosan/carbon-based nanomaterials as scaffolds for tissue engineering, Pages 381-397, P.R. Sivashankari, M. Prabaharan
Index, Pages 399-407

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Woodhead Publishing Series in Biomaterials

Biopolymer-Based Composites Drug Delivery and Biomedical Applications

Edited by

Sougata Jana Sabyasachi Maiti Subrata Jana

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 © 2017 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. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-08-101914-6 (print) ISBN: 978-0-08-101915-3 (online) For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisition Editor: Gwen Jones Editorial Project Manager: Charlotte Cockle Production Project Manager: Poulouse Joseph Designer: Mark Rogers Typeset by TNQ Books and Journals

List of contributors

Ayan K. Barui Department of Chemical Biology, CSIR-Indian Institute of Chemical Technology, Hyderabad, India; Academy of Scientific and Innovative Research (AcSIR), Chennai, India Vishnu S. Bollu Department of Chemical Biology, CSIR-Indian Institute of Chemical Technology, Hyderabad, India; Academy of Scientific and Innovative Research (AcSIR), Chennai, India Diana M. Buitrago Universidad El Bosque, Bogotá, Colombia Weiren Cheng Institute of Materials Research and Engineering, A*STAR, Singapore Lucimara G. de La Torre State University of Campinas (UNICAMP), Campinas, Brazil Ismail Es¸ State University of Campinas (UNICAMP), Campinas, Brazil Yenny M. García Universidad El Bosque, Bogotá, Colombia José F. Ibla Universidad El Bosque, Bogotá, Colombia Keerti Jain National Institute of Pharmaceutical Education and Research (NIPER), Raebareli, India Sougata Jana Gupta College of Technological Sciences, Asansol, India Subrata Jana Indira Gandhi National Tribal University, Amarkantak, India Ting Jiang Fudan University, Shanghai, China Kai Jin Fudan University, Shanghai, China Rajesh Kotcherlakota Department of Chemical Biology, CSIR-Indian Institute of Chemical Technology, Hyderabad, India; Academy of Scientific and Innovative Research (AcSIR), Chennai, India Gloria I. Lafaurie Universidad El Bosque, Bogotá, Colombia Xianpping Liu Fudan University, Shanghai, China Ye Liu Institute of Materials Research and Engineering, A*STAR, Singapore

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List of contributors

Rahul Maheshwari BM College of Pharmaceutical Education and Research, Indore, India Sabyasachi Maiti Gupta College of Technological Sciences, Asansol, India Nishi Mody Dr. H.S. Gour Vishwavidyalaya, Sagar, India Sandra J. Morantes Universidad El Bosque, Bogotá, Colombia Thays F. Naves State University of Campinas (UNICAMP), Campinas, Brazil Susheel K. Nethi Department of Chemical Biology, CSIR-Indian Institute of Chemical Technology, Hyderabad, India; Academy of Scientific and Innovative Research (AcSIR), Chennai, India Aline F. Oliveira State University of Campinas (UNICAMP), Campinas, Brazil Zhiqing Pang Fudan University, Shanghai, China Jenny E. Parraga Tampere University of Technology, Tampere, Finland Chitta R. Patra Department of Chemical Biology, CSIR-Indian Institute of Chemical Technology, Hyderabad, India; Academy of Scientific and Innovative Research (AcSIR), Chennai, India Amanda C.S.N. Pessoa State University of Campinas (UNICAMP), Campinas, Brazil M. Prabaharan Hindustan Institute of Technology and Science, Chennai, India Rajeev Sharma Dr. H.S. Gour Vishwavidyalaya, Sagar, India Caroline C. Sipoli State University of Campinas (UNICAMP), Campinas, Brazil; Federal University of Technology – Paraná (UTFPR), Apucarana, Brazil P.R. Sivashankari Hindustan Institute of Technology and Science, Chennai, India Muktika Tekade Technocrats Institute of Technology Campus, Bhopal, India Rakesh K. Tekade National Institute of Pharmaceutical Education and Research (NIPER) – Ahmedabad, Gandhinagar, India Franciele F. Vit State University of Campinas (UNICAMP), Campinas, Brazil Micaela T. Vitor State University of Campinas (UNICAMP), Campinas, Brazil Suresh P. Vyas Dr. H.S. Gour Vishwavidyalaya, Sagar, India

Editors’ biographies

Prof. Sougata Jana is working at the Department of Pharmaceutics, Gupta College of Technological Sciences, under Maulana Abul Kalam Azad University of Technology, West Bengal, India. He is an MPharm (Pharmaceutics) from Biju Patnaik University of Technology, Odisha, India. He is engaged in research for 10 years and in teaching for 9 years. He qualified the Graduate Aptitude Test in Engineering (GATE) examination in the year 2005. He received “Gold Medal” from West Bengal University of Technology, Kolkata, for standing first at the undergraduate level. The Indian Pharmaceutical Association (IPA) Bengal branch, Kolkata, India, conferred on him the “M.N. Dev Memorial Award” for securing the highest marks in the state of West Bengal in the year 2005. He bagged the “Best Poster Presentation Award” at the 21st West Bengal State Science and Technology Congress, 2014, Burdwan, and the “Outstanding Paper Award” at the first Regional Science and Technology Congress, 2016, Bardhaman Division, organized by the Department of Science and Technology (DST), Government of West Bengal, and Bankura Christian College, West Bengal, India. He has published 30 research and review papers in different national and international peer-reviewed journals. He has edited four books published by Elsevier, Springer, and Pharmamedix India Publication Pvt. Ltd. More than 25 book chapters are also in his credit in Elsevier, Wiley VCH, CRC Press, and Taylor and Francis group. He is a reviewer of various international journals published by Elsevier, Wiley, Springer, Taylor and Francis, Dove Press, etc. He is a life member of the Association of Pharmaceutical Teachers of India (APTI) and Associateship (A.IC) from Institution of Chemists, India. He has successfully guided 16 postgraduate students for their research projects. He is working in the field of drug delivery science and technology, including modification of synthetic and natural biopolymers, microparticles, nanoparticles, semisolids, and interpenetrating network system for controlled drug delivery.

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Editors’ biographies

Dr. Sabyasachi Maiti is an MPharm, PhD, from Jadavpur University, Kolkata, India. He is currently working as Professor at the Department of Pharmaceutics, Gupta College of Technological Sciences, West Bengal, India. His research interest includes modification of natural polysaccharides and design of novel drug delivery systems. He has published more than 50 research and review papers in various international journals of repute. He has penned 20 book chapters for various international publishers and edited two books for CRC Press and InTech. He is also working as coeditor for an Elsevier book. He received research grants from national funding agencies such as All India Council for Technical Education (AICTE), Science and Engineering Research Board (SERB)-DST. He is associated with several professional bodies, such as IPA, APTI, and Indian Science Congress Association (ISCA). He is a fellow of the Indian Chemical Society and Institute of Chemists (India). He is a reviewer of Science Direct, Bentham Science, Springer, Informa Healthcare, and other journals.

Prof. Subrata Jana obtained his PhD in organic chemistry from Indian Institute of Engineering Science and Technology, Shibpur, India. His doctoral work was based on the design and synthesis of abiotic receptors for the recognition of biologically active neutral molecules and ions along with the development of synthetic methodologies. He also studied the recognition process in the solution phase and in the solid state. Overall, he has extensively studied the supramolecular behavior of the host–guest

Editors’ biographies

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interaction. Besides, he has worked on the development of synthetic methodologies for substituted heterocyclics, such as pyrimidines, naphthyridines, quinolone, and diazepines by exploiting the microwave protocol for green chemical synthesis. He moved to University of Victoria, Canada, to work with Dr. Fraser Hof on supramolecular and medicinal chemistry as a postdoctoral fellow, where he worked on the synthesis of different receptors targeting N-methylated protein residue along with anions. He then worked with Dr. Kenneth J Woycechowsky at University of Utah, USA, on protein engineering and enzyme catalysis as a postdoctoral research associate. He studied enzyme activity when the enzyme is encapsulated inside the capsid, which is a nanocarrier and an excellent delivery vehicle for important biological substrates, including drug molecules. Presently, he is working as Associate Professor at Department of Chemistry, Indira Gandhi National Tribal University, Amarkantak, Madhya Pradesh, India, and his current research focuses on the design and synthesis of artificial receptors for the recognition of anions, cations, and N-methylated protein residue. He is also working on biodegradable polymeric-based carrier systems for the delivery of drug molecules by collaboration with pharmaceutical scientists. So far he has published about 40 research papers in peer-reviewed international journals and contributed more than 10 book chapters in different books published by internationally renowned publishers. He also serves as editorial board member for the Journal of PharmaSciTech (ISSN: 2231-3788) and International Journal of Scientific and Engineering Research (ISSN: 2229-5518), as well as reviewer for International Journal of Biological Macromolecule (Elsevier) and Journal of PharmaSciTech and Current Pharmaceutical Design (Bentham).

Biocomposites in therapeutic application: current status and future

1

Sabyasachi Maiti1, Sougata Jana1, Subrata Jana2 1Gupta College of Technological Sciences, Asansol, India; 2Indira Gandhi National Tribal University, Amarkantak, India

1.1   Introduction From the health care perspective, biomaterials can be defined as the materials that possess some novel properties that make them appropriate to come into immediate contact with the living tissue without eliciting any adverse immune rejection reactions. The biomaterials can be divided into the following categories: (1) synthetic (metals, polymers, ceramics, and composites); (2) naturally derived (animal and plant derived); (3) semisynthetic or hybrid materials. The ensuing developments of these biomaterials have enhanced their utility in health care. Natural polymers are widely used in the areas of health care for the fabrication of drug delivery systems. As naturally derived biomaterials have limited mechanical strength, their applications as drug carriers are restricted. Therefore such materials are being modified chemically to improve their mechanical properties. Biomaterials for the delivery of drugs or bioactive molecules are showing newer developments with time. During fabrication of these biomaterials, a balance between the physical and mechanical properties together with minimal toxicity to host tissue must be maintained [1,2]. Complementing synthetic polymers, naturally derived biopolymers are on the way of becoming equally important. However, the biomaterials must be obtained with sufficient and reproducible purity and quality. To meet clinical needs in terms of biodegradation, drug loading, and drug release, natural biopolymers may be endowed with additional functional moieties. Because of their exceptional biocompatibility, biodegradability, and nontoxic products of degradation, natural polymer-derived materials have been extensively studied in biomedical engineering [3,4]. Biopolymer-based hydrogels have attracted a great deal of interest in tissue engineering and drug delivery applications [5]. Despite the known advantages and wide applicability of biomaterials, there are several limitations such as poor mechanical properties and low stability in aqueous environments that restrict their use for biomedical applications [6,7]. However, the hydrogels possess a wide variety of functional groups, including hydroxyl, amino, and carboxylic acid groups, which can further be cross-linked and conjugated with cell-targeting ligands. The naturally occurring biopolymers are polysaccharides, and typical examples are chitosan, hyaluronic acid, dextran, pullulan, and alginate. The covalent cross-linkers interconnect molecules, Biopolymer-Based Composites. http://dx.doi.org/10.1016/B978-0-08-101914-6.00001-6 Copyright © 2017 Elsevier Ltd. All rights reserved.

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Biopolymer-Based Composites

increase the molecular weight, and generally provide higher mechanical strength and improved stability. However, cross-linking may also decrease the degradability and the availability of functional groups in the cross-linked polymer and changes the rheology of the polymers, leading to subsequent processing difficulties and potential increase in cytotoxicity [6]. For a widely used cross-linker glutaraldehyde, up to 8% has been shown to be noncytotoxic [8]. Because individual polymers alone are not sufficient in providing biomaterials with the desired properties, the use of blends of different polymers has been suggested [9]. For instance, porous collagen/chitosan blend scaffolds were developed and treated with glutaraldehyde/genipin to improve their mechanical properties and stability [10,11]. Reduced swelling and decrease in enzymatic degradation were found for collagen and hyaluronic acid composite materials without any impact on the cell viability [12]. Similar to the binary blends, a ternary composition of collagen, hyaluronic acid, and poly (caprolactone) was used to generate sponge-like dense membranes using ultraviolet irradiation and carbodiimide coupling [13]. The hybrid cross-linking systems were able to preserve the native structure of collagen, and the inclusion of poly (caprolactone) provided the ability to control the degradation and mechanical properties. This made the sponges suitable for the development of wound dressings and periodontal membranes. Composite materials or composites are engineered materials made up of two or more constituent materials with significantly different physical or chemical properties, which remain separate and distinct on a macroscopic level within the finished structure. Thus composites are always heterogeneous [14]. In any composite, there are two major categories of constituent materials: a matrix (or a continuous phase) and a dispersed phase(s). The continuous phase is responsible for filling the volume as well as surrounding and supporting the dispersed material(s) by maintaining their relative positions. The dispersed phase(s) is (are) usually responsible for enhancing one or more properties of the matrix. Most of the composites target an enhancement of the mechanical properties of the matrix, such as stiffness and strength; however, other properties, such as erosion stability, transport properties (electrical or thermal), or biocompatibility might also be of great interest. This synergism produces properties that are unavailable from the individual constituent materials [15,16]. By controlling the volume fractions of the dispersed phase, the properties and design of composites can be varied and tailored to suit the necessary conditions. Higher volume fractions of reinforcement phases tend to improve the mechanical properties of the composites. The uniform distribution of the dispersed phase is also desirable, as it imparts consistent properties to the composite [15,16]. In general, composites might be simple (homogeneous dispersion of one dispersed phase throughout a matrix), complex (homogeneous dispersion of several dispersed phases throughout one matrix), graded (inhomogeneous dispersion of one or several dispersed phases throughout one matrix), or hierarchical. A hierarchical composite refers to those cases in which fine entities of either a simple or a complex composite are somehow aggregated to form coarser ones (granules or particles), which afterward are dispersed inside another matrix to produce the second hierarchical scale of the composite structure [17]. To prepare any type of composite, at least two different materials must be mixed. However, the interfacial strength among the phases is a very important factor, because lack of adhesion among

Biocomposites in therapeutic application: current status and future

3

the phases will result in an early failure at the interface and thus cause a decrease in the mechanical properties, especially the tensile strength [18]. Several types of interactions can exist among the components: strong covalent, coordination, ionic, and weak interactions, such as van der Waals forces, hydrogen bonds, and hydrophilic– hydrophobic balance [19]. Polymer nanocomposites are a new class of materials that are characterized by superior properties compared with macro- and micropolymer composites. They consist of a macromolecular matrix and nanofiller, which in most cases is a mineral or a metallic material (clay minerals, metal, and carbon nanotubes) [20]. The nanofillers are mostly used because of their high surface area, colloidal dimensions of their particles, and other characteristics [21]. The decoration of silver, gold, iron oxide nanoparticles with biomaterials are more effective in the field of health science for targeting the delivery of drugs and for biomedical applications, with a variety of roles like improving solubility, enhancing in vivo stability, and optimizing the biodistribution and pharmacokinetics of drugs [22,23]. Considering the importance and uniqueness of biocomposite systems, this chapter highlights the fundamentals and developments in this field under different heads such as stimuli-responsive composites, mucoadhesive composites, hydrogel and inorganic composites, and theranostic composites.

1.2  Composites and their therapeutic application 1.2.1   Stimuli-responsive composites Bionanocomposites are novel materials with drastically improved mechanical properties due to the incorporation of a small amount ( 1° amine >> 2° amine (formed) [12,22,23]. PEGylation of hyperbranched poly(amido amine)s can be performed in two different ways. One is to conjugate PEG chains in a one-pot two-step synthesis. After double molar BAC is reacted with AEPZ in methanol at 50°C for 6 days, α-amino-ω-methoxy-PEG with Mn 750 is added into the reaction solution at 60°C, and the reaction is left for another week. As shown in Fig. 2.2 [19], PEG is conjugated to the hyperbranched poly(amido amine) core via the Michael addition reaction of the primary amine of α-amino-ωmethoxy-PEG with the vinyl terminals. The other method of PEGylation is after reaction, the vinyl terminal group of hyperbranched poly(BAC2-AMPD1) is converted into amine via the reaction with excess AMPD in ethanol/dimethyl sulfoxide mixture [11]. The disappearance of the vinyl peaks at 125.5 and 130.5 ppm in Fig. 2.1(e) indicates the depletion of the vinyl ends by the reaction with AMPD. Scheme 2.1 shows that there are three possible reactions between the vinyl ends and AMPD, i.e., the reaction between the vinyl group and either the 2° amine (original) or the 1° amine and the AMPD cyclic reaction with two vinyl terminals. Because the reactivity sequence of the three amines of AMPD is as follows: 2° amine (original) > 1° amine >> 2° amine (formed) [11,12,23], when

Redox-responsive hydrogels

HN

N AEPZ

NH2

+

35

H N O

H N

S S

O

1:2 O

BAC

PEG-NH2

O N H

O

SH N

S

S

O

O N H

S

N H

N

Red ellipsoid H N

N O

S

S

H N

H N

NH2

n=16 or 17 Mn: 750

O O Green helical chains

O n

S Cyan rhombi

Poly(BAC2-AEPZ1)-PEG

Figure 2.2  Synthesis of the branched poly(BAC2-AEPZ1)-PEG. AEPZ, 1-(2-aminoethyl) piperidine; BAC, N,N-cystaminebis(acrylamide); PEG, poly(ethylene glycol). Reprinted with permission from Wu DC, Loh XJ, Wu YL, Lay CL, Liu Y. ’Living’ controlled in situ gelling systems: thiol-disulfide exchange method toward tailor-made biodegradable hydrogels. Journal of the American Chemical Society. 2010;132:15140–3. Copyright (2010) American Chemical Society.

excess AMPD is presented, the vinyl ends should react exclusively with the 2° amine (original) instead of the 1° amine. This is supported by the presence of the characteristic peak such as c2 at 29.4 ppm in Fig. 2.1(e). Meanwhile, the peaks corresponding to the reaction between vinyl and 1° amine such as a3 and h3 cannot be observed. Therefore, it is evident that most of the vinyl ends react with the 2° amine (original) of AMPD forming hyperbranched poly(BAC2-AMPD1) with dNH2 terminals, hyperbranched poly(BAC2-AMPD1)-NH2. PEG with molecular weight of 2000 is then conjugated to the dNH2 terminals via the formation of urethane bond as shown in Scheme 2.1. The molecular weight of hyperbranched poly(BAC2-AMPD1)-PEG is 125.6 ± 2.6 kDa obtained from Zimm plot in methanol. Thus each hyperbranched poly(BAC2-AMPD1) macromolecule is conjugated with ∼25 PEG chains.

2.2.2  Synthesis of thermoresponsive hyperbranched poly(amido amine)s The hydrogel precursor, thermoresponsive hyperbranched poly(amido amine)s, S-HPAA, is similarly prepared via the Michael addition polymerization of BAC and N,N-dimethyldipropylenetriamine (DMDPTA) [24]. A typical procedure is 1.0 mmol of DMDPTA and 1.0 mmol of BCA are added into 3 mL of methanol/water mixture (1/1 in volume). The reaction is performed at 45°C for 72 h to produce S-HPAA with a molecular weight of 9500 and polydispersity index of 1.2. The 13C nuclear magnetic resonance spectrum of S-HPAA verifies the hyperbranched structure. In the previous polymerization shown in Scheme 2.1, hyperbranched poly(amido amine)s are achieved with AMPD and double molar BAC, whereas for this polymerization, equal molar quantities of DMDPTA and BAC

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are used instead to yield poly(amido amine)s with hyperbranched structure. This is because the reactivity sequence of the three types of amines in DMDPTA differs from that of AMPD and AEPZ; therefore, the reaction pathways are different and different monomer molar ratios are applied.

2.3  Redox-responsive hydrogels 2.3.1  Redox-responsive poly(amido amine) hydrogels Redox-responsive poly(amido amine) hydrogels can be formed via intermolecular thiol-disulfide exchange reaction of PEGylated hyperbranched poly(amido amine)s. The gelation process can be activated, controllably terminated, and interrupted and reinitiated anytime [19]. In this way, the extent of cross-linking of the in situ hydrogels can be precisely controlled. Thiol-disulfide exchange is a pH-responsive reaction depending on the pKas of the involved thiols, and the reacting species are the thiolates [25]. Protonation of thiolates and deprotonation of thiols can be reversibly manipulated by adjusting the system pH, resulting in thiol-disulfide exchange with an “on/ off” function. So, the thiol-disulfide exchange reaction could be utilized to develop the gelling system from a branched precursor with at least three cross-linkable disulfide-containing branches. When 10 wt% poly(BAC2-APEZ1)-PEG solution (Fig. 2.3(a)) is basified to pH 12, a white loose hydrogel (Fig. 2.3(b)) is formed about 1 h later. For a clear demonstration of this sol–gel transition, rheological measurement is performed to obtain the storage modulus G′ and viscosity η (associated with dissipative modulus G″). Fig. 2.4 (a)

(b)

(c)

pH12

pH12

1h

24h

pH7

(b*)

pH7 pH12

Figure 2.3  Schematic representation of basification of branched poly(BAC2-AEPZ1)-PEG (BAP) solution and derived methodology for producing loose and compact hydrogels. (a) 10 wt% BAP solution, (b) loose hydrogel after basification for 1 h, (b*) inert neutralized loose hydrogel after 1 h of cross-linking, and (c) compact hydrogel after basification for 24 h. Reprinted with permission from Wu DC, Loh XJ, Wu YL, Lay CL, Liu Y. ’Living’ controlled in situ gelling systems: thiol-disulfide exchange method toward tailor-made biodegradable hydrogels. Journal of the American Chemical Society. 2010;132:15140–3. Copyright (2010) American Chemical Society.

Redox-responsive hydrogels

37

shows that the viscosity of the solution remained constant initially, but after about 24 min, an abrupt increase of both G′ and η is observed, indicating gelation of the polymer solution. The formation of stable loose hydrogel is further confirmed by the storage modulus G′ which is higher than the dissipative modulus G″ regardless of change of oscillatory frequency. The loose hydrogel shrinks further into a compact hydrogel (Fig. 2.3(c)) after 1 day of aging. However, there is no further shrinkage in the diameter of the compact hydrogel when the aging period is extended to 1 week. The transition from loose to compact hydrogels can be interrupted at any point of time by neutralization and restarted by rebasification (Fig. 2.3). The neutralized hydrogel (Fig. 2.3(b)) has good stability and can retain its shape for over half a year in mild aqueous 103 102

102

101

101

100 10–1 10–2

100

24 mins

Solution

η (Pas)

G' (Pa)

103

Slippage Hydrogel

10–1 0

10

20 Time (min)

30

40

Figure 2.4  Storage modulus G′ and viscosity η of 10 wt% BAP solution after basification to pH 12 as a function of time. Oscillatory frequency: 1 rad/s. Reprinted with permission from Wu DC, Loh XJ, Wu YL, Lay CL, Liu Y. ’Living’ controlled in situ gelling systems: thiol-disulfide exchange method toward tailor-made biodegradable hydrogels. Journal of the American Chemical Society. 2010;132:15140–3. Copyright (2010) American Chemical Society.

6.0 mm

5.2 mm

5.0 mm

4.2 mm

4h

6h

8h

24 h

Figure 2.5  Variation of hydrogel shrinkage during basification of 10 wt% BAP solution as a function of time. Reprinted with permission from Wu DC, Loh XJ, Wu YL, Lay CL, Liu Y. ’Living’ controlled in situ gelling systems: thiol-disulfide exchange method toward tailor-made biodegradable hydrogels. Journal of the American Chemical Society. 2010;132:15140–3. Copyright (2010) American Chemical Society.

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environments. Fig. 2.5 illustrates the macroscopic shrinkage of the hydrogels during basification and reveals that an increased aging time leads to increased gel shrinkage. The formation of hydrogel with poly(BAC2-AEPZ1)-PEG is so robust that a polymer solution of 0.5 wt% is sufficient to form the hydrogel. However, to completely encapsulate the water in the initial solution, 2.0 wt% polymer solution is required. On the basis of the thiol-disulfide exchange mechanism, the formation and structures of the loose and compact hydrogels are explained, and the “living” controlled process of hydrogel formation is also demonstrated (Fig. 2.6). The terminal thiols of the hyperbranched poly(BAC2-AEPZ1)-PEG (represented by purple cones in Fig. 2.6-I), which are formed from the reduction of disulfide bonds during the preparation and purification procedures, are converted to thiolates (represented by blue cones in Fig. 2.6-II) at pH 12. Then, these thiolates trigger disulfide exchange reactions and bond two hyperbranched polymers by the formation of cross-linked disulfide bonds (represented by blue rhombi) and the release of reassembled PEG shells. The PEG shell, which has decent solubility in water, is a good leaving group, so the driving force is toward the formation of stable cross-linked products through the release of PEG shells. After the system is aged for an hour, the loose hydrogel is formed when all hyperbranched polymers are bonded together by in situ buildup of cross-linked disulfide bonds. If the system is allowed to age for another day, further disulfide exchange and hydrophobic interactions drive the integration of all the individual polymer cores to form a large hydrophobic core (purple cylinder built from a lot of red ellipsoids linked by cyan and blue rhombi, Fig. 2.6(c)), and at the same time releasing most of the reassembled PEG chains into the aqueous phase. The highly hydrophobic cylindrical core expels a large amount of encapsulated water and shrinks into a compact gel. The depleting PEG shells and the highly dense and hydrophobic structure of the compact gel seriously limit the access of attacking thiolates, thus inhibiting further disulfide exchange reaction. This explains the lack of shrinkage observed in the compact hydrogel after 1 day. The thiol-disulfide exchange reaction can be interrupted and reinitiated anytime during the transformation from loose to compact hydrogels by tuning the pH of the system. Under neutralization, the reactive thiolates can be deactivated to thiol and the resultant hydrogel is highly stable. By simply adjusting the pH of the system back to 12, the thiols can be reactivated to form thiolates to continue the thiol-disulfide exchange. This demonstrates the “living” capability of the hydrogel system. As shown in Fig. 2.7, hyperbranched poly(BAC2-AEPZ1)-PEG exhibits low cytotoxicity and because of the presence of ample disulfide bonds the polymer is also biodegradable by thiol-containing species such as cysteine and GSH. Fig. 2.8 shows that both loose and compact hydrogels do not undergo degradation in phosphate buffered saline (PBS) or PBS with 10 μM dithiothreitol (DTT). However, in the presence of 10 mM DTT, the loose hydrogel formed with 1 h of cross-linking takes 3 h to degrade completely, whereas total degradation of the compact hydrogel formed after 24 h of cross-linking takes 80 h. Both doxorubicin (DOX) and paclitaxel can be loaded into the redox-responsive hydrogel of poly(BAC2-AEPZ1)-PEG. Fig. 2.9 illustrates the tailor-made hydrogels allow a fine modulation of drug release to facilitate preparation of smart drug vectors that are stable in the extracellular medium and decomposable in the intercellular environment to trigger drug release [26].

Redox-responsive hydrogels

39

Solution

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Figure 2.6  Schematic representation of the formation and structures of loose and compact hydrogels. (a) 10 wt% BAP solution, (b) loose hydrogel after 1 h of cross-linking, and (c) compact hydrogel after 24 h of cross-linking. Reprinted with permission from Wu DC, Loh XJ, Wu YL, Lay CL, Liu Y. ’Living’ controlled in situ gelling systems: thiol-disulfide exchange method toward tailor-made biodegradable hydrogels. Journal of the American Chemical Society. 2010;132:15140–3. Copyright (2010) American Chemical Society.

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Biopolymer-Based Composites BHK21, 1 day extraction BHK21, 3 day extraction HEK293, 1 day extraction HEK293, 3 day extraction L929, 1 day extraction L929, 3 day extraction L929, 3 day extraction, 10 mM DTT

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Figure 2.7  Cytotoxicity of the loose hydrogel after 4 h of cross-linking in BHK21, HEK293, and L929 cells after incubation for 1 and 3 days (values were presented as mean ± standard deviation, n = 4) [19].

2.3.2  Other redox-responsive hydrogels To afford redox responsivity in hydrogels, the most obvious approach is the use of disulfide-containing cross-linkers, and generally, there are several methods to achieve it. One is by cross-linking precursor polymers with disulfide linkages [27–29]. For these hydrogels, loading of small-molecular-weight active species, which depends on either physical encapsulation or noncovalent interactions, can be performed during or after the cross-linking process. However, for macromolecular compounds, unless they can interact with the hydrogel surface noncovalently, it is more advisable to load them during cross-linking, as these molecules may be too large to diffuse and penetrate the polymer network. Covalent attachment of loads to the hydrogels via disulfide bonds is another popular way to ensure high specificity and loading capacity. For example, it was reported that lysozyme can be covalently bound along the redox-cleavable crosslinkers of dextran hydrogels and still retain its enzymatic activity [28]. As a result, in the presence of GSH, which reduces the disulfide bonds, the immobilized lysozyme is rapidly released. Redox-responsive hydrogels can also be easily prepared in a one-pot synthesis. Typically, the monomers, cross-linkers, and initiators such as photoinitiators [30,31] are all added in one reaction to kick off polymerization and hydrogel formation simultaneously. The loading of active species in these types of hydrogels can be done before or after polymerization [32], and each method has its associated limitations. Owing to the potentially harsh and reactive conditions of polymerization, such as free radical polymerization, it is less favorable to load fragile active species like proteins before the reaction. Moreover, loading before polymerization may prohibit the purification of the

Redox-responsive hydrogels

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Figure 2.8  (a) Amounts of the hydrogels remaining (i) after 1 h of cross-linking in 1× phosphate buffered saline (PBS), (ii) after 24 h of cross-linking in 1× PBS, (iii) after 1 h of cross-linking in 1× PBS containing 10 μM dithiothreitol (DTT), and (iv) after 24 h of cross-linking in 1× PBS containing 10 μM DTT at 37°C as a function of time. (b) Amounts of the hydrogels remaining (v) after 1 h of cross-linking in 1× PBS containing 10 mM DTT and (vi) after 24 h of cross-linking in 1× PBS containing 10 mM DTT at 37°C as a function of time (values were presented as mean ± standard deviation, n = 3) [19].

resultant hydrogels and thus lead to the retention of potentially cytotoxic compounds such as initiators in the hydrogels. On the other hand, loading after polymerization faces the same issue highlighted in the previous paragraph; macromolecular loads can be too large to enter the polymer network and thus can only be adsorbed on the surface of the hydrogels, which is not as desirable, as they may be more prone to attacks from

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Figure 2.9  a) Cumulative amounts of doxorubicin (DOX) released from the DOX-loaded hydrogels: (i) after 1 h of cross-linking in 1× PBS, (ii) after 24 h of cross-linking in 1× PBS, (iii) after 24 h of cross-linking in 1× PBS containing 10 mM dithiothreitol (DTT), and (iv) after 24 h of cross-linking in 1× PBS containing 10 μM DTT at 37°C as a function of time. (b) Cumulative amounts of paclitaxel (PTX) released from the PTX-loaded hydrogels: (i) after 1 h of cross-linking in 1× PBS, (ii) after 24 h of cross-linking in 1× PBS, (iii) after 24 h of cross-linking in 1× PBS containing 10 mM DTT, and (iv) pure PTX in 1× PBS at 37°C as a function of time (values were presented as mean ± standard deviation, n = 3) [19].

the immune system. Therefore, to prepare hydrogels in this approach, milder polymerization techniques are usually preferred. As such many redox-responsive hydrogels were designed to deliver a range of active compounds such as small-molecular-weight drugs [32–34] and macromolecules [35]. In particular, a study demonstrated that

Redox-responsive hydrogels

43

amino acid–based hydrogels are applicable for oral drug delivery [36]. To overcome the major barrier of oral delivery, the hydrogels collapse and form compact structures to protect the drugs at acidic pH in the stomach. Once the hydrogels reach the intended target, the physiological pH and reductive environment can trigger rapid release of the loaded drugs. Being extremely hydrophilic, biodegradable, and capable of delivering actives upon stimulation to facilitate cellular activities, redox-responsive hydrogels can provide a very conducive environment for cell regeneration and proliferation; thus they are attractive candidates in tissue engineering. One of such studies is the design of disulfide-cross-linked hyaluronan–gelatin hydrogels, which were fabricated by cross-linking thiolated hyaluronan and gelatin derivatives in air, to mimic the extracellular matrix and the hydrogels [37]. Their observations showed that the presence of hyaluronan slowed the degradation of the hydrogels, which can be altered to match tissue regeneration by collagenase both in vitro and in vivo, whereas the increase in gelatin improves the cell attachment and proliferation. Another similar work that used thiolated four-armed PEG as the hydrogel precursor showed more rapid release of bovine serum albumin upon GSH exposure, and when loaded with recombinant human bone morphogenetic protein-2, the hydrogels promoted bone formation and favorable blood cell growth [38,39]. In another similar area, some hydrogels have demonstrated self-healing capability, which is essential to the development in organ repair. Harada, Nakahata, and coworkers have contributed significantly in this area using host–guest polymers [40], and one of their works involved redox-responsive self-healing hydrogels that used cyclodextrins and ferrocene as host–guest to achieve redox sensitivity [41,42]. The spreading of GSH solution on the cut surfaces of two hydrogel pieces allowed reduced ferrocene to insert back into the cavity of cyclodextrin and spontaneously adhere to each other. Another acylhydrazone- and disulfide-based hydrogel is able to self-heal under basic conditions because of the facilitated disulfide exchange reaction by high pH [29]. Encapsulation of cells within the hydrogel provides more uniform cell distribution favorable for growth and proliferation than seeding cells on the hydrogel surface. However, to encapsulate cells in hydrogels, requirements such as purely water-based process and mild reaction conditions must be strictly adhered to in order to ensure cell viability and functionality after encapsulation. In one study, horseradish peroxidase was used to cross-link thiolated PEG in the absence of hydrogen peroxide to form redox-responsive hydrogels [43]. This hydrogel could degrade readily under reductive environment and the viability of their recovered cells was over 98%. Another group used water-based polymerization of monomer, carboxybetaine methacrylate, and cross-linker N,N′dimethacryloylcystine to encapsulate cells and form the redox-responsive hydrogels [44] without damaging the cells. To prepare desirable three-dimensional scaffolds is highly challenging, and in one study, tubular scaffold, which is structurally similar to blood vessels, was successfully fabricated. These mechanically strong hydrogels synthesized via photopolymerization are able to self-roll upon exposure to GSH [30]. Owing to the uneven redox-induced shrinkage stress on both the surfaces of the bilayered hydrogels, the hydrogels curl up to form a tubular structure and encapsulate cells at the same time.

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The ability to gel in situ is one of the reasons why hydrogels are ideal biomaterials, and this is especially easy to achieve in redox-responsive hydrogels. By simply oxidizing thiols-containing polymers, hydrogels can be formed in situ with any desirable shape and size. Two groups have independently designed injectable redox-responsive PEG- and poly(aspartic acid) (PASP)-based hydrogels, respectively. PEG-based hydrogels can be formed within 1 min by mixing a predetermined amount of PEG-SH and H2O2 in vitro [38,39]. Furthermore, when the hydrogel formation is studied in vivo, gels harvested from sacrificed rats at 30 min postinjection still remain intact and smooth, indicating fast in situ gelation. The swelling ratio and degradation rate of the PEG-based hydrogels are dependent on the polymer concentration. At low concentrations of GSH, the PEG-based hydrogels experience gradual redox-induced degradation and sustainable protein release. The study also demonstrated that the PEG-based hydrogels functionalized with cell-targeting peptides and osteoinductive cytokines are able to support cell ingrowth and promote bone regeneration. On the other hand, the PASP-based hydrogels are capable of undergoing reversible redox-induced sol–gel transition. Even after several cycles of transformation, the hydrogels still retain their initial mechanical properties [34,45]. Polymerization that occurs in mild conditions and does not produce side products can also be employed to induce in situ gelation. An example is the phosphoramidate–hyaluronan hybrid-based hydrogels [46]. These hydrogels are prepared by in situ Michael addition polymerization of acrylated hyperbranched polyphosphoramidate (HPPAE-AC) and thiolated hyaluronan (HA-SH), and the gelation can take place as quickly as 4 min. As expected, the HPPAE-HA-based hydrogels show DTT-dependent degradation rate.

2.4  Redox-responsive nanogels/microgels Nanoparticles, dubbed as the “magic bullet,” were the mainstream of scientific research on drug delivery for the past two to three decades. Nanogels, which are hydrogels shrunk to nanoscale, are ideal for drug delivery because of the combined properties of hydrogels and nanoparticles, such as hydrophilicity, protection of drugs, lengthening blood circulation time, and passive/active targeting capability [47,48].

2.4.1  Redox-responsive poly(amido amine) nanogels/microgels from emulsion with surfactants As hydrogels of hyperbranched poly(BAC2-APEZ1)-PEG can be easily molded into different shapes and sizes, with appropriate templates, nanogels of hyperbranched poly(BAC2-AEPZ1)-PEG can be fabricated effortlessly as shown in Fig. 2.10 [20]. Emulsion has been frequently used to prepare uniform nanoparticles. Here a span80/ tween80-stabilized water-in-oil (W/O) emulsion with water/decane can assist in the formation of nanogels of hyperbranched poly(BAC2-APEZ1)-PEG. In these W/O emulsions, under basification, the polymers in the surfactants stabilized dispersed water phase undergo in situ gelation to form loose hydrogel particles with the size close to the diameter of the water droplets. Therefore, by tuning the surfactants

Redox-responsive hydrogels

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Figure 2.10  (a) Schematic illustration of the core/shell separation process, dissociation of the shells, and cross-linking of the cores and (b) schematic depiction of the synthetic approach to controlled formation of (multilayered) hydrogel particles. W/O, water in oil. Reprinted with permission from Zhang J, Yang F, Shen H, Wu D. Controlled formation of microgels/nanogels from a disulfide-linked core/shell hyperbranched polymer. ACS Macro Letters. 2012;1:1295–9. Copyright (2012) American Chemical Society.

concentration to vary the size of water droplets formed, hydrogel particles of different dimensions can be easily obtained. Furthermore, the use of cyclohexane instead of decane as the oil phase can also generate hydrogel particles in the micron range. Similarly, the mechanism of gelation is the pH-dependent thiol-disulfide exchange reaction accompanied by the dissociation of hydrophilic PEG shells. If the aging is allowed to continue, further dissociation of the PEG shells leads to shrinkage of the loose hydrogel particles into smaller and more compact ones. As mentioned previously, through neutralization during the transition from loose to compact hydrogel particles, the conversion of thiolates to thiols can halt the thiol-disulfide exchange reaction and

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preserve the size of the hydrogel particles. Another approach to control the size of the hydrogel particles is to vary the polymer concentration. At the early stage, regardless of polymer concentration, the dimension of loose hydrogel particles is determined by the size of the water droplets generated. However, if cross-linking continues, a high polymer concentration facilitates the formation of larger hydrogel particles as each water droplet contains more polymer cores. The gelling duration affects not only the size of the hydrogel particles formed but also the degree of cross-linking. As the duration of gelling increases, the swelling ratio of the hydrogel particles decreases because the greater degree of cross-linking limits the distension of the polymer networks. Multilayer hydrogel particles are gaining much interest because of their unique shell/core structures that have separate properties and respond to different stimuli. However, preparation methods to independently adjust the different layers with differing structures and properties are usually highly complex. Using the aforementioned controlled in situ gelation method in association with a seed emulsion technique, each on-demand layer can be produced, and controlled formation of multilayered hydrogel particles with flexibly designable structures and properties of the respective layer can thus be achieved. To prepare these multilayered hydrogel particles, first, suitable hydrogel particles that are utilized as seeds are fabricated with a predetermined gelling time. Then, fresh polymer solution is added to the W/O emulsion, with the preprepared hydrogel particles serving as the core for the subsequent polymer deposition and cross-linking of the shell. Interestingly, this approach even allows the deposition of the third layer. Fig. 2.11 shows the transmission electron microscopy (TEM) images of the multilayered hydrogel particles, with each layer having different cross-linking density. The hydrogel particles are subjected to redox-induced degradation when exposed to GSH. In low GSH concentration, which mimics the redox condition of the extracellular matrix, the hydrogel particles are stable and no obvious degradation is observed even after 2 days of incubation. On the other hand, in GSH concentration similar to the intracellular redox environment, obvious degradation of the hydrogel particles is observed. Hydrogel particles with longer gelling time and higher cross-linking density degrade less rapidly than those aged for a shorter duration, which have lower cross-linking density. When the multilayered hydrogel particles also faced a similar reductive condition, it was noted that the disassociation of the core/shell happens more readily than the degradation of the individual core and shell layer. This is because the disulfide linkage across the core/shell interface is not as robust as that located in the core or shell layer. Furthermore, the study also shows that the shell layers tend to degrade much faster than the core; therefore the system is promising for further development of hierarchically controlled drug delivery.

2.4.2  Redox-responsive poly(amido amine) nanogels/microgels from emulsion without surfactants The use of nanogels/microgels prepared by most of the reported approaches may not always favor application in the biomedical field. This is because majority of these methods are complex and frequently require the use of potentially cytotoxic surfactants,

Redox-responsive hydrogels

47

(a)

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Figure 2.11  (a–f) Transmission electron microscopy (TEM) images of core/shell doublelayered nanogels after different cross-linking times: (a) 6 h/1 h, (b) 6 h/2 h, (c) 10 h/2 h, (d) 10 h/4 h, (e) 24 h/4 h, and (f) 24 h/8 h. (g, h) TEM images of (24 h/8 h/2 h) core/double shells triple-layered nanogels. Reprinted with permission from Zhang J, Yang F, Shen H, Wu D. Controlled formation of microgels/nanogels from a disulfide-linked core/shell hyperbranched polymer. ACS Macro Letters. 2012;1:1295–9. Copyright (2012) American Chemical Society.

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[20,49–52] cross-linkers [53–56] or initiators, catalysts, and monomers for reaction or polymerization [55,57–62]. As a result, additional exhaustive purification steps are needed to remove these surfactants, by-products of cross-linking, initiators, catalysts, and unreacted monomer before the nanogels are safe for use. Drugs can be loaded in the nanogels either during or after formulation. For methods that load drugs during formulation, the preparation technique usually involves mild reaction or polymerization to avoid damaging the drugs loaded [59,63]. As a result of this challenge, a large portion of the studies focuses on loading drugs after formulation, and this process of loading drugs within or on the nanogels can be highly inefficient and limited. One of the major problems is to load macromolecular drugs within the nanogels. The dense polymer network may prevent large drugs from penetrating into the nanogels. Therefore, very often large drugs can only be adsorbed on the nanogel surfaces, subjecting them to potential degradation by the environment. Another limitation to load drugs after formulation is that certain form of interactions, such as electrostatic and hydrophobic interactions, van der Waals force, or hydrogen bonding, between the polymers and drugs must exist for loading to be possible. This greatly limits the range of drugs that can be loaded in a particular nanogel system, e.g., neutrally charged active compounds have difficulties loading into positively charged nanogels. Therefore, the drug-loading efficiency and capacity are largely affected by the compatibility of the polymers and drugs, which is often very poor and limited. Therefore, using a similar PEGylated hyperbranched poly(amido amine), poly(BAC2AMPD1)-PEG, a surfactant-free emulsion-based preparation technique to obtain redoxresponsive nanogels was designed [64]. The advantages of this surfactant-free approach are that no purification processes are needed as no potential cytotoxic compounds like surfactants and initiators are used and efficient loading of neutrally charged, highmolecular-weight compound that can be released upon exposure to high redox potential environment. As seen in Fig. 2.12, a W/O emulsion can be generated feasibly from the mixture of water and solution of hydrogel precursor, poly(BAC2-AMPD1)-PEG in chloroform. The water droplets dispersed in chloroform are stabilized and filled with poly(BAC2-AMPD1)-PEG, and the nanogels are formed via intermolecular disulfide exchange reaction of poly(BAC2-AMPD1)-PEG. Water-soluble species can be loaded into the nanogels by simply dissolving it in water before emulsification. In Fig. 2.13(a) and (b), a W/O emulsion was generated from a mixture of water and poly(BAC2-AMPD1)-PEG in chloroform solution spontaneously without any energy input. Then by adding more water and removing chloroform, a stable aqueous solution of nanogels can be yielded. TEM image (Fig. 2.13(c)) of these nanogels showed near-spherical solid particles with diameter of approximately 90 nm. This result also indicated that unlike typical surfactants in emulsion, poly(BAC2-AMPD1)-PEG not only reduces the interfacial tension along the fluid boundary of the water droplets and chloroform but also fills up the water droplets. When poly(BAC2-AMPD1)-PEG are brought sufficiently close to each other in the water droplets, intermolecular disulfide exchange reaction occurred to form the cross-linked polymer gel network. As seen from Fig. 2.13(d), the addition of fluorescein isothiocyanate (FITC)-dextran before emulsification did not significantly affect the formation, and the loaded nanogels

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Figure 2.12  Schematic illustration of preparation of fluorescein isothiocyanate (FITC)-dextran loaded poly(BAC2-AMPD1)-PEG nanogels [64]. AMPD, 4-(aminomethyl)piperidine; BAC, N,N-cystaminebis(acrylamide); GSH, glutathione; PEG, poly(ethylene glycol); W/O, water in oil. Reproduced by permission of John Wiley and Sons.

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Figure 2.13  Photographs of the mixture of 1 part of water and 20 parts of a solution of poly(BAC2-AMPD1)-PEG in chloroform at (a) 0 h and (b) 20 h. Transmission electron microscopy images of (c) nanogels formed via spontaneous emulsification; (d) fluorescein isothiocyanate–dextran loaded nanogels obtained via spontaneous emulsification; nanogels obtained via homogenization at 13 500 rpm (e) without heat treatment and (f) after heat treatment at 50°C for 3 h [64]. AMPD, 4-(aminomethyl)piperidine; BAC, N,N-cystaminebis(acrylamide); PEG, poly(ethylene glycol). Reproduced by permission of John Wiley and Sons.

formed are not observed to be noticeably different from nonloaded nanogels. Agitation such as shaking, ultrasonication, and homogenization can also facilitate the generation of the W/O emulsion and produce structurally similar nanogels (see Fig. 2.13(e)). The rate of homogenization has an insignificant effect on the average size of the water droplets in chloroform and hydrodynamic diameter (Dh) of the nanogels dispersed in water, presented in Fig. 2.14(a) and (b). The rising average scattering intensity collected during dynamic light scattering measurement in Fig. 2.14(c) indicated that a higher rate of homogenization can result in the formation of more emulsified water droplets. Furthermore, 45% more poly(BAC2-AMPD1)-PEG was recovered from the polymer-filled water droplets in emulsion generated at 24,000 rpm than at 9500 rpm. This observation also signifies that high loading efficiency of water-soluble drugs can be facilitated by generating the emulsion with greater homogenization rate. During emulsification, poly(BAC2-AMPD1)-PEG are brought close together in the water droplets to allow some intermolecular disulfide exchange reaction to occur. To facilitate denser cross-linking via the exchange reaction, the nanogel solution was heated at 50°C for various durations. From the TEM image in Fig. 2.13(f) it is visually clear that heating did not significantly alter the structure of the nanogels formed.

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Figure 2.14  Dependence of (a) Dh of poly(BAC2-AMPD1)-PEG nanogels, (b) average size of water droplet in chloroform without heat treatment, and (c) average scattering intensity of emulsion generated on the rate of homogenization with water/polymer ratio of 2 [64]. AMPD, 4-(aminomethyl)piperidine; BAC, N,N-cystaminebis(acrylamide); PEG, poly(ethylene glycol). Reproduced by permission of John Wiley and Sons.

Furthermore, as shown in Fig. 2.15(a), during heating, the swelling of the nanogels decreased for the first 6 h and stabilized thereafter approximately at 2.5 times. The heat-treated nanogels were found to be more stable against salt out, which is likely to be attributed to the denser cross-linking. Akin to typical surfactant, the water droplets and nanogels showed a decrease in size with decreasing water to poly(BAC2-AMPD1)-PEG ratio (Fig. 2.15(b)). With a lower ratio, more poly(BAC2-AMPD1)-PEG is available to compensate the higher interfacial tension associated with smaller water droplets. However, this trend stabilizes when the ratio is smaller than 2.5. The smallest dimension attainable is probably limited by the size and shape of poly (BAC2-AMPD1)-PEG. The degradation of poly (BAC2-AMPD1)-PEG by a reducing agent, GSH, was monitored by dynamic light scattering (Fig. 2.16(a)). Upon exposure to GSH, for the first 45 min, the nanogels experienced a significant drop in size, whereas the average scattering intensity remained relatively constant. Thereafter, the nanogels started to swell and the average scattering intensity started to fall rapidly. These observations indicate that during the initial exposure, the nanogels underwent further cross-linking, which was facilitated by a GSH-induced intermolecular disulfide exchange reaction, resulting in the decrease in size and constant scattering intensity. However, once the maximum cross-linking density was achieved, the nanogels started to degrade, which is supported by the rapid increase in size and fall in scattering intensity. The nanogels consist of two areas of high and low cross-linking density. In the absence of GSH, only FITC-dextran encapsulated in the low cross-linking density area is released, resulting in the release profile shown in Fig. 2.16(b,i). On the contrary, in

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Figure 2.15  (a) Dependence of (i) Dh of poly(BAC2-AMPD1)-PEG nanogels dispersed in aqueous solution and (ii) average diameter of the dried nanogels measured from transmission electron microscopy images on the heat treatment time at 50°C. (b) Dependence of (i) Dh of poly(BAC2-AMPD1)-PEG nanogels dispersed in aqueous solution and (ii) average size of water droplets in chloroform before heat treatment on the ratio of water (mL)/poly(BAC2-AMPD1)PEG (mg) generated at 24,000 rpm [64]. AMPD, 4-(aminomethyl)piperidine; BAC, N, N-cystaminebis(acrylamide); PEG, poly(ethylene glycol). Reproduced by permission of John Wiley and Sons.

the presence of GSH, as explained in the earlier paragraph, GSH-induced intermolecular disulfide exchange reaction expanded the area of high cross-linking density and tightly trapped more FITC-dextran. As a result, lesser FITC-dextran from the shrunk area of low cross-linking density is released rapidly in the first 10 h (Fig. 2.16(b,ii)). When the degradation of the nanogels started to dominate, more FITC-dextran is gradually released over the next 6 weeks.

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Figure 2.16  (a) Dependence of (i) average scattering intensity and (ii) Dh of poly(BAC2AMPD1)-PEG nanogels dispersed in aqueous solution on the incubation time with 10 mM GSH at 37°C. (b) Release profiles of fluorescein isothiocyanate (FITC)-dextran loaded from poly(BAC2-AMPD1)-PEG nanogels in 10× PBS (i) without and (ii) with 10 mM GSH [64]. AMPD, 4-(aminomethyl)piperidine; BAC, N,N-cystaminebis(acrylamide); PEG, poly(ethylene glycol). Reproduced by permission of John Wiley and Sons.

This study shows that the surfactant-free emulsion-based approach to redoxresponsive nanogels is a promising and feasible technique to produce biodegradable carriers for delivery of neutrally charged drugs, such as carbohydrate-based drugs. Furthermore, more work can be done to expand this approach to other classes of polymers for other biomedical applications.

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2.4.3  Redox-responsive nanogels/microgels from thermoresponsive poly(amido amine)s As most of the approaches reported to prepare nanogels and microgels involve surfactants, monomers, initiators, or cross-linkers, which are potentially cytotoxic, extensive purification steps are always required to remove these compounds. Therefore, simple methods to prepare nanogels and microgels are highly desired. One such approach uses the thermosensitive hyperbranched poly(amido amine)s, S-HPAA, described in Section 2.2, to yield nanogels and microgels [24]. Because S-HPAA has a lower critical solution temperature of about 35°C in water, when the S-HPAA solution in water– methanol (v/v = 9:1) is heated to 50°C, the dehydration of the N,N-dimethylamine units leads to the formation of nanoparticles. At this point, the nanoparticles are stabilized by weak intermolecular van der Waals interactions; therefore if the solution is quickly cooled to 25°C, the nanoparticles can disassemble and the polymers can dissolve back in the solution. However, keeping the heated solution at 50°C for another 20 min does not yield the same observation even after cooling the solution to 0°C. The intermolecular disulfide exchange induced during the prolonged heating cross-linked the nanoparticles; thus the nanoparticles remained stable even at low temperature. The transmittance of S-HPAA solution decreases with increasing temperature, but no change in transmittance is observed upon cooling the solution back to even −20°C, indicating this process is irreversible. The size of the nanogels/microgels can be easily controlled by varying the polymer concentration, with the size increasing with the polymer concentration. There is no formation of nanogels when the polymer concentration falls below 0.1 mg/ mL, whereas bulk hydrogel is formed when the concentration is above 30.0 mg/mL. Moreover, the nanogels/microgels can undergo redox-induced degradation when subjected to 10 mM DTT.

2.4.4  Other redox-responsive nanogels/microgels Since nanogels were reported for drug delivery in the late 1990s [53,65], various preparation methods have been developed, including via polymerization of monomers in homogenous or heterogeneous environment, cross-linking polymer precursors, and using template-assisted nanofabrication [47,66]. Typically, to form redox-responsive nanogels through polymerization, particle templates where the polymerization will occur, must be generated in the continuous solvent phase. In the case of emulsion polymerization [67,68], emulsifiers are used to form and stabilize particle templates, which are the water droplets in the continuous oil phase. In these water droplets, the polymers will be synthesized and cross-linked to form nanogels. Because the nanogels are formed within the emulsifier-stabilized water droplets, their sizes are highly dependent on the dimension of the water droplets. Furthermore, using emulsion as templates ensures that the nanogels obtained are highly uniform in size. On the other hand, precipitation polymerization tends to yield less uniform and regular redoxresponsive nanogels because of the lack of stabilizers [69,70]. In precipitation polymerization, the monomers, initiators, and cross-linkers are soluble in the initial

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continuous phase; however, upon initiation, the polymers formed are insoluble and start to precipitate. After precipitation, the polymerization continues by absorbing the monomers, initiators, and cross-linkers into the particle templates [71]. In the case of both emulsion and precipitation polymerization, the active species are usually loaded after the redox-responsive nanogel formation [67–69]. Cross-linking the polymer precursor is also another approach to yield redox-­ responsive nanogels. Generally, non-cross-linked nanoparticles are first formed via self-assembly, like nanoprecipitation or electrostatic complexion, followed by intermolecular cross-linking of the polymers. Polymers that have intrinsic disulfide bonds can be cross-linked either by adding a catalytic amount of reducing agents [56] or heating [72] to facilitate the thiol-disulfide exchange reaction to form redox-responsive nanogels. As for other polymers, cross-linkers such as cystamine or BAC can be added to form the three-dimensional polymer networks [73,74]. Besides depending on self-assembly of polymers, in some interesting works, researchers employ layerby-layer assembly on sacrificial silica microspheres to form hydrogel microcapsules [75]. Alternative layers of thiol-functionalized poly(methacrylic acid) (PMASH) and poly(vinylpyrrolidone) (PVPON) are deposited on the sacrificial templates under predetermined conditions to promote hydrogen bonding between the layers. Then, under oxidative condition, the PMASH are cross-linked and PVPON are removed. Lastly, the silica templates are also removed to form the hydrogel capsules. Detailed cellular study on the hydrogel capsules showed that the capsules can be internalized by human colon cancer cells and DOX-loaded capsules are 5000 times more cytotoxic than free DOX [76]. Because the shape of the micro/nanogels depends on the sacrificial templates, cubic hydrogels are also successfully fabricated [74]. Using Mn2O3 cubic particles, PMA and PVPON are deposited layer by layer inside the porous template. After cross-linking the polymer layers with cystamine and dissolving the template, nanoporous and biodegradable PMA cubic hydrogel particles are formed. So far, the methods discussed to prepare nanogels are all via the “bottom-up” approach. However, a novel “top-down” photolithographic approach known as Particle Replication in Nonwetting Templates (PRINT) was reported in 2005 [77,78]. By using nonwetting molds, the formation of residual interconnecting film between molded objects can be eliminated and the formation of isolated, monodispersed nanogels allowed. This technique allows strict control over the particle size, shape, composition, and surface functionality. Using PRINT, in one of their studies, they managed to fabricate redox-responsive cylindrical small interfering RNA–conjugated hydrogel nanoparticles, which under reductive environment can release their genetic materials for gene silencing in vivo [79,80].

2.5  Conclusions Development biomaterials based on stimuli-responsive polymers so as to achieve the requirements of different applications is highly promising. As we gain deeper understanding of the human physiological redox environment, redox-responsive polymers become more relevant and attractive to the biomedical scientific community. As a

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result, numerous approaches to prepare redox-responsive hydrogels have been formulated to produce suitable biomaterials for different applications, including drug delivery and tissue engineering. In our study, well-defined redox-responsive hydrogels are produced via the intermolecular disulfide exchange reaction of hyperbranched poly(amido amine)s, prepared feasibly via Michael addition polymerization of trifunctional amines with disulfide-containing diacylamide. Owing to the unique features of the intermolecular disulfide exchange reaction, the cross-linking degree of the redox-responsive poly(amido amine) hydrogels can be well tuned by adjusting the pH value of the system. Furthermore, making use of this intermolecular disulfide exchange reaction, redox-responsive micro/nanogels can be prepared with/without emulsion as template. A surfactant-free approach that uses thermosensitive poly(amido amine)s that undergo heat-induced self-assembly and self-cross-linking can also produce nanogels. In a holistic view, this chapter has shown that redox-responsive poly(amido amine) hydrogels can be explored for many biomedical applications, including controlled and targeted drug delivery and tissue engineering, because of the feasible preparation processes, good biocompatibility, well-defined structure, and redox-response performances.

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Stimuli-responsive guar gum composites for colon-specific drug delivery

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Sougata Jana1, Sabyasachi Maiti1, Subrata Jana2 1Gupta College of Technological Sciences, Asansol, India; 2Indira Gandhi National Tribal University, Amarkantak, India

3.1   Introduction In the last three decades, synthetic and natural polymers are widely used in the biomedical and pharmaceutical fields [1–6]. Natural polymers are advantageous over synthetic polymers in pharmaceutical and biomedical research because of their inexpensiveness, biodegradability, biocompatibility, nontoxicity, and ease of surface modification [7–10]. Among natural polymers, the polysaccharides, a class of monosaccharide (sugars) polymers, are available in versatile structures and properties. Natural polysaccharides have a large number of reactive functional groups, which are amenable to easy chemical and biochemical functionalization to form novel materials. These materials are mainly used for the design and development of drug delivery systems for improving bioavailability of the drugs [11–14]. Among the polysaccharides, guar gum (GG) has been extensively used in pharmaceutical fields as a promising therapeutic carrier for colon delivery [15,16]. Guar gum (GG) is isolated from the seeds of Cyamopsis tetragonolobus (Leguminosae family) and chemically consists of linear chains of (1, 4)-β-d-mannopyranosyl units with α-d-galactopyranosyl units attached via (1, 6) linkages (Fig. 3.1) [17]. It is a low-cost, biodegradable, biocompatible, and nontoxic polymer [18]. GG produces a highly viscous solution in aqueous medium at lower concentration and finds applications in chemical industries such as food, textile, paper, petroleum, and pharmaceuticals [19,20]. Nevertheless, native GG has some limitations, such as high swelling, uncontrolled rates of hydration, thickening effect, high chances of microbial attack, instability during storage, and almost uncontrollable viscosity due to its fast biodegradation [21]. GG has found application as a binding and disintegrating agent in hydrophilic matrix systems for prolonged release of drug via the oral route [22–24]. The gelling property and enzymatic degradation of GG has enabled the fabrication of potential drug delivery carrier to colon [25,26]. The various functional groups, such as dCONH2, dNH2, dCOOH, dOH, and dSO3H, of natural polysaccharides have been chemically modified for designing pH-responsive, sustained drug release systems [27–29]. Composites may be defined as substances that consist of two or more physically and chemically different phases that are separated by a distinct interface [30]. The different polymeric systems are combined to achieve a composite system with desired mechanical Biopolymer-Based Composites. http://dx.doi.org/10.1016/B978-0-08-101914-6.00003-X Copyright © 2017 Elsevier Ltd. All rights reserved.

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Biopolymer-Based Composites CH2OH O H OH H H OH

H

O

OH CH2

H

H H O

O

OH

OH H

H

H

O

H

H

OH

OH H

H H O CH2OH

Figure 3.1  Chemical structure of guar gum.

strength. Scientists are trying to fabricate a new class of green biodegradable composites by combining with biodegradable polymers. The biobased composites with their constituents have been developed from natural resources for potential biological applications [31,32]. Currently, the biopolymer-based colon drug delivery systems are not restricted to local treatment and are also used for systemic therapeutic effect. Colonic drug delivery of drugs may be essential when prolonged drug absorption is required from therapeutic point of view, e.g., arthritis angina [33,34]. Various systems are being used for drug targeting to the colon, such as time-dependent systems [35], pH-sensitive polymeric systems [36], and enzyme-dependent systems [37]. In this chapter, an overview of stimuli-responsive GG-based composites is discussed with special reference to their application in colon-specific drug delivery.

3.2  Drug delivery applications of guar gum composites The stimuli-generated response is a basic property of living biological systems. Various types of systems that respond to external stimuli such as pH [38], temperature [39], light [40], chemicals, and ionic strength [41] have been developed.

3.2.1  pH-responsive composites The most important and challenging job of research scientists is to develop pH-­sensitive colon-specific drug delivery systems (CSDDSs), where the drug release is protected from the acidic environment (stomach pH). Therefore, the formulation of successful pH-sensitive CSDDSs could deliver the drug to the target site for desired therapeutic effect. Seeli et al. (2016) designed barium ion (Ba2+) cross-linked ibuprofen (IBU)loaded GGS-NaALG (guar gum succinate–sodium alginate) microbeads. Infrared analysis suggested the formation of strong cross-links between GGS and NaALG by Ba2+ ions and hydrogen bonding between the polymer chains. X-ray diffraction showed that the drug crystallinity was decreased after encapsulation into the microbeads and the

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drug was molecularly distributed throughout the polymer blend. Scanning electron microscopy (SEM) revealed that IBU-loaded GGS-NaALG beads were 1.4 mm in diameter and spherical in nature. The release of IBU from GGS-NaALG beads was only 20% in 3 h at pH 1.2 and 90% of the initial drug content within 2 h in pH 7.4 (after changing buffer). This result suggested that the formulated beads could prevent premature release of IBU in the stomach pH and specifically deliver IBU to the colon. The drug release kinetics of GGS-NaALG beads shifted from Fickian diffusion (pH 1.2) to non-Fickian diffusion mechanism (pH 7.4). The swelling of GGS-NaALG beads was pH dependent; a higher degree of swelling was observed in alkaline pH than that in acidic pH because of the existence of anionic groups in the polymer chains. The IBUloaded GGS-NaALG beads did not show any cytotoxic effect on cultured mouse mesenchymal stem cells [42]. The design and development of pH-sensitive hydrogel systems are important for oral delivery of protein drugs. George et al. (2007) fabricated alginate–guar gum (ALG-GG) hydrogel for delivery of bovine serum albumin (BSA), a model protein. The pH-sensitive hydrogel beads are prepared by the glutaraldehyde (GA) cross-linking method. The release of BSA from ALG-GG was ∼20% at pH 1.2 and was significantly higher, ∼90%, at pH 7.4 [43]. Das et al. (2015) fabricated dexamethasone (DX)-loaded GG–poly (acrylic acid)–β-cyclodextrin (GG–PAA–CD) hydrogel systems for intestinal drug delivery. The hydrogel was prepared by cross­ linking of tetraethyl orthosilicate. The formulated hydrogels showed pH-responsive swelling and different drug release rates at pH 2.0 and pH 7.0. MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] assay revealed noncytotoxicity and biocompatibility of DX-GG–PAA–CD hydrogel in fibroblast cell (L-929 cells) [44]. Bosio et al. (2014) developed 25-μm ALG–carboxymethyl guar gum (ALG–CMGG) microspheres containing Congo red (CR). The encapsulation of CR was 89.7% at an ALG:CMGG ratio of 7:1 in the presence of 50 mM calcium chloride. ALG–CMGG microspheres did not liberate their load at pH 1.2 (within 25 min); however, they released 68.2% CR up to 8 h [45]. Dziadkowiec et al. (2017) synthesized cationic guar gum–montmorillonite nanocomposites containing IBU for controlled drug delivery. IBU is a nonsteroidal anti-inflammatory drug (NSAID) with antipyretic and analgesic activity. In vitro IBU release from the nanocomposites was performed at pH 7.4 using the dialysis method and a sustained release pattern was shown [46]. Sinha et al. (2004) fabricated 5-fluorouracil (5-FU)-loaded GG-based compression coated tablet for colorectal cancer. The coating of xanthan gum:GG (10:20) resulted in 67% and 80% 5-FU release in the presence of 2% and 4% cecal content, respectively, in 19 h [47]. Furthermore, in 2005, Sinha et al. formulated a technetium-labeled GG-based tablet for colon drug delivery. In vitro release of technetium from the tablet was evaluated at pH 1.2 for 2 h and at pH 6.8 up to 25 h. Gamma (γ) scintigraphy experiments were performed in six healthy human volunteers. It was found that the tablet remained intact during its transit through the upper gastrointestinal tract (GIT) [48]. Celkan et al. (2016) prepared theophylline-containing GG-based compression coated tablets for the treatment of nocturnal asthma. In vivo scintigraphy studies confirmed that the prepared tablet was distributed throughout the GIT in healthy volunteers [49]. Krishnaiah et al. (2001) developed mebendazole-loaded GG-based matrix tablets by wet granulation method for colon delivery. Thermal ­analysis showed no interaction between mebendazole and GG. In vitro release was performed in 0.1 N HCl (pH 1.2) and pH 7.4

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Sorensen’s phosphate buffer. Approximately, 50% of mebendazole was released in buffer solution (pH 7.4) containing rat cecal. The accelerated stability study (45°C/75% relative humidity for 3 months) did not reveal any changes in the physical appearance of the tablet [50]. In another experiment, they evaluated the pharmacokinetics of mebendazole-loaded matrix tablet in healthy volunteers (Krishnaiah et al., 2003). In vitro drug release study was performed at pH 1.2 for 2 h and thereafter continued at pH 6.8 up to 19 h in the presence of cecal content. The in vivo study design was based on crossover design and compared between immediate release (IR) and formulated matrix tablet. The peak plasma concentration (Cmax) was 25.72 ng/mL at Tmax 9.46 h, whereas IR tablets showed Cmax 37.26 ng/mL at Tmax 3.46 h. The area under plasma concentration time curve from time zero to infinite (AUC0-∞) was 528 ng/mL/h for IR tablets and 554 ng/ mL/h for colon-targeted matrix tablets. The result confirmed that the formulated matrix tablet prolonged the release of mebendazole in the colon without releasing any significant amount of the drug at stomach pH and small intestine pH [51]. Al-Saidan et al. (2004) fabricated GG-based three-layer matrix tablets containing metoprolol tartrate. The hardness of the matrix tablets was 5.63 kg that was able to control the release of metoprolol tartrate. The different pharmacokinetic parameters were (1) Cmax(ng/mL) 165.5 for IR tablets and 61.3 for matrix tablets, (2) Tmax (h) 2 for IR tablets and 4.7 for matrix tablets, and (3) AUC0-∞ (ng/mL/h) 1444.1 for IR tablets and 2035 for matrix tablets. The elimination half-life (t1/2) of metoprolol tartrate was 4.2 for IR tablets and 19.4 for matrix tablets. In vivo pharmacokinetic parameters confirmed that the matrix tablet released metoprolol tartrate in a controlled manner for a prolonged period [52]. Furthermore, Al-Saidan et al. (2005) formulated rofecoxib-loaded GG-based matrix tablets by wet granulation method for the treatment of colorectal cancer. Formulated matrix tablets released only 5%–12% rofecoxib in the simulated gastric fluid (SGF) and simulated intestinal fluid (SIF). They evaluated matrix tablets containing 70% GG in vivo and established the colon-targeting capacity of this system. In vivo experiment in human volunteers revealed decreased plasma Cmax and prolonged absorption time, which indicated the delivery of matrix tablets to the colon [53]. Singh et al. (2012) investigated nitrofurantoin-loaded GG-based matrix tablets for colon targeting. In vitro release was evaluated at pH 1.2 for 2 h and phosphate buffer pH 7.4 and the data showed that only 10% nitrofurantoin was released in 5 h [54]. Potu et al. (2011) formulated fenoprofen calcium (FC)-loaded GG-based compression coated tablet. FC is a propionic acid derivative NSAID, used to reduce mild to moderate pain, fever, and inflammatory diseases, such as ankylosing spondylitis, rheumatoid arthritis, osteoarthritis, and juvenile arthritis. In vivo X-ray investigation demonstrated the drug release in colon [55]. Generally, 5-amino salicylic acid (5-ASA) is absorbed in the upper part of the GIT and is not available in the colon. Sawarkar et al. (2015) incorporated sulfasalazine into GG-based tablets for the treatment of inflammatory bowel disease. Sulfasalazine is a prodrug that liberates the active form 5-ASA after metabolic conversion. In vivo antiinflammatory activity was investigated in chemically (2,4,6-trinitrobenzenesulfonic acid and acetic acid) induced colitis rat model, and a significant reduction of the inflammation was noticed in rat colon [56]. GG-based tablets containing vancomycin HCl have also been reported by Elkhodairy et al. (2014) for colon targeting. In vitro microbiological assay in Staphylococcus aureus isolates showed promising inhibitory action [57].

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Vats et al. (2012) prepared piroxicam-loaded GG microspheres and subsequently compressed them into matrix tablets and provided eudragit-S100 coating. Most of the loaded piroxicam (97%) was released from the optimized tablet in simulated colonic fluid (SCF) [58]. Hashem et al. (2011) formulated GG-based tablet containing prednisolone. The tablets were then coated with hydroxypropyl methylcellulose for colon delivery. They evaluated the mechanical properties, in vitro prednisolone release, and in vivo behaviors in human volunteers. In pH 7.4 PBS containing rat cecal content (2% w/v), the tablets showed the highest release rate. In vivo γ-scintigraphic study in human volunteers was conducted using technetium-99 m-diethylenetriamine pentaacetic acid as a tracer and confirmed that the formulated tablets remained intact in the stomach and small intestine and completely released the tracer in the colon [59]. Curcumin is a natural substance used in colon cancer treatment, but it is poorly absorbed from the upper GIT. Elias et al. (2010) fabricated curcumin-loaded GG-based matrix tablets by wet granulation techniques for enhancing curcumin bioavailability. The higher GG content in the tablets led to a reduction of the drug release rate. The curcumin release from 40% GG-containing formulation reached 91% in 24 h in the presence of rat cecal contents, but the same was 82% for the tablets containing 50% GG [60]. Quercetin, a natural antioxidant, exhibits poor absorption in the upper GIT and is used for the treatment of colon cancer. Singhal et al. (2011) formulated a GG-based matrix tablet for enhancing oral bioavailability of quercetin [61]. Lai et al. (2010) prepared lansoprazole-loaded GG-based matrix tablets for colon-targeted delivery of a poorly aqueous-soluble drug lansoprazole. In vitro release of lansoprazole in SIF (pH 6.8) after 5 h was  alginate > gellan gum > xanthan gum > carrageenan. Besides the slow diffusion of the drug through the gel matrix, there might also be ionic interactions between the anionic polymer chains of gellan gum, xanthan gum, and carrageenan and the positively charged PH, slowing down the release of the drug even further. There might have even been repulsion between the positively charged amino groups of the chitosan backbone and the drug, facilitating the diffusion of the drug through the chitosan gel matrix as seen by the comparably high value for the chitosan release rate constant. The relatively high release rate for the alginate formulation may be explained by the fact that it is unaffected by the addition of monovalent cations, which might also result in less interaction with PH, exhibiting only one positive charge. This is the opposite for the carrageenan system, whose phase behavior is more affected by the addition of monovalent cations, which may cause stronger ionic interactions between the negatively charged polysaccharide chains and the positively charged drug, explaining the slow release rate of pilocarpine from this system. Unlike the biphasic clearance pattern of the solution, the polymer formulations based on gellan gum, xanthan gum, carrageenan, and chitosan seemed to drain at a slower and almost constant rate from the rabbit eye [36]. Because the gellan gum, xanthan gum, carrageenan, and chitosan formulations were equally viscous (15 ± 1 mPa s) before instillation, the prolonged retention could be attributed to the increased viscosity of the formulations due to gelling once in contact with the cations present in the tear fluid and the mucoadhesive effect of the polymers. In general, both anionic and cationic charged polymers demonstrated better mucoadhesive properties than nonionic polymers. The alginate formulation exhibited poor performance. If the cation concentration is too high, alginate forms gel upon addition of divalent Ca2+ ions and may result in precipitation of the polymer. Therefore the rapid tear turnover, which may have further enhanced the viscosity of the formulations based on gellan gum, xanthan gum, and carrageenan, may have resulted in precipitation of alginate in front of the eye and exhibited a similar drainage profile to that of the solution. The miotic response profiles for pilocarpine hydrochloride released from the various formulations over a period of 3 h revealed that the AUC0–3h was the same for all anionic polysaccharide formulations. However, the AUC values and the miotic response at 2 h were increased by 2.5-fold after administration of gellan gum, xanthan gum, and carrageenan formulations compared with an aqueous solution. The miotic response was in the order of gellan gum > xanthan gum > carrageenan, suggesting that

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the gellan gum formulation formed the strongest gel and was therefore retained in the precorneal area the longest. The drug is initially available at the same time from all formulations. However, the chitosan formulation exhibited a rather low tmax value (peak time to reach maximum drug concentration) of 25.0 ± 0.91 min compared with the other formulations, indicating an earlier onset of pharmacodynamic response. This could be attributed to the repulsion between the positively charged drug and the cationic polymer backbone, facilitating the initial drug release from the chitosan formulation, while the ionic interactions of drug with the anionic polymer systems would lead to a slightly higher tmax value because of increased precorneal retention of drug resulting from the charge–charge interactions. All the tested in situ gelling systems were practically nonirritant, as there were no signs of vascular response, such as hyperemia, hemorrhage, and coagulation. Ion-activated in situ gelling systems of gellan gum and HPMC were also investigated for treating fungal keratitis [37]. The ocular in situ gels could enhance the drug residence time for better ocular absorption of fluconazole. The formulation was therapeutically efficacious and stable and provided sustained release of drug over a period of 8 h. A combination of Pluronic F-127 and Pluronic F-68 with chitosan (pH-sensitive polymer as a permeation enhancer) was used in temperature- and pH-triggered in situ gelling systems, whereas gellan alone was used in ion-activated in situ gelling systems for ofloxacin [38]. In situ gel-forming ability of the developed systems significantly controlled precorneal drainage as studied by gamma scintigraphy and were nonirritant. The Cmax of the in situ gelling formulation was found to be 1.5 times higher than that of the marketed eye drops at the similar tmax of 1 h. Another in situ ophthalmic drug delivery system of antibacterial agent, linezolid, was based on ion-activated guar gum derivatives [39]. Novel polymers such as cationic guar with hydroxypropyl guar were used as gelling as well as viscosity-enhancing agents. The in situ gel sustained the release of the drug up to 12 h. In vivo ocular toxicity studies revealed nonirritant and nontoxic nature of the formulation. Therefore the guar gum derivative-based ophthalmic in situ gel by virtue of its prolonged corneal residence time and sustained drug release could help in achieving higher ocular bioavailability. The composite of 1% gellan gum and 0.5% kappa-carrageenan (4:1) showed the best mechanical and mucoadhesive properties [40]. The composite polymer solution did not exhibit either cytotoxicity or irritability. There were no alterations or signs of toxicity after being administered every 12 h for 3 months in rats. During this period, no changes in corneal thickness were detected and tissue damage was observed by direct visual observation or immunohistochemical techniques. Ocular delivery system of levofloxacin, based on the concept of ion- (sodium alginate) and pH- (chitosan) activated in situ gelation concept was reported by Gupta et al. [41]. Owing to their elastic properties, in situ gels resisted the ocular drainage of drug, leading to longer contact times with the ocular surface. Both the drug release studies (over 12 h) and in vivo precorneal retention studies by gamma scintigraphy indicated better therapeutic efficacy of this system than the standard eye drops. Novel pH-responsive in situ gel of alginate and HPMC for ophthalmic delivery of an antibiotic ciprofloxacin hydrochloride was also reported for clinical interventions of corneal ulcers [42].

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They utilized the instantaneous calcium alginate gel-forming ability of mucoadhesive alginate polymer by virtue of its interaction with divalent cation (Ca2+) present in lachrymal fluid. HPMC was further incorporated to achieve the desired consistency for sustained drug release. Only 67.02% drug was released in phosphate buffer (pH 7.4) in 7 h. The addition of HPMC exhibited more pseudoplasticity as compared with the formulation prepared without HPMC.

6.2.3  Nanoparticles The association of an active molecule with a nanocarrier allows the molecule to interact intimately with specific ocular structures and thus overcome ocular barriers and prolong its residence in the target tissue [43]. Furthermore, this technology offers promising solution for formulating various poorly soluble drugs in the form of eye drops [44]. Because the cornea and conjunctiva have a negative charge, it was thought that the use of mucoadhesive polymers would increase the drug residence time after interaction with these extraocular structures [45]. Polymeric nanoparticles (NPs) could be the best drug delivery tools for treating ocular diseases. The self-assembly of amphiphilic CS into NPs has attracted increasing interest in pharmaceutical areas because of the simplicity relative to those prepared by covalent cross-linking, ionic cross-linking, and desolvation method [46]. Moreover, the hydrophobic microenvironments formed by the association of hydrophobic components enable the NPs to reserve hydrophobic drugs. Yuan et al. [47] synthesized amphiphilic cholesterol-modified chitosan (CS-CH) as carrier for hydrophobic cyclosporine A, generally used for the treatment of dry eye disease. The synthesis of CS-CH involved mixing of cholesterol-hemisuccinate in N,N-dimethyl formamide and 1% chitosan– hydrochloric acid solution followed by the dropwise addition of 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (Fig. 6.3). The diameters of the NPs were 50–200 nm. The monomodal size distribution of nanospheres was noted at a degree of cholesterol substitution of 1.7. Bimodal size distribution was observed at higher degrees of cholesterol substitution. A low critical aggregation concentration (0.01–0.04 mg/mL) value of the copolymer indicated stability of NPs at dilute solution. The loading efficiency was poor and was about 42%. Only 60% of the initial drug was released from NPs in 4 h and then continued up to 48 h, releasing 95% of the entrapped drug in saline solution. 99mTc-labeled CS NPs showed good precorneal retention after topical administration; however, no radioactivity was detected in the posterior segment. Motwani and coworkers [48] designed mucoadhesive CS-sodium alginate polyelectrolyte complex NPs to prolong ophthalmic delivery of an antibiotic, gatifloxacin. Briefly, an aqueous alginate solution was sprayed into the chitosan solution containing Pluronic F-127 under continuous magnetic stirring. Pluronic F-127 (0.50% w/v) was added to aid in drug solubilization. The drug encapsulation efficiency and particle size varied from 61% to 82% and 205 to 572 nm, respectively. In STF, pH 7.4, the NPs released only 10%–12% during the first hour, followed by a gradual drug release over 24-h periods. The poor aqueous solubility of econazole nitrate, an antifungal drug, creates a barrier for the treatment of ophthalmic fungal infection [49]. The situation demands a

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HCl CH3CH2

N C N

(CH2)3 N

CH3

CH3

HOOCCH2CH2COO

O OCCH2CH2COO HCl CH3CH2

NH

(CH2)3 N

C N

CH3 O

CH3

OCCH2CH2COO NH2

O

O

OH O CH2OH

NH

NHCOCH3

n

O O

OH O CH2OH

m

NHCOCH3

O

OH O CH2OH

n

O OH O

m

CH2OH

Figure 6.3  The scheme of conjugation of chitosan and cholesterol-hemisuccinate. Reprinted from Yuan X, Li H, Yuan Y. Preparation of cholesterol-modified chitosan self-aggregated nanoparticles for delivery of drugs to ocular surface. Carbohydrate Polymers 2006;65:337–45, copyright (2006) with permission from Elsevier.

suitable system that could deliver the drug in an effective concentration to the eye. Mahmoud et al. [50] combined the mucoadhesive property of cationic CS nanocarriers and drug-holding ability of the anionic sulfobutylether (SBE)-β-cyclodextrin (CD) for sustained delivery of the drug to the eye. The hydrophobic interaction with the interior cavity of β-CD along with electrostatic interaction between the cationic drug and anionic SBE-β-CD were presumed to be responsible for drug-SBE-β-CD complex formation [51]. The diameter of nanosystems ranged from 90.8 to 461.2 nm depending on the concentration of CS and SBE-β-CD. The lowest concentration of CS and SBE-β-CD corresponded to the lowest size and drug content of the NPs (23.97%). The maximum value, i.e., 45.67%, corresponded to the highest CS and SBE-β-CD concentrations. The NPs prepared using 360 kDa CS showed lesser drug content and loading capacity than those prepared using 150 kDa CS. Transmission electron microscopy image indicated that CS/SBE-β-CD NPs were predominantly spherical in shape with an irregular surface. About 50% of the drug content was released in phosphate buffer solution (pH 7.4) in 8 h. The characteristic melting endothermic peak of drug disappeared in drug-loaded NPs, indicating the amorphous dispersion of econazole in CS-NPs. The sharp crystal X-ray diffraction peaks of drug were overlapped with the noise of the coated polymer and disappeared, indicating its complete and successful encapsulation into the core of the CS-NPs. Upon mixing of the CS-NPs with mucin, the zeta potential value reduced, and this confirmed the mucoadhesive properties of the NPs. Perhaps the ionic interaction between negatively charged sialic acid residues in mucin and positively charged amino groups in CS-NPs caused a significant reduction of the surface charge of the NPs.

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Sterile filter paper discs were placed under the eyelid of rabbit for 1 min at 1-h intervals up to 8 h following a single instillation of 50 μL NPs dispersion equivalent to 0.2% econazole in the conjunctival sac of rabbits. The discs were then placed in the inoculated Sabourand dextrose broth tubes. Then the inoculated broth was incubated at 27 ± 0.5°C for 24 h. The area under the percentage inhibition of yeast growth versus time curve for CS:SBE-β-CD (0.15:2.5) NPs was 1.34 times higher than that for CS:SBE-β-CD NPs (0.2:3). The antifungal effect of LC3-2 NPs increased gradually up to 4 h and then decreased gradually at low ratios of CS:SBE-β-CD. In contrast, the NPs showed almost constant antifungal effect at high ratios of CS and β-CD complex. Despite similar drug concentration, the tested NPs differed in their mucoadhesive properties, and hence, the sustained effect of the NPs might be stronger at higher ratios of CS:SBE-β-CD than at lower ratios. Natamycin, a polyene antifungal drug is considered as the drug of choice for treating filamentous fungal keratitis. Currently, the drug is available as 5% (w/v) ophthalmic suspension instilled in the conjunctival sac at hourly or two-hourly intervals for at least 4–6 weeks to get complete relief [52]. To reduce the frequency of application and improve effectiveness, Bhatta and coworkers [53] prepared sustained release mucoadhesive lecithin/CS NPs for ocular instillation. The NPs were prepared by injecting methanolic lecithin/drug solution into acetic acid solutions of chitosan under magnetic stirring at 1000 rpm. An increase of the lecithin/chitosan ratio from 5:1 to 20:1 caused a significant improvement of drug entrapment efficiency from 60.06% to 78.29%. The surface charge of NPs reduced as a result of the ionic interaction between mucin and NPs after incubation with mucin for 6 h. This suggested mucoadhesive property of the NPs. A biphasic release pattern was evident in phosphate buffer solution (pH 7.4). The burst release of 10%–15% was attributed to desorption and diffusion of the drug from the NP surface, followed by slow release of ∼64.22% drug in 7 h. However, almost all the drug was released within 2 h from the drug suspension. The MIC90 (minimum inhibitory concentration of drug at which 90% of the isolates are inhibited) of NPs was similar to that of free drug against Candida albicans (3.12 μg/mL) and Aspergillus fumigatus (1.56 μg/mL). No signs of irritation or damaging effects in the cornea, conjunctiva, or iris of male rabbits were reported for the NPs. Compared with the marketed suspension, the NPs exhibited ∼1.47-fold higher bioavailability and ∼7.4-fold lower clearance. The MRT of the NPs was about five times higher than that of the pure drug. Thus the positively charged NPs can provide a binding force to the eye surface. It was stated that the bioadhesion might also be promoted by the formation of hydrogen bonds between free NH2 and OH groups of CS molecules and mucin components, such as sialic acid residues of the ocular surface [54]. A long-term treatment schedule was suggested for mitigating corneal and conjunctival intraepithelial neoplasia and squamous cell carcinoma [55]. The acute and chronic side effects of 1% 5-fluorouracil (5-FU) are more frequent because of high concentration and systemic absorption. This encouraged the development of 5-fluorouracil-loaded NPs to achieve high bioavailability even at lower doses and reduce the dosing frequency to minimize systemic absorption [56]. Nagarwal et al. [57] proposed chitosan-coated alginate–chitosan NPs for the ocular delivery of this drug. Chitosan coating increased the size of the NPs (chitosan:alginate = 1:2) from 412 to 505 nm.

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A higher encapsulation efficiency of 26.66% was obtained. No interaction was found between alginate–chitosan NPs and mucin. Upon CS coating, the NPs exhibited a significant increase in the viscosity of mucin, suggesting a considerable interaction and mucoadhesion characteristics of coated NPs. Chitosan coating failed to suppress the burst release as compared with uncoated alginate NPs. The drug-loaded alginate-chitosan NPs (SA–CH DNPs) and chitosan-coated alginate-chitosan NPs (CH-SA–CH DNPs) released 74.15% and 81.20% of entrapped 5-FU in 8 h, respectively, in phosphate buffer solution (pH 7.4). However, the drug permeation rate through excised rabbit eye from coated NPs was higher than alginate NPs and drug solution. The ocular bioavailability of coated and uncoated NPs was 5.81- and 3.74-folds higher than that of the aqueous solution. More precisely, the coated particles exhibited 1.55 times higher bioavailability than the uncoated particles. The maximum ocular drug concentration (Cmax) was also found to be higher for the coated particles. The NPs showed low concentration of 5-FU in vitreous humor because of high availability of drug in aqueous humor (Fig. 6.4). The plasma drug levels were negligible thus minimizing systemic side effects of the drug. No damage to cornea, iris, and conjunctiva of New Zealand albino rabbits was reported up to 12 h, thus proving good tolerability of the formulation. Tropicamide-loaded Ca2+-carboxymethyl tamarind gum NPs were reported [58]. The particle size of 339 nm and encapsulation efficiency of 15.57% was found with

5-FU concentration (µg/mL)

30

5-FU

SA-CH DNPs

CH-SA-CH DNPs

25 20 15 10 5 0 Vitreous humor

Aqueous humor 2h

Aqueous humor

Vitreous humor 4h

Aqueous humor

Vitreous humor 8h

Time (h)

Figure 6.4  Comparison of 5-fluorouracil (5-FU) level in aqueous humor and vitreous humor of rabbit eye at different time intervals. CH-SA-CH DNP, chitosan-coated alginate-chitosan nanoparticle; SA-CH DNP, drug-loaded alginate-chitosan nanoparticle. Reprinted from Nagarwal RC, Kumar R, Pandit JK. Chitosan coated sodium alginate–chitosan nanoparticles loaded with 5-FU for ocular delivery: in vitro characterization and in vivo study in rabbit eye. European Journal of Pharmaceutical Sciences 2012;47:678–85, copyright (2012) with permission from Elsevier.

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the NPs of ovoid morphology. Drug permeation from the NPs across isolated goat cornea was comparable with that of aqueous drug solution. The NP suspension adsorbed 87.67% mucin, hinting their excellent bioadhesion potential. Perhaps the carboxylic groups formed hydrogen bonds with oligosaccharide chains of mucin and exhibited bioadhesive potential. The nonirritating potential coupled with bioadhesiveness supported that the nanoparticulate formulation could provide sustained delivery of drug as a consequence of greater precorneal residence. One of the most effective antiviral agents against herpes simplex virus is acyclovir. However, limited transmucosal absorption and ocular retention time of its conventional formulation constitute a big challenge for effective antiviral therapy. Microparticles (1–10 μm) were prepared by spray-drying technique. Briefly, the drug was dissolved in tripolyphosphate (TPP) aqueous solution (pH 4.5). This solution was then added to the chitosan hydrochloride solution under magnetic stirring, and the resulting dispersion was spray dried. CS-NPs (30–300 nm) were prepared by ionic gelation of TPP and chitosan hydrochloride. High encapsulation efficiency (75%) was observed for the microparticles, whereas the same was low, only 16%, for the NPs. The microparticles liberated their content in PBS solution (pH 7.4) at a faster rate than NPs (75%–80% in 24 h). Calderón and associates [59] examined the irritation potential of the particulate system based on slug (Arion lusitanicus) mucosal irritation assay. This slug produces mucus when exposed to irritating substances. It was noticed that slugs treated with drug-loaded chitosan particles produced an increased amount of mucus compared with the negative control (waxy maize starch). The amount of mucus produced by the slugs was a measure of irritation. Based on this observation, the drug-loaded chitosan particles were considered as a moderately irritating substance. The ionotropic gelation technique is mostly used for the preparation of drug-embedded CS NPs. In this method, the NH2 moiety of CS is protonized to NH3 + in acidic medium, which interacts with an anion moiety, PO4 −2 group of TPP, by electrostatic interaction and forms NPs [60]. Ameeduzzafar et al. [61] synthesized carteolol-loaded NPs of 168.90 nm size at an optimized CS and TPP concentration of 0.15% and 0.17%, respectively. The NPs had a mean drug entrapment efficiency of 69.57%. The particles showed an initial burst release of 28% in 1 h followed by 88% release in 24 h in STF, pH 7.4. In contrast, about 84% drug was dissolved in 1 h and almost completely released in 4 h from aqueous suspension of pure drug. Hence, the CS-NPs exhibited pronounced sustained release. The transcorneal permeation and flux of the drug through goat cornea was significantly increased (73.20% and 125.1 μg/cm2/h) compared with free drug (34.48% and 58.20 μg/cm2/h) in 6 h. Owing to hydrogen bond formation between the positively charged amino group of chitosan and oligosaccharide chains of mucin [62], the degree of NP adsorption on pig mucin glycoprotein was found to be 90.65%, thus demonstrating a strong mucoadhesive nature of CS-NPs. Rhodamine B–loaded CS-NP suspension penetrated the cornea up to a length of 94.04 ± 3.26 μm as demonstrated by a high fluorescence intensity, whereas the drug solution penetrated 36.92 ± 4.53 μm during 8 h of confocal laser scanning microscopy (CLSM). Histopathological examination revealed that CS-NPs did not perturb the cellular structure of the cornea. Hen’s egg test (HET)-chorioallantoic membrane (CAM) test is a semiqualitative method for assessing potential irritancy of compounds

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by observing adverse changes that occur in the CAM of the egg after exposure to test chemicals [63]. Because the CAM of an embryonated hen’s egg is similar to the vascularized mucosal tissues of the human eye, the HET-CAM test can provide information on the effects that may occur in the conjunctiva following exposure to a test substance in vivo. The HET-CAM test showed that the formulation was well tolerated and nonirritant. Gamma-scintigraphy revealed that the CS-NP suspension was retained in the precorneal area for a period longer than aqueous CRT solution because there was a quick fall in 99mTc radioactive counts on the corneal surface. The drug solution was cleared rapidly from the corneal region and reached into the systemic circulation by the nasolachrymal drainage system after 6 h of ocular administration. As a consequence, the drug-loaded NPs significantly reduced the intraocular pressure than the aqueous drug solution at an equivalent dose and extended the hypotensive effect up to 24 h following instillation into rabbits. Katiyar and groups [64] prepared and optimized the dorzolamide-loaded CS-NPs for effective glaucoma therapy. At 0.175% concentration of CS and TPP, the particles were 164 nm in size and spherical and had a drug entrapment efficiency of 98.10%. The NPs were then dispersed in 1.25% sodium alginate solution with a reasonable viscosity. The formulation released about 74.90% drug in STF (pH 7.4). The dispersion of NPs in 1.25% alginate solution released 58% drug in 8 h, thus making this system superior to only NPs. Only 35.80% drug permeated through goat cornea because of slow release of drug from the NPs and gelling polymer matrix in 2 h. The corneal mucoadhesion of nanocomposite hydrogels increased by 1.5-fold compared with only NPs perhaps due to the cumulative effect of chitosan and alginate. The HET-CAM test revealed that in situ gelling NPs were practically nonirritant and therefore safe. Gamma scintigraphy showed that the 99mTc-labeled formulation was retained very well in the precorneal region during the first 25 min of instillation on albino rabbits. The formulation was retained at the corneal surface even after 2 h. Probably the mucoadhesive CS and alginate prolonged the residence time on the corneal surface and consequently improved the bioavailability of dorzolamide. To better modulate the morphology and physical stability of particles, Andrei et al. [65] introduced gelatin in the CS NPs to lessen the concentration of toxic covalent cross-linker used to stabilize TPP-induced gelled particles. The cross-linking mechanism is depicted in Fig. 6.5. Glutaraldehyde cross-linking of dual polymer particles offered cefuroxim encapsulation efficiency of around 47%. After an initial burst effect for 1 h, the composite systems slowly released the drug over 48 h in alkaline environment (pH 7.4). The ocular biodistribution of NPs was assessed after intravitreal administration into Wistar rats. The fluorescein-labeled NPs reached the lens and retina in large quantities 24 h after administration. The particles were still present at the retina, lens, sclera, and cornea, although at a low level, at 72 h, as was evident from fluorescence intensity. The biodistribution profiles of the particles clearly demonstrated their potential to migrate to the cornea and retina after a longer time. Amphotericin B, a natural polyene antifungal antibiotic, is used to treat fungal keratitis, which if neglected, may lead to permanent vision loss. The drug dissolved in lecithin solution was dispersed into aqueous phase containing Pluronic F-127 under

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stirring at 58°C for evaporation of organic solvent. Chhonker et al. [66] reported a decreasing tendency of particle size and drug entrapment efficiency of NPs with decreasing molecular weight (MW) of CS. The size ranged from 161.9 to 282.7 nm and the drug entrapment efficiency persisted between 70.0% and 74.6%. Low viscosity grade CS demonstrated rapid desorption and diffusion of drug molecules at the initial hour. This would be beneficial in terms of antifungal activity as it helps achieve the therapeutic concentration of drug in minimal time period. Low-MW CS NPs released 88% drug at 10 h. The same was 83% and 86% for high- and medium-MW CS NPs, respectively. The zone of inhibition against C. albicans and A. fumigatus was maximal for medium-MW CS NPs. The low-MW CS formulation exhibited significantly low clearance and high AUC as compared with high-MW CS NPs. Probably an increase in MW resulted in increased particle size, which eventually increased ocular clearance due to excessive tear flow and eye blinking. An alternative way of administering poorly soluble natural flavanone naringenin was investigated by Zhang et al. [67] for the treatment of age-related macular degeneration [68]. An inclusion complex of naringenin and SBE-β-CD was prepared to

Chitosan Gelatin

Figure 6.5  Double cross-linking mechanism. TPP, tripolyphosphate. Reprinted from Andrei G, Peptu CA, Popa M, Desbrieres J, Peptu C, Gardikiotis F, Costuleanu M, Costin D, Dupin JC, Uhart A, Tamba BI. Formulation and evaluation of cefuroxim loaded submicron particles for ophthalmic delivery. International Journal of Pharmaceutics 2015;493:16–29, copyright (2015) with permission from Elsevier.

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enhance its aqueous solubility. Ionically gelled spherical CS-SBE-β-CD NPs of 446.4 nm size prolonged the release of drug in STF for 5 h. The formulation was nonirritant to the rabbit’s eye on long-term study. The concentration of the drug in the aqueous humor of rabbit was significantly higher than that from the suspension. The ocular bioavailability of the drug after application of NPs was about 4.6-fold higher than that of the suspension. As mentioned earlier, the NPs might have interacted with the cornea and conjunctiva because of charge attraction and could therefore provide an improved ocular pharmacokinetics. Dexamethasone sodium phosphate (DEX) is used to treat inflammations caused by injury, surgery, or other infectious conditions in the eyes. Despite serious systemic side effects, it has been advised that the controlled delivery of DEX at the targeted site via nanocarriers could suppress the ocular inflammatory conditions, circumventing the side effects [69]. Kalam [70] designed mucoadhesive CS NPs through ionotropic gelation technique for improving its precorneal retention and corneal permeability. The mucoadhesive and free-flowing properties were imparted to the NPs by coating with HA. The NPs had a solid dense spherical structure with smooth surfaces. The uncoated NPs with TPP:chitosan (2:3) were little cloudy (Fig. 6.6(a)), but HA coating improved the appearance of NPs and caused a slight increase in the size of the CS-NPs (Fig. 6.6(b)). The mean diameter of the CS-NPs increased from 305.25 to 400.57 nm after coating. The physical stability and mucoadhesive potential of the CS-NPs was predicted from high negative zeta potential (31–33 mV) values. The electrostatic repulsion between similarly charged NPs prevented agglomeration of the particles. A high negative surface charge prolonged precorneal retention of the CS-NPs because of possible interaction of HA-coated particles with the hyaluronan receptors present on the corneal and conjunctival epithelial cells [71]. Rapid and efficient adsorption of negatively (a)

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Figure 6.6  Scanning electron micrographs of dexamethasone sodium phosphate–chitosan nanoparticles (DEX-CS-NPs): (a) without hyaluronic acid (HA) coating and (b) with HA coating. Reprinted from Kalam MA. Development of chitosan nanoparticles coated with hyaluronic acid for topical ocular delivery of dexamethasone. International Journal of Biological Macromolecules 2016;89:127–36, copyright (2016) with permission from Elsevier.

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charged HA onto the surface of CS-NPs resulted in a relatively higher negative surface charge density on HA-coated CS-NPs, which also attributed to a significant increase in the hydrodynamic diameters of the coated CS-NPs. Furthermore, fast dissolution of DEX adsorbed onto the huge specific surface area of uncoated CS-NPs caused a rapid and initial burst release of DEX (40%–50% in 2 h) from NPs in STF medium (pH 7.4). The surface coating prevented the fast dissolution of DEX into the release medium to a limited extent (35% in 2 h). The release rate became slower and sustained thereafter over a minimum period of 12 h. The viscosity of the formulations lied in the range of 31.25–38.23 mPa s for easy dispersion and sterilization by filtration. Relatively higher viscous formulation than normal solutions (20 mPa s) could prolong the corneal retention without any blurred vision and discomfort due to any particulate sensitivity. Furthermore, they evaluated transcorneal permeation, in vivo ocular irritation potentials, and bioavailability following application of uncoated and coated CS-NPs in rabbit eyes [72]. The amount of drug that permeated through the excised rabbit cornea was 52.86 and 55.01 μg/cm2 at 6 h, respectively, for the uncoated and coated NPs. The developed CS-NPs were nonirritant to the rabbit eyes. The topical administration of uncoated and coated NPs into rabbit’s eye enabled quantification of the drug in aqueous humor up to 24 h. Obviously, 1.83 and 2.14 times higher bioavailability was observed with the uncoated and HA-coated NPs as compared with DEX solution. The Cmax of DEX from CS-NPs and HA-CS-NPs were found to be 1.075 and 1.082 times higher than that of DEX solution, respectively (Fig. 6.7). The data suggested that HA-coated CS-NPs performed better than uncoated CS-NPs possibly because of the higher mucoadhesive nature of HA as well as its interaction with hyaluronan receptors on corneal epithelia. Rosmarinic acid, an antioxidant, could

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Figure 6.7  Dexamethasone (DEX) concentrations in aqueous humor after topical administration of different DEX-containing formulations into rabbit eyes (mean ± SD, n = 3). CS, chitosan; HA, hyaluronic acid; NP, nanoparticles. Reprinted from Kalam MA. The potential application of hyaluronic acid coated chitosan nanoparticles in ocular delivery of dexamethasone. International Journal of Biological Macromolecules 2016;89:559–68, copyright (2016) with permission from Elsevier.

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be a potent inhibitor of retinal neovascularization and may be applied in the treatment of vasoproliferative retinopathies [73]. Some workers encapsulated rosmarinic acid in chitosan and sodium tripolyphosphate (7:1) NPs [74]. The NPs appeared to be safe without relevant cytotoxicity against retina pigment epithelium (ARPE-19) and human cornea cell line (HCE-T). The NPs were 280 nm in size and had an encapsulation efficiency of 60.2%. The monitoring of zeta potential values revealed that the surface charge decreased probably because of the electrostatic interaction between positively charged chitosan and anionic mucin [75]. The formation of microaggregates after mucin incubation also suggested an interaction occurs between NPs and mucin. This mucoadhesive nature caused an increased retention of CS-NPs over the ocular mucosa after instillation and hence was found promising for ocular application in oxidative eye conditions.

6.2.4  Composite films/inserts Ocular inserts are sterile, thin, multilayered, drug-impregnated solid or semisolid devices placed into the cul-de-sac or conjunctival sac for therapeutic application [76]. The thin polymeric film should also be nontoxic, biocompatible, and biodegradable. Ophthalmic films are currently being developed for overcoming the ocular barriers and preventing loss of drugs through the lacrimal drainage system. Controlling compositions of polymers of different grades has facilitated the modification of key characteristics of thin films such as drug release rate, mucoadhesive properties, mechanical strength, and other related properties [77]. Ophthalmic thin films are generally applied to treat diseases of the anterior segment such as conjunctivitis, glaucoma, and chronic dry eye syndromes. Mucoadhesive films may facilitate in extending residence time at the application site, leading to prolonged therapeutic effects. Most of the thin films having mucoadhesive properties are hydrophilic in nature, undergo swelling, and form a chain interaction with the mucin [78]. Among several polymers, mucoadhesive properties are exhibited by chitosan, hyaluronan, cellulose derivatives, polyacrylates, alginate, and gelatin. Compared with nonionics, the cationic and anionic polymers facilitate strong interaction with mucus. A cross-linked gelatin insert was used by Attia et al. [79] to enhance bioavailability of dexamethasone in the rabbit eye. The level of dexamethasone in the aqueous humor was found to be fourfold greater than dexamethasone suspension. Saettone et al. [80] prepared ophthalmic pilocarpine nitrate inserts using poly(vinyl alcohol) (binder):glyceryl behenate (lubricant) at a ratio of 1.6:1 and one of the biopolymers such as xanthan gum, jota-carrageenan, and HA. The inserts were subsequently coated with a mixture of Eudragit RL and RS. The coated inserts released half of drug content in 3–5 h in phosphate buffer (pH 7.4); however, the uncoated inserts did the same within half an hour. Coating of xanthan gum cores provided faster drug release than the other two polymers (∼80% drug release in 9 h). The medicated inserts were applied into the lower conjunctival sac of nonanesthetized male New Zealand albino rabbits, and the best overall miotic activity was demonstrated by HA matrix. Aburahma and Mahmoud [81] used the film-casting method for making polymeric ocular inserts of PVP K-90 and biopolymer containing an ocular antihypertensive drug brimonidine. Alginate-based ocular inserts were coated with hydrophobic ethyl

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cellulose polymer. The highest swelling capacity was observed for chitosan-based uncoated ocular inserts. Low-MW and medium-MW alginate retained their integrity after swelling, although the swelling index was still lower than that of chitosan-based ocular inserts. Low-MW chitosan and alginate inserts showed higher swelling ability than that exhibited by higher-MW biopolymers. Increased cross-linking between the high-MW polymer chains limited water penetration and thus decreased the swelling of the matrix system [82]. The CS-based ocular inserts emptied the entire content in less than 1 h, whereas alginate-based inserts delayed the drug release rate and liberated almost all drug after 6 h. It has to be mentioned that the extent of drug release from the medium-MW alginate-based films was relatively low. The low viscosity grade alginate contains shorter polymer chains that form a loose network and allow more water penetration. This causes faster swelling and erosion of the polymer matrix, thereby facilitating rapid release of the drug [83]. An increase in polymer molecular weight assisted in greater entanglement of the polymer chains and reduced the polymer dissolution rate. This collectively led to slow diffusion of drug molecules and therefore prohibited drug release [84]. One-sided ethyl cellulose coat on low-MW alginate-based ocular inserts remained intact and dramatically delayed brimonidine release for more than 24 h. Differential scanning calorimetry thermograms of the medicated film indicated complete drug amorphization and good distribution of brimonidine in the film matrices. The adverse and/or damaging effects of the ocular inserts were evaluated by observing the conjunctiva and cornea of rabbits’ eyes. Uncoated ocular alginate inserts and one side–coated ocular insert applied in the lower conjunctival sac of rabbits were in general well tolerated and did not show any visible redness or inflammation in the conjunctiva and cornea of rabbit’s eyes. The ocular insert of low-MW alginate with ethyl cellulose coating exhibited better pharmacodynamic performance in terms of decreasing the intraocular pressure compared with eye drops. Ocular insert of gatifloxacin, sodium alginate, polyvinyl alcohol, and glycerin was also reported [85]. Eudragit RL-100 coated and calcium chloride (CaCl2) cross-linked ocular insert showed maximum drug permeation, i.e., 89.53% in 11 h through cellulosic membrane. Ca2+ cross-linking and RL-100 coating sustained the drug release up to 12 h and demonstrated satisfactory ocular tolerance. A recent literature indicated that bimatoprost-loaded inserts could lower the intraocular pressure for 4 weeks following a single application and thus seemed suitable for glaucoma management [86]. Another report described that xyloglucan ocular films could control the release of ciprofloxacin for 24 h, releasing 98.85% drug at the end. The safety of the formulation to ocular mucosa was also assured [87].

6.2.5  Biopolymer solid lipid nanoparticles Solid lipid nanoparticles (SLNs) are submicron-sized lipid emulsions where the liquid lipid (oil) has been substituted by a solid lipid. In some cases, solid lipids are mixed with small amounts of liquid lipids (oils) to achieve high drug load called nanostructured lipid carriers (NLCs). SLNs have gained increasing interest in the field of

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ocular drug delivery because of their ability to encapsulate and protect the lipophilic drug, enhance ocular tolerance, improve penetration efficiency, and increase corneal uptake [88]. Resina Draconis, a deep red resin has been traditionally used in ocular inflammations. However, the poor solubility of Resina Draconis limits its therapeutic efficacy and clinical application. Hao et al. [89] fabricated a thermosensitive in situ gelling formulation. The melt-emulsion ultrasonication and low temperature-solidification method was used for the fabrication of Resina Draconis–loaded SLNs. Resina Draconis was dissolved in the molten lipid phase and subsequently emulsified into a hot water-surfactant solution. The ultrasonication of warm emulsion led to the formation of nanoemulsion, which was rapidly cooled to yield uniform dispersion of NPs (150 nm). More than 90% of the drug was loaded into the lipid NPs because of the high lipid solubility of Resina Draconis. The thermosensitive composite hydrogels were designed by dispersing SLNs into poloxamer solution. The permeation rate across rabbit cornea was steady, and this phenomenon was most likely related to the mucoadhesive properties, which enhanced retention time and corneal adsorption and thus formed a drug reservoir for sustained permeation. Coumarin-6-labeled SLNs deeply penetrated into the cornea after application of poloxamer liquid gels. The corneal CLSM images observed at various times after administration of hydrogel are shown in Fig. 6.8. The green fluorescent strip mainly concentrated on the corneal epithelium at 0.5 h and then moved gradually into the inside of the cornea. The fluorescence intensity became more homogeneous over time and supported that SLNs deeply penetrated into the cornea. They reasoned that surfactant poloxamer reduced the surface tension and easily adsorbed onto the ocular surface. The natural lipophilic character of the outer epithelium favored SLN permeation. After deposition of the SLNs in the corneal epithelium, the passive transport of drug molecules was initiated for entry into the aqueous humor and posterior segment of the eye. This fact coupled with nonirritant behavior of the polymer system could be useful for effective transcorneal delivery of lipophilic drugs. Bacterial keratitis, if left untreated, often leads to progressive tissue destruction, with corneal perforation. Ofloxacin (OFX), a fluoroquinolone antibacterial agent, is highly active against both gram-positive and gram-negative bacteria. Ustündağ-Okur et al. [90] studied the effect of chitosan oligosaccharide lactate (COL) modification on lipid formulation in enhancing corneal residence time and corneal bioavailability of the drug. Ofloxacin-loaded NLCs were produced by microemulsion and high-shear homogenization method. The molten mixture of oleic acid and Compritol lipids was mixed with aqueous phase containing Tween 80 and ethanol to form a transparent oil-inwater (o/w) microemulsion. NLC dispersions were produced by dispersing the warm o/w microemulsion into cold water (2–3°C) under mechanical mixing at equal microemulsion:water ratio. In another method, the aqueous phase was homogenized with lipid phase and subsequently, the particles were dispersed in ultrapure water at 4°C. For drug loading into NLCs, the drug was added to the lipid phase. To promote corneal residence and mucoadhesion, a cationic COL polymer solution was added to the aqueous phase of NLCs. The microemulsion method produced smaller NLCs (11.35 nm) than those produced by the homogenization method (157.6 nm).

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Figure 6.8  Confocal laser scanning micrograph of rabbit corneal tissues after the administration of Coumarin 6-labeled solid lipid nanoparticle-based in situ gelling system at various times (A: corneal epithelium, B: cornealstroma, C: corneal endothelium). Reprinted from Hao J, Wang X, Bi Y, Teng Y, Wang J, Li F, Li Q, Zhang J, Guo F, Liu J. Fabrication of a composite system combining solid lipid nanoparticles and thermosensitive hydrogel for challenging ophthalmic drug delivery. Colloids and Surfaces B: Biointerfaces 2014;114:111–20, copyright (2014) with permission from Elsevier.

Biocomposites in ocular drug delivery Groups / Zone inhibition ring 0.3 % OFX solution/ 0.75 % COL solution

S. aureus ATCC 29213 (grampositive)

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S. aureus ATCC 29213 (grampositive)

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Figure 6.9  Zone inhibition diameters of formulations. COL, chitosan oligosaccharide lactate; OFX, ofloxacin. Reprinted from Ustündağ-Okur N, Gökçe EH, Bozbıyık Dİ, Eğrilmez S, Ozer O, Ertan G. Preparation and in vitro–in vivo evaluation of ofloxacin loaded ophthalmic nano structured lipid carriers modified with chitosan oligosaccharide lactate for the treatment of bacterial keratitis. European Journal of Pharmaceutical Sciences 2014;63:204–15, copyright (2014) with permission from Elsevier.

Microemulsion-based NLCs (NM) released the drug at a faster rate during the first 3 h than NLCs prepared by the homogenization method (NH) because of their smaller size and related higher surface area; however, they extended the drug release up to 12 h in PBS. NH-COLOFX exhibited the highest antibacterial activity than other formulations against Staphylococcus aureus. The antibacterial activities of drug-loaded NM-COLOFX and NH-COLOFX against Escherichia coli were almost similar. The interaction between positively charged chitosan molecules and negatively charged microbial cell membranes led to the disruption of the microbial membrane and subsequently caused cell death due to the leakage of proteinaceous and other intracellular constituents [91]. The drug-free NM-COL and NH-COL showed a significant antibacterial effect, which would be advantageous for ocular drug administration (Fig. 6.9). After 48 h of the contact period, the amount of drug that permeated through the cornea from NM NLCs was higher than that from NH NLCs. The modification of NLCs with COL decreased the amount of penetration through rabbit cornea. The increased viscosity following addition of COL was thought to lessen NLC motion in the formulation. NH-COL4OFX showed the longest residence time of 1 h on the eye surface. The efficacy of NLCs was investigated in bacterial keratitis induced in rabbit eyes. The infected groups were treated with 50-μL formulations twice daily over 1 week. The clinical presentations of eyes infected with S. aureus after treatment with formulations are shown in Fig. 6.10. It was observed that the corneal opaque area decreased faster with

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0. day

7. day

NMCOL4OFX

NH-COL4OFX

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Figure 6.10  Photographs demonstrate clinical presentation of eyes infected with S. aureus after 0 and 7 days treatment with NM-COL4OFX, NH-COL4OFX, and commercial solution. COL, chitosan oligosaccharide lactate; OFX, ofloxacin. Reprinted from Ustündağ-Okur N, Gökçe EH, Bozbıyık Dİ, Eğrilmez S, Ozer O, Ertan G. Preparation and in vitro–in vivo evaluation of ofloxacin loaded ophthalmic nano structured lipid carriers modified with chitosan oligosaccharide lactate for the treatment of bacterial keratitis. European Journal of Pharmaceutical Sciences 2014;63:204–15, copyright (2014) with permission from Elsevier.

the NLC formulations. The rabbits applied NH-COL4OFX formulation demonstrated lower conjunctival redness and corneal opacity than NM-COL4OFX-applied rabbits. As was evident from aqueous humor concentration–time profiles of OFX following ocular instillation in rabbits, the peak concentration of drug in the aqueous humor was higher with COL-modified NLC formulations in comparison with the commercial solution. The bioavailability was greater with the NH-COL4OFX formulation in comparison with NM-COL4OFX NLCs. Betaxolol hydrochloride is used to slow down the progression of glaucoma by IOP. Montmorillonite (MMT), a clay mineral, has an alumina octahedral sheet sandwiched

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by two silica tetrahedral sheets on both sides in a single layer. An acid-treated montmorillonite (acid-MMT) was intercalated with betaxolol hydrochloride (BH) in the interlayers, and this nanocomposite was further used for encapsulation by SLNs (MMT-BH-SLNs) using an emulsion evaporation–low temperature solidification method [92]. The drug entrapment efficiency was 75.2% for the 150-nm particles, which is favorable for cornea permeation. The release profile of BH-SLNs liberated 39.7% drug within the initial 2 h but led to 70.2% release within 12 h. In the case of MMT-BH-SLNs, the initial release and the cumulative release decreased to 33.1% and 58.3% at the respective times. The intercalation of BH into the interlayer space of MMT significantly suppressed BH release. The cation exchange process is responsible for sustained release of drug for up to 12 h in artificial tear fluid. In rabbit ocular irritation tests, the MMT-BH-SLNs caused lower irritation than BH solutions. MMTBH-SLNs possessed slow drug release properties and therefore could significantly increase the drug corneal permeability.

6.3  Conclusion The composite materials can surmount shortcomings of individual materials and may give rise to synergistic functions that are absent in either of the components. A critical review of a plethora of research reports indicated that the nanocomposites, in situ forming hydrogels, SLNs, and ocular films had great promise in enhancing corneal penetration of the drug. The preliminary data provided indications of good ocular tolerance and acceptability. Although nanotechnology has shown promise in ocular applications, a greater emphasis must be placed on quantitative correlation of the dose and exposure of NPs with their penetration, therapeutic efficacy, and toxicity. Despite a vast potential of composite materials as ocular drug delivery device, there still remains a lot of challenges that must be overcome before their clinical use. However, detailed pharmacological and toxicological studies are needed to demonstrate their therapeutic potentials. Therefore, in future studies, the long-term toxicity study as well as biological properties such as protein adsorption, cell adhesion, tissue compatibility, and overall in vivo effect of such composite systems should clearly be addressed.

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Keerti Jain National Institute of Pharmaceutical Education and Research (NIPER), Raebareli, India

7.1   Introduction Poor bioavailability is one of the main obstacles that reduce therapeutic effectiveness of various conventional therapeutic moieties as well as new chemical entities. Hence the development of a suitable formulation is one of the critical steps in drug design and development. In the modern era, nanomaterials are extensively being investigated to develop optimum formulation of drug molecules to improve bioavailability as well as to obtain controlled drug delivery attributed to their unique properties such as extraordinarily small size and high surface/mass ratio. A large number of nanocarriers, such as nanoparticles, nanoemulsion, carbon-based nanomaterials including carbon nanotubes (CNTs) and graphenes, dendrimers, and quantum dots, are being investigated to develop the nanoformulation with the ability to provide controlled drug delivery including brain delivery with an improved bioavailability and solubility profile as well as a high therapeutic index. The size range of these nanomaterials varies from 10 to 100 nm or sometimes more in which drug molecules are dissolved, conjugated, intercalated, or encapsulated. The choice of a nanomaterial for the development of a formulation depends on a number of factors, including the ability of the formulation to (1) selectively deliver the drug molecule to the target site or site of action or diseased part, (2) protect the drug from the environment or factors such as pH, enzymes, and metabolic degradation, which can degrade it, (3) facilitate the finely tuned release of payload at the target site in the desired concentration to achieve therapeutic concentration, (4) improve potency, (5) improve therapeutic index, (6) improve patient compliance, (7) cause higher accumulation at target site, (8) facilitate therapeutic effect at a low dose, and (9) being economic. Nanotechnology is also being explored for the development of vaccine and in immunotherapy of diseases such as cancer and autoimmune disorders [1–10]. Dendrimers are also known as dendritic polymers, which are defined as threedimensional polymeric macromolecules with a highly branched tree-like structure having unique physicochemical properties, such as monodispersity, solubility, large number of surface functional groups, and well-defined size, shape, and molecular weight. The unique properties of dendrimers lie in their architecture composed of three parts, namely, core, internal branches, and surface functional groups (Fig. 7.1). Polyamidoamine (PAMAM) dendrimers were the first dendrimers Biopolymer-Based Composites. http://dx.doi.org/10.1016/B978-0-08-101914-6.00007-7 Copyright © 2017 Elsevier Ltd. All rights reserved.

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Multifunctional core

Dendrimer generations

Internal cavity spaces

Internal branching points

Figure 7.1  Diagrammatic representation of dendrimeric architecture.

designed by scientists in 1980–90. Since then a large number of different types of dendrimers and variations of dendritic polymers have been designed and reported by many scientists [11–15].

7.1.1  History The history of dendrimers starts with the discovery of hyperbranched molecules in the 1980s by two different groups of scientists, Vogtle and coworkers and Tomalia and coworkers. They synthesized these dendritic hyperbranched polymers as lowmolecular-weight dendrimers. Before the actual synthesis of these hyperbranched and dendritic polymers, the theoretical concept was proposed by scientist P. Flory in the 1950s. First, the divergent synthesis method was used in 1980–90 to synthesize a complete dendrimeric architecture. Later on at the end of the same decade, dendrimers were also synthesized by the convergent method of synthesis. Apart from the term dendrimers, some other terms, including “arborols,” which means trees, cascade molecules, etc., have also been used by scientists. But dendrimers is the most common term for these monodisperse macromolecular architectures, which is frequently used by scientists [11,16,17].

7.1.2  Dendrimers versus polymers In comparison with polymers, dendrimers have some superior advantages in various biomedical as well as industrial applications owing to the nearly monodispersed architecture. Furthermore, dendrimers also offer a dense population of surface functional groups, enabling simultaneous conjugation or labeling with different bioactives, diagnostic agents, targeting ligands, solubilizing agents, etc. Furthermore, the presence of this large number of surface functional groups or attached moieties

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Figure 7.2  Schematic representation of four classes of polymeric architecture, including linear, cross-linked, branched, and dendritic architecture. Reprinted from Kesharwani P, Jain K, Jain NK. Dendrimer as nanocarrier for drug delivery. Progress in Polymer Science 2014;39(2):268–307. Copyright (2014) with permission from Elsevier.

facilitates biological interactions leading to increased penetration across the biological membrane [14,18–21]. Broadly these hyperbranched and dendritic polymers could be divided in four classes, including random hyperbranched polymers, dendrigrafts, dendrons, and dendrimers (Fig. 7.2). Size, shape, surface chemistry, architecture, elemental composition, and flexibility/rigidity are the six parameters that have been designated as critical nanoscale design parameters, and it has been emphasized by scientists that dendrimers with highly controlled architecture should have at least these six properties [11,16,17,22,23].

7.1.3   Synthesis Different synthesis methods have been investigated for the synthesis of dendrimers and used to design different types of dendrimers. Dendrimers have been synthesized mainly by two methods, namely, divergent and convergent methods. In the divergent method, dendrimers are synthesized from a multifunctional core point in contrast to

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the convergent method in which synthesis starts from dendrons toward the core. In the convergent method, dendrons are attached to a multifunctional core to design a dendrimer. Other methods such as double exponential growth method, Lego chemistry, click chemistry, and hypercores and branched monomers growth approaches have also been explored for designing different types of dendrimers. A summary of the different methods of synthesis of dendrimers is given in Table 7.1 [9,16,17,24,25].

7.1.4  Properties Dendrimers have some characteristic properties owing to their nanometric size and well-defined monodisperse architecture. The most important unique properties of dendrimers that are particularly relevant to their drug delivery applications are listed below: •  Monodispersity: Dendrimers represent a special class of polymers synthesized via step-bystep controlled chemical reactions that give them a uniform structure with defined molecular weight resulting in a nearly monodisperse drug delivery system [9,17,25–27]. •  Polyvalency: Presence of multiple functional groups renders a polyvalent surface to dendrimers. This polyvalency of dendrimers enables modification of surface groups of dendrimers by a variety of ligands and makes possible formation of drug–dendrimer, antibody–dendrimer, targeting agent–dendrimer, imaging agents–dendrimer, etc. conjugates or complexes [25,28,29]. •  Self-assembly: Dendrimers have shown the properties getting self-assembled into vesicles and micelles. Scientists have prepared amphiphilic Janus dendrimers with hydrophilic and hydrophobic groups that have the propensity to self-assemble into vesicles, dendrimersomes, disks, helical ribbons, tubular vesicles, and cubosomes. This self-assembly of nanoarchitectures could be exploited in various bioinspired applications, including control of drug release efficiently by external triggers [30–32]. •  Electrostatic interactions: Another characteristic feature of dendrimeric nanoarchitectures is the presence of charges on the surface, which enables self-assembly into aggregates of different forms by combination of electrostatic and π-π interactions, interactions with the drug molecules and biological lipid membranes, and complexation with nucleic acids to form dendriplexes [24,30,33–37]. •  Chemical and physical stability: The unique architecture of dendrimers gives them excellent stability against chemical and physical stimuli, as well as makes possible formation of stable complexes of drug with dendrimers [38,39]. •  Unimolecular micelles: The presence of hydrophilic surface groups and hydrophobic interior cavities gives the properties of unimolecular micelles to dendrimers. This property of dendrimers makes them an important nanocarrier candidate to increase the solubility and oral bioavailability of drugs that are poorly soluble and that are not absorbed on oral administration. This property of dendrimers also assists in its cellular accumulation, making dendrimers a prospective candidate vehicle of drugs to reverse the problem of resistance [11,40–42]. •  Solubility: Again the presence of a hydrophobic interior and hydrophilic exterior makes dendrimers a promising solubilizing agent for poorly soluble drugs. The solubility efficiency of dendrimers has been found to be influenced by the pH conditions and concentration of dendrimers. By improving the solubility of poorly soluble drugs, dendrimers can significantly improve the drug delivery efficiency as a drug delivery system [43–46].

Table 7.1  A

summary of different synthetic approaches for dendrimers [9,16,24] Miscellaneous

1. In this method reaction starts with core and consists of two steps coupling and activation. a. First, two or more moles of branching monomers with protecting groups are reacted with multifunctional core b. Removal of protecting group or (hydrogenation of nitrile groups to amino groups as in PPI dendrimers) is the second step in which 1.0G of dendrimers is synthesized c. Then these two steps are repeated to get the desired generation of dendrimers 2. Dendrimers synthesized by this method are: a. PPI dendrimers b. PAMAM dendrimers c. PLL dendrimers d. Phosphorus dendrimers up to 7.0G e. Linear dendritic poly (ester)-blockpoly(ether)-block-poly(ester) copolymers f. Aliphatic ester dendrimers g. Polyphenylene dendrimers h. Siloxane and carbosiloxane dendrimers 3. Drawbacks: a. Excess quantity of reagents is required b. Difficulty in synthesizing dendrimer with structural uniformity 4. Advantages a. Large-scale production of high-generation dendrimers

1. In contrast to divergent growth method convergent synthesis starts at surface toward the multifunctional core a. In this method dendrons are synthesized by linking surface monomer units b. Then these dendrons are linked to multifunctional core to synthesize dendrimers c. This method represents stepwise inward assembly of building blocks to synthesize dendrimers 2. Dendrimers synthesized by this method are: a. Poly (aryl ether) dendrimer synthesized by Frechet and coworkers b. Poly (aryl alkyne) dendrimers c. Poly (phenylene) dendrimers d. Poly (alkyl ester) dendrimers e. Poly (aryl alkene) dendrimers f. Poly (alkyl ether) dendrimers 3. Drawbacks: a. Time consuming due to stepwise synthesis b. Scale-up is difficult 4. Advantages a. Dendrimers with structural uniformity and active functional groups could be synthesized by this method

Some other approaches such as (1) hypercores and branched monomers, (2) double exponential growth method, (3) Lego chemistry, and (4) click chemistry have been designed to speed up the rate of synthesis of dendrimers • Hypercores and branched monomers In this method preassembled oligomeric moieties are linked to synthesize large quantity of dendrimer in fewer steps. Insolubility of hypercore and difficulty in separating the by-products were the main drawbacks of this method • Double exponential growth method In this method dendrimers are synthesized rapidly using combination of divergent and convergent methods using a single focal point • Click chemistry In this method dendrimers could be synthesized with comparatively high purity using a metal catalyst • Lego chemistry This method could be described as straight forward synthesis of dendrimers, which could be tailored easily using two types of monomer. The advantages of this method include requirement of fewer steps and the nontoxicity of by-products

PAMAM, polyamidoamine; PLL, poly-l-lysine; PPI, poly(propyleneimine).

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Divergent growth method

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7.2  Types of dendrimers The special architecture of dendrimers is the most unique property that enabled scientists to synthesize a large number of different types of dendrimers with different branching monomer units and multifunctional core molecules to fulfill a variety of industrial and biomedical applications [16,26,47–53] (Fig. 7.3).

7.2.1  PAMAM (Starburst) dendrimers PAMAM dendrimers are one of the most investigated dendrimers for biomedical applications. They have been investigated for drug delivery applications as well as for some intrinsic therapeutic activity, including antimicrobial activity as topical

Tecto dendrimer

Triazine dendrimer

Chiral dendrimer

Liquid crystalline dendrimer

PPI dendrimer

Mesogenic dendrimer Peptide dendrimer

Carbosilane dendrimer

Polyether dendrimer Types of dendrimers

Hybrid dendrimer

Metallodendrimer

Glycodendrimer

PAMAM dendrimer

Polyester dendrimer

PAMAMOS dendrimer

Fulleropyrrolidine dendrimer

Figure 7.3 Different types of dendrimers. PAMAM, polyamidoamine; PAMAMOS, poly(amidoamine-organosilicon); PPI, poly(propyleneimine). Reprinted from Kesharwani P, Jain K, Jain NK. Dendrimer as nanocarrier for drug delivery. Progress in Polymer Science 2014;39(2):268–307. Copyright (2014) with permission from Elsevier.

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microbicides and as immunomodulators and antiangiogenic agents after surface engineering. Some hyperbranched dendrimers have also been designed for increasing this intrinsic antimicrobial activity of PAMAM dendrimers [48,54]. PAMAM dendrimers were commercially supplied under the name Starburst. This name was arrived at from the Starburst effect shown by PAMAM dendrimers. The increase in generation of dendrimers causes the increase in branching, which leads to a highly dense and packed architecture of higher-generation dendrimers to such an extent that it hinders and resists further growth to further higher generation. This effect is known as Starburst effect [25,49]. PAMAM dendrimers have been investigated for drug delivery applications, to develop biosensors, for gene delivery, and as nanoscaffold to deliver imaging and contrast agents [44,46,55–58].

7.2.2   Poly(propyleneimine) dendrimers Poly(propyleneimine) (PPI) dendrimers were first synthesized on large scale by two scientists, Ellen M.M. deBrabander-van den Berg and E.W. Meijer in 1993. They synthesized PPI dendrimers by sequential repetition of two steps using 1,4-diaminobutane as the core; the first step was double Michael addition of acrylonitrile to primary amines, and the second step was heterogeneously catalyzed hydrogenation of the nitrile groups, which results in doubling of the number of primary amines on the surface of dendrimers [59]. Ethylene diamine has also been used as a core molecule for the synthesis of PPI dendrimers. PPI dendrimers are also known as POPAM dendrimers, which stands for poly (propylene amine). These dendrimers have surface amino groups and have been explored for a large number of biomedical applications, including delivery of drug, nucleic acids, and imaging and diagnostic agents. They have also been evaluated for some intrinsic therapeutic activities after surface engineering [25,46,52,58,60–62].

7.2.3  Amino acid dendrimers Being part of the biological system, dendrimers made of amino acids are anticipated to be biocompatible. Hence amino acids such as lysine, ornithine, and arginine have been used to engineer the surface of dendrimers as well as in designing of new dendrimers. Both amino acid-conjugated and amino acid dendrimers have been used for the delivery of drug molecules as well as dendriplexes to deliver nucleic acids. As a transfection agent, amino acid–based dendrimers have shown high efficiency in preclinical research. Poly-l-lysine (PLL) dendrimers are one of the most explored amino acid dendrimers in biomedical applications. PLL dendrimers have also shown promising antiangiogenic activity in cell culture assays as well as in animal studies and promising results in drug delivery and gene transfection researches. The amino acid–based dendrimers as well as peptide dendrimers have shown promising antimicrobial activity, including antiviral activity, in preclinical investigation. One of the products of Star Pharma, VivaGel (SPL7013), is an example of dendrimers (4.0G naphthalene disulfonic acid surface group bearing PLL dendrimer) made up of amino acid and it is currently in clinical trials for its antimicrobial activity, including antiviral activity [41,47,50,63–67].

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7.2.4  Triazine dendrimers Triazine dendrimers are mostly based on 1,3,5-triazine units and were designed to develop biodegradable dendrimers for various drug delivery applications as well as some industrial applications. The presence of hydroxyl groups at their surface makes them biocompatible, and they have been evaluated by scientists to increase the aqueous solubility of hydrophobic drug molecules, as drug delivery vehicles, and as antimicrobial agents also. They have also shown a comparatively safer profile in view of toxicity, as a single dose up to 10 mg/kg did not produce any toxic effects on kidney and liver. Triazine dendrimers have also shown antimicrobial activity after surface engineering with groups such as 1,4-diazabicyclo[2.2.2]octane [27,68–70].

7.2.5  Phosphorous dendrimers Phosphorous dendrimers and carbosilane dendrimers are examples of two emerging inorganic dendrimers. Phosphorous-containing dendrimers, such as phosphorhydrazones, have shown some promising results in drug delivery applications. Bifunctional phosphorous dendrimers could have some advantages in biomedical applications, because the presence of two different types of functional groups on the surface of dendrimers can render the opportunities to explore these dendrimers for various biological applications [26,71–74]. It has been observed that phosphorous-containing amino-bis(methylene phosphonate)-capped dendrimers showed some efficiency in the treatment of a chronic inflammatory disease, multiple sclerosis. The investigator concluded that these phosphorous-containing dendrimers could be explored as a promising drug molecule for the treatment of multiple sclerosis and other chronic inflammatory diseases because of their ability to prevent the development of autoimmune encephalomyelitis as well as to inhibit the progression of disease. These dendrimers have also been found to redirect pathogenic myelin-specific CD4(+) T-cell response toward interleukin (IL)-10 production [71]. The phosphorous-containing dendrimers are recognized by specific receptors present on the surface of human monocytes and have shown some promising antiinflammatory and immunomodulatory activity in animal and in vitro testing [72]. For biomedical applications, interaction of dendrimers with biological components play vital role. The 3.0 and 4.0G phosphorous dendrimers with surface cationic charge have been observed to interact and form complexes with human immunodeficiency virus (HIV)-derived peptides, including aspartate transaminase, alkaline phosphatase, and l-lactic dehydrogenase, leading to significant change in the secondary structure of these proteins [74].

7.2.6  Carbosilane dendrimers Carbosilane dendrimers represent one of the inorganic dendrimers that are currently being explored particularly for its antimicrobial activity, including antiviral activity. Carbosilane dendrimers with sulfate or naphthyl-sulfonated surface groups has shown blockade of HIV-1 as entry inhibitor and by inhibiting cell-to-cell transmission of HIV-1 in a mice model. These dendrimers have also shown binding affinity for surface protein of herpes simplex virus (HSV)-2 and showed synergistic anti-HSV-2 activity

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with tenofovir and acyclovir. These polyanionic carbosilane dendrimers have also shown efficiency in preventing hepatitis C virus infection in cell culture and anticancer activity on incorporation of metals. These dendrimers have also shown ability as mucoadhesive to improve ocular administration of drug [51,63,75–79].

7.2.7   Peptide dendrimers Different peptide dendrimers have been designed by scientists for various biomedical applications, including conjugation of peptide molecules to the surface of dendrimers for targeted delivery of genes, drugs, and diagnostic agents and development of biosensors [80–83]. Peptide dendrimers have also been examined for their interactions with biological membrane. In a study with model lipid membrane, peptide dendrimers were observed to interact with lipid membranes via electrostatic interactions (between positive charges of dendrimers and negative charge of the membrane) and hydrophobic interactions (between hydrophobic branches of dendrimers and acyl chains of lipid membrane) without affecting the integrity of the lipid membrane [84]. These interactions with biological membrane led to the investigation on the antimicrobial property of peptide dendrimers with positive results [53,85].

7.2.8   Polyglycerol dendrimers Polyglycerol dendrimers have been developed by scientists as water-soluble dendrimers for various biomedical applications, such as delivery of anticancer drug, delivery of siRNA, antiinflammatory agent to inhibit progression of knee joint cartilage degradation, and biosensor [86–90]. Polyglycerol dendrimers with a large number of hydroxyl end groups can solubilize the drug molecule easily [91]. Schneider et al. [88] investigated dendritic polyglycerol sulfate to modulate the degradation of knee joint cartilage as a strategy to lessen the progression of degenerative joint disease, osteoarthritis. They investigated the efficacy of polyglycerol dendritic molecule in a rat model of osteoarthritis in which it was observed that these dendritic molecules had chondroprotective effect on osteoarthritic knee joints. In addition, the dendritic molecule was found to be biocompatible without significant cytotoxicity [88]. Stefani and coworkers have designed polyglycerol-based core-multishell carriers and investigated them as drug delivery vehicles. The developed core–shell hyperbranched nanocarrier based on polyglycerol dendritic polymer showed capacity for loading of drug molecules, including finasteride and dexamethasone, without toxicity at low concentration ranging up to 0.05 mg/mL [91]. Report is also available on the immunosuppressive activity of dendritic polyglycerol sulfate molecules, and polyglycerol dendritic molecules were found to show significant in vivo (mouse model of severe acute graft-versus-host disease) immunosuppressive effect against severe immune reaction graft-versus-host disease after hematopoietic stem cell transplantation [90].

7.2.9  Gallic acid–based dendrimers Gallic acid–triethylene glycol (GATG) dendrimers with surface amino groups have been synthesized by scientists. These dendrimers have been conjugated with

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polyethylene glycol (PEG) as well. Their efficiency to form dendriplexes with nucleic acids has been explored to develop biocompatible transfection agent to deliver complex genetic material. The amino groups of GATG dendrimers with surface cationic charge can interact with bacteria and these dendrimers have also shown antibacterial and antiviral activity. These dendrimers could be explored in future as biocompatible transfection agent, as prospective drug and gene delivery vehicle, as antiviral and antibacterial agent, for treatment of neurodegenerative disorders, and for diagnosis and study of multivalent carbohydrate recognition and dendrimer dynamics [92,93].

7.2.10  Domino dendrimers Domino or self-immolative dendrimers represent smart polymeric systems that disassemble in response to a stimulus. The presence of stimuli causes the self-immolative dendrimer to disassemble into its building blocks with special properties, including fluorogenic molecules. These dendrimers have been explored primarily as stimuliresponsive drug delivery system, for detection of biological and chemical activities, for development of functional molecular polymers with specific reactivity and spectroscopic properties, and for design of biosensor with ability for ultrasensitive detection of pathologically important antigens [94,95].

7.2.11  Miscellaneous dendrimers The most unique property of dendrimeric nanocarriers is their architecture, and to exploit the advantages of this unique architecture scientists have designed many different types of dendrimers by using different multifunctional cores and branching units. The examples of such dendrimers include polyglycerol dendrimers, mannose dendrimers, tricarboxylato dendrimers, citric acid dendrimers, cyclodextrin dendrimers, and poly(oxypropylene) triamine (Jeffamine)–based dendrimers [31,32,96–101]. Tricarboxylato and Jeffamine dendrimers have shown some promising antimicrobial activity comparable with some antibiotics available in clinical market, such as penicillins, tetracyclines, aminoglycosides, and antifungal antibiotics [96,97]. Mannose dendrimers have been synthesized as Janus glycodendrimers by scientists and shown affinity for lectin receptors [31,32,98]. Huang et al. [100] developed a series of hydroxyl-rich hyperbranched polyaminoglycosides of gentamicin, tobramycin, and neomycin using redox-responsive disulfide bonds. These hyperbranched polyaminoglycosides showed promising antibacterial activity estimated on microorganisms, Escherichia coli and Staphylococcus aureus, with good biocompatibility measured by performing hemolysis assay. These redox-responsive hyperbranched polyaminoglycosides with disulfide linkages also showed excellent in vitro gene transfection efficiency and promising in vivo antitumor activity [100]. Metals have been included in dendrimeric architectures as a part of the structure or conjugated at the surface or have been used to construct whole dendrimers for various biomedical as well as industrial applications [53,102–104]. Ferrocenoyl or ruthenocenoyl groups were included by Hoffknecht et al. [53] into peptide dendrimers to evaluate the effect of these metallic groups on the antibacterial activity of dendrimers.

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They observed that these metallic groups had an additive effect on the antimicrobial activity of peptide dendrimers [53]. Toomari and Namazi [101] designed pH-sensitive dendrimers using β-cyclodextrin (βCD) residue and formed inclusion complexes of methotrexate. These β-CD dendrimers were found to be biocompatible and showed negligible toxicity in the methylthiazol tetrazolium assay performed on T47D breast cancer cells. The conjugate also showed pH-sensitive drug release in in vitro drug release kinetic studies [101]. Chen et al. [99] reported the synthesis of reduction-cleavable hyperbranched polymers of doxorubicin (DOX) with disulfide bonds synthesized by copper-catalyzed azide–alkyne click chemistry reaction. These porous hyperbranched polymers represent molecularly imprinted polymers with high specific surface area enabling high encapsulation efficiency and showed a cumulative release of DOX more than 2 mg under optimized pH conditions, suggesting its promising potential in controlled and targeted drug delivery [99].

7.3  Surface engineering of dendrimers The presence of multiple end groups on the surface of dendrimers allows attachment of multiple functional moieties, including drug molecule, diagnostic agent, targeting ligand, genetic materials, antigen, and antibody. The purposes of surface engineering could be alleviation of toxicity by masking the hemolytic and cytotoxic groups on the surface of dendrimers, targeted delivery, gene delivery, intracellular delivery, increased retention time, solubility, development of biosensor, or incorporation of some therapeutic activity such as antimicrobial or antiangiogenic activity [8,27,29,60,105–113].

7.3.1   Biocompatibility Several reports are available on engineering of dendrimer surface groups with biocompatible ligands, such as mannose, lysine, folic acid, galactose, and maltose. All these ligands are part of biological systems and hence are expected to reduce the toxic effects speculated to be due to the presence of surface charges on dendrimers. Reports are also available on significant reduction of toxicity of dendrimers after conjugation with such ligands. PEGylation, glycosylation, and acetylation are examples of surface engineering strategies that have reduced the toxicity and increased the biocompatibility of dendrimers [28,29,109–112,114–116].

7.3.2   Conjugation chemistry Conjugation chemistry plays important roles in the surface engineering of dendrimers. In most of the researches, surface engineering of dendrimers was based on conjugation of surface groups of dendrimers with the ligand being used for surface engineering of dendrimers. Many ligands, including carbohydrates, amino acids, peptides, vitamins, PEG, and surfactants, have been conjugated with the functional end groups of dendrimers [28,29,66,106,113,117,118].

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7.3.3  Surface engineered dendrimers A large number of ligands for various drug delivery applications have been conjugated to the surface of dendrimers. For targeted drug delivery, dendrimers have been engineered at the surface with the ligands for which receptors are overexpressed in certain disease conditions or on some specific body tissues. For example, some transporters and receptors are overexpressed on cancer cells and the blood–brain barrier (BBB). The transporter GLUTs and receptors for sialic acid are overexpressed on cancer cells and the BBB. Patel et al. [119] performed a one platform, i.e., PPI dendrimers, to compare the brain-targeting efficiency of sialic acid, concanavalin A, and glucosamine to deliver the anticancer drug paclitaxel (PTX) selectively to brain tumor, which is a substrate for the P-glycoprotein (P-gp) efflux system present in brain, decreasing the intracellular trafficking of PTX in brain tumor cells. The in vivo studies showed that the ligand-conjugated dendrimers increased the delivery of PTX in the brain as compared with free PTX. Furthermore, among the investigated targeting ligand, sialic acid was found to be the most superior followed by glucosamine and concanavalin A in brain-targeting efficiency [119]. In one study folic acid–conjugated dendrimers showed promising efficiency in solubilizing a very potent and hydrophobic anticancer agent, 3,4-difluorobenzylidene diferuloylmethane, with targeted delivery to folate receptor–overexpressing HeLa and SKOV3 cancer cells [44]. A summary of different surface engineering strategies with their prospective applications is given in Table 7.2. More than one surface engineering strategy and conjugation of more than one ligand have also been explored by the scientists to develop versatile dendrimers for various biomedical applications. Some examples are simultaneous conjugation of IL-6 antibody and arginine-glycine-aspartic acid (RGD) peptide, sialyl Lewis X (Slex) antibodies and fluorescein isothiocyanate, and RGD peptide and alpha-tocopheryl succinate, and multivalent conjugation of two antibodies, EpCAM and Slex antibodies, simultaneously for preventing cancer metastasis, increasing targeting efficiency, and for theranostic application [8,107,108,120].

7.4  Dendrimers as scaffolds: guest–host relationship The compact and globular architecture of dendrimers with the presence of multiple surface functional groups and a number of inner cavity spaces (Fig. 7.1) facilitates the interaction with the drug molecule. Drug molecule could be loaded into dendrimers via three approaches, including conjugation, electrostatic interactions, and physical encapsulation [24,121]. Sanyakamdhorn et al. [121] reviewed the binding and loading efficacy of some anticancer drugs and their metabolites with some synthetic polymers including dendrimers. Anticancer drugs doxorubicin and tamoxifen, and metabolites of tamoxifen, including 4-hydroxytamoxifen and endoxifen, were included in this report. PEG, mPEG-anchored 3.0G PAMAM dendrimers, and 4.0G PAMAM dendrimers were included as synthetic polymers in this study. In this report it was stated that H-bonding and hydrophobic interactions were the main binding forces between the drug and synthetic polymer. Furthermore, 4-hydroxytamoxifen formed a more

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Table 7.2 

Surface engineering strategies for dendrimers with their prospective bioinspired applications Surface engineering

Dendrimers

Prospective applications

References

PEGylation

PPI, PLL, PAMAM, GATG dendrimers, etc.

• Reduced

toxicity biocompatibility • Enhanced efficacy of drugs such as anticancer drugs due to prolonged circulation • Promising drug and gene delivery vehicle • Improved pharmacokinetic profile • Theranostic applications • Development of stimuli (pH, redox)-responsive delivery system • Less cytotoxicity • Gene delivery applications with good transfection efficiency • Stimuli (temperature, pH)-sensitive delivery • Increased drug-loading efficiency • Promising drug delivery applications, including anticancer drug • Improved pharmacokinetic profile • As antagonists for toll-like receptors could be explored for treatment of infectious, inflammatory, and malignant diseases • siRNA delivery • Potential drug and gene delivery vehicle • Reduced hemolytic activity and cytotoxicity • Targeted delivery • Designing of biocompatible and biodegradable dendrimers • Increased biocompatibility, reduced toxicity of dendrimers as well as loaded drug molecule • Higher payload

[92,111,112, 155,171]

• Theranostic

[66,106,117, 147,175]

Acetylation

PAMAM, triazine, PPI dendrimers, etc.

Glycosylation [carbohydrate (mannose, maltose, galactose, etc.) conjugation]

Various types of dendrimers with different surface groups

Vitamins (folic acid, riboflavin, tocopherol)

PAMAM, PPI, PLL, etc.

• Enhanced

applications drug delivery • Improved efficacy of anticancer drug • Improved biocompatibility • Targeted

[115,116,122]

[29,31,32,114, 162,174]

Continued

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Table 7.2 

Surface engineering strategies for dendrimers with their prospective bioinspired applications—cont’d Surface engineering

Dendrimers

Prospective applications

References

Peptide [arginineglycine-aspartic acid (RGD), short oligopeptides of phenylalanine, arginine histidine, muramyl dipeptide (MDP), TAT peptide, etc.]

PAMAM, PPI dendrimers, PEGylated dendrimers, etc.

• Theranostic

[28,106,113, 118,176]

Antibody-conjugated dendrimers [antiHER-2 antibody (Herceptin), monoclonal anti-interleukin-6 antibody, trastuzumab (Herceptin)]

PAMAM, polyglycerol dendrimers, PPI, etc.

Aptamer

PAMAM, pyrrolemodified dendrimers, DNA dendrimers

applications drug delivery • Delivery of contrast agent • High transfection efficiency (early endosomal escape) • Could assist in designing of biocompatible, biodegradable dendrimers with negligible toxicity • Targeted imaging and diagnosis of diseases such as cancer • High intracellular uptake, better retention in systemic circulation • Transdermal delivery • Vaccine delivery • As a photosensitizer for photodynamic therapy • As nanodiagnostic or nanotherapeutic or nanotheranostic agent for targeting and imaging cancer cells • As bioimaging probe • Immune therapy of cancer cells • Increased cytotoxicity toward cancer cells • Suppression of tumor growth • Targeted delivery • Reduced toxicity including hemolytic-, hepato-, and nephrotoxicity • Treatment of drug abuse (e.g., methamphetamine) • Controlled and targeted delivery of drugs (e.g., anticancer drugs) • To inhibit growth, migration, invasion of cancer cells • Induction of apoptosis of cancer cells • Could be developed as aptasensor for diagnostic applications [detection of thrombin, ochratoxin A (a carcinogenic mycotoxin that contaminates food such as cereals, wine, and beer)] • Targeted

[5,124,177]

[47,55,57, 178–180]

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Table 7.2 

Surface engineering strategies for dendrimers with their prospective bioinspired applications—cont’d Surface engineering

Dendrimers

Prospective applications

References

Chitosan

PAMAM, magnetic cored dendrimers

• Enhanced

[156,158, 181,182]

Fluorophore/diagnostic/imaging agents

PLL, PAMAM, polyester, dendrimers

efficiency of anticancer drugs • pH-sensitive drug release • Relatively more hemocompatible carrier than plain dendrimers • Oral delivery of drugs • As amphoteric adsorbant • Removal of heavy metalions • To track pathways for cellular entry, and internalization and trafficking • Effect of conjugation site and mechanism on cellular trafficking • As imaging and diagnostic agent • As delivery agents for nucleic acids (siRNA, antisense oligonucleotides, gene) • Biocompatible theranostic biosensors • Target-specific delivery of imaging and contrast agents

Amino acid (lysine, arginine, ornithine, histidine, tryptophan)

PAMAM, PPI, PLL, polyglycerol, polyester dendrimers

• High-efficiency

[47,60, 67,184]

Glycoprotein/lipoprotein (transferrin, LDL, HDL)

PAMAM, PPI, telodendrimers

transfection agents • Safe and effective drug and gene delivery vehicle with negligible toxicity • Biocompatible and biodegradable drug delivery system • Antiangiogenic activity • siRNA delivery • Dendrimer-based antidotes to treat cases of poisoning (e.g., in pesticide dichlorvos poisoning to maintain acetylcholinesterase activity) • Targeted therapy of cancer • Targeted gene delivery • Targeted delivery to brain tumor and brain-related diseases

[124,153,183]

[185–189]

Continued

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Table 7.2 

Surface engineering strategies for dendrimers with their prospective bioinspired applications—cont’d Surface engineering

Dendrimers

Prospective applications

References

Dexamethasone

PAMAM

• Enhanced ocular permeability, could

[159,190,191]

Cyclodextrins (CDs) (glucuronylglucosyl-β-CDs, α-CD)

PAMAM, thioalkylated mannose-modified dendrimer

be explored in retinal delivery • Gene therapy of diseases such as ischemic stroke • Gene transfection • Enhanced antiinflammatory effect • Improved stability and transfection efficiency of dendriplexes • siRNA delivery

[157,192,193]

GATG, gallic acid–triethylene glycol; HDL, high-density lipoprotein; LDL, low-density lipoprotein; PAMAM, polyamidoamine; PEG, polyethylene glycol; PLL, poly-l-lysine; PPI, poly(propyleneimine); siRNA, small interfering RNA.

stable complex than the other metabolite endoxifen and anticancer drug tamoxifen, whereas the strongest binding occurred with 4.0G PAMAM dendrimers followed by mPEG-anchored 3.0G PAMAM dendrimers and then PEG-6000. Among all the anticancer drugs and metabolites included in this report, doxorubicin showed the highest binding affinity for 4.0G PAMAM dendrimers [121]. The interactions between the host molecule, i.e., dendrimers, and guest molecule, i.e., drug, gene, depends on many factors, including charges of dendrimers and the host molecule, surface engineering, generation number, and type of dendrimers.

7.4.1  Dendrimers as vesicles: physical encapsulation of drug Presence of multiple internal voids or cavities in the dendrimers allows encapsulation of many guest molecules into a single molecule of dendrimer (Figs. 7.1 and 7.4). Kavyani et al. [58] investigated the structural properties and encapsulation efficiency of core (PPI)–shell (PAMAM) hybrid dendrimers for drug delivery via molecular dynamics simulation. In this study, the researcher used pyrene as the guest molecule for loading efficiency. The outcome of this research showed that the guest molecule, pyrene, was encapsulated in the dendrimers with little and no influence of PPI core size. Furthermore, the author speculated that the branching chain of PPI dendrimers hinders the entry of guest molecule into PPI, which could be sorted out by adding PAMAM dendrimers on the surface of PPI dendrimers, which resulted in the enhanced encapsulation efficiency for pyrene in this research work [58]. The high drug-loading capacity of dendrimers has also been confirmed by computational studies. The effect of structural properties and pH condition on loading efficiency of dendrimers has been investigated by scientists [122,123]. In a comparative molecular dynamic simulation study between acetyl-terminated and amine terminated 4.0G PAMAM dendrimers, the effect of pH conditions and structure of dendrimers on the encapsulation efficiency for a Biopharmaceutical Classification System class

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II antidiabetic drug, nateglinide, was investigated [122]. The encapsulation efficiency was found to be influenced by pH, and higher loading was observed at low pH for both dendrimers. At low pH, drug molecules were found to be located in the interior hydrophobic pockets of dendrimers, whereas at neutral pH, drug molecules were found to be distributed at the surface and interior hydrophobic cavities of acetylated and amine-terminated dendrimers, respectively, as investigated using center-of-mass distance analysis. Although the loading efficiency was not found to be influenced significantly by location of drug molecules in dendrimers, it was found to be affected by intermolecular interactions, charges on dendrimers and drug molecules, changes in the conformation of dendrimers induced by pH, and the surface groups of dendrimers [122].

7.4.2  Dendrimers as template: chemical conjugation of drug The presence of multiple surface functional groups gives the opportunity for the functionalization of dendrimers with multiple molecules (Figs. 7.4 and 7.5). Fig. 7.5 shows simultaneous anchoring of dendrimers on graphene oxide sheets and conjugation with gadolinium diethylenetriaminepentaacetate (Gd-DTPA) and prostate stem cell antigen (PSCA) monoclonal antibody [83,124]. Kolhatkar et al. [83] developed a dendritic hexapeptide with four arms with the aim of developing a biodegradable system that can specifically release the highly cytotoxic drugs in niche of tumor site by its own degradation. To design such type of system Kolhatkar and coworkers used a tetrapeptide sequence that is recognized by cathepsin B enzyme. This site-specific self-degradable system was developed to deliver an inhibitor of heat shock protein 90, geldanamycin, selectively to cancer cells. Geldanamycin was conjugated to terminal end groups of this degradable dendritic hexapeptide, and the drug was preferentially released in the presence of cathepsin B. Cathepsin B is overexpressed by various cancer cells, and hence the investigator concluded that such type of system could be very helpful in designing of a proper controlled drug delivery system for cancer therapy [83].

Figure 7.4  Figure showing encapsulation of guest molecule into interior cavities and conjugation/electrostatic interaction with surface groups into dendrimer host molecule.

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Biopolymer-Based Composites

&22+ ('&1+6

'73$GD

*G

1(W

*O XW

DUD OGH K\ GH

&22+

2 2 1 +

1

2 1

2 2

2

1 *G

2+

P$E36&$

2

2 2

2+

Figure 7.5  Schematic illustration of hybrid of graphene oxide sheets with dendrimers conjugated with contrast agent Gd-diethylenetriaminepentaacetate (DTPA) (for imaging) and monoclonal antibody (mAb) as targeting moiety. EDC, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride; NHS, N-hydroxysuccinimide; PSCA, prostate stem cell antigen. Reprinted from Guo L, Shi H, Wu H, Zhang Y, Wang X, Wu D, et al. Prostate cancer targeted multifunctionalized graphene oxide for magnetic resonance imaging and drug delivery. Carbon 2016;107:87–99. Copyright (2016) with permission from Elsevier.

Nanaware-Kharade et al. [5] investigated dendrimers as a platform for binding antimethamphetamine antibody (antimethamphetamine scFv7F9Cys) fragments and to improve their pharmacokinetic profile to devise a strategy for promising treatment of methamphetamine abuse. PEG was used as a linker to attach the antibodies to the surface of dendrimers. The antibody-conjugated dendrimers, dendribodies, were given to male Sprague-Dawley rats (which were already administered subcutaneous infusion of methamphetamine) by intravenous route as bolus injection. As a conjugate with dendrimers the volume of distribution and clearance of antibodies were decreased 1.6 and 45 times, respectively, whereas the terminal half-life was increased by about 20-fold, suggesting that dendribodies improved the pharmacokinetic behavior of antibodies [5].

7.5  Characterization of dendrimers Pertaining to step-by-step synthesis by chemical reaction and polymeric nature, dendrimers have been characterized for their morphology, size, shape, surface area, molecular weight and distribution, surface functional groups, purity, and molecules

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conjugated to dendrimers by different methods including spectroscopy, spectrometry, microscopy, chromatography, and rheological evaluation [120,125,126]. Different techniques are used for the characterization of dendrimers at various levels, including • Techniques for confirming synthesis of dendrimers • Characterization of synthesized dendrimers for shape, size, generation, molecular weight, functional groups • Characterization of surface engineered dendrimers • Methods for analyzing the interactions between dendrimers and drug molecules • Characterization of dendriplexes, dendrimersomes, dendribodies, and domino dendrimers

A large number of analytical and characterization techniques have been used in dendrimer chemistry, such as spectroscopic methods (Fourier transform infrared, nuclear magnetic resonance, and Raman and adsorption spectroscopy) microscopy methods (transmission electron microscopy, scanning electron microscopy, atomic force microscopy, and confocal laser scanning microscopy), chromatographic methods (liquid chromatography, adsorption chromatography, gel permeation chromatography, high-performance liquid chromatography, liquid chromatography-mass spectrometry, and gas chromatography), elemental analysis, thermal analysis (differential scanning calorimetry, differential thermal analysis, thermogravimetric analysis), scattering technique (dynamic light scattering, small-angle X-ray scattering, and small-angle neutron scattering), electrophoresis, mass spectrometry, rheological evaluation, and circular dichroism [28,103,126–131]. A summary of the different characterization methods used in dendrimer chemistry is given in Table 7.3.

7.6  Dendrimer hybrids with other nanocarriers Scientists are exploring combination of two or more nanomaterials for various drug delivery applications. Dendrimers are also being explored in combination with other nanomaterials to develop a prospective hybrid nanocarrier system.

7.6.1   Carbon-based nanomaterials–dendrimers The hybrids of carbon nanomaterials (CNTs, carbon nanohorns, graphene sheets, fullerenes) and dendrimers are extensively being explored for various biomedical and industrial applications [23,124,132–137]. Yang et al. [23] investigated single-walled carbon nanotubes (SWCNTs) with surface-anchored dendrimers for drug delivery and biological applications. In this investigation scientists conjugated dendrimers bearing zinc-phthalocyanine chromophores on the surface of SWCNTs. Investigation carried out to examine photoinduced electron transfer and photophysical properties with ultraviolet–visible and fluorescence spectroscopy showed these hybrid conjugates to be superior over individual nanocarriers, suggesting their promising potential in drug delivery as biological labeling ­applications [23].

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Table 7.3 

Brief summary of analytical techniques used for characterization of dendrimers [125,126,130,131] Method of characterization Spectroscopic methods

Fourier transform infrared spectroscopy

Raman spectroscopy Ultraviolet–visible spectroscopy Nuclear magnetic resonance (NMR)

Fluorescence spectroscopy

Circular dichroism and optical rotation

Scattering and diffraction techniques

X-ray diffraction

Laser light scattering

Important applications in dendrimer’s characterization Characterization of branching monomer unit, core molecule, surface end groups, synthesized dendrimers, surface engineered dendrimers, surface transformation [conversion of nitrile groups to amino groups during synthesis of poly(propyleneimine) dendrimers], interaction with drug molecule, H-bonding • Synthesis of dendrimers • Degree of cyclodehydrogenation Synthesis and purity of dendrimers • Step-by-step

synthesis of dendrimers, including high-generation dendrimers • Using special NMR techniques [spin-lattice relaxation time (T1) of protons, T2 (spin–spin) relaxation times] size, morphology, and dynamics of dendrimers could also be determined • Characterization of dendrimer-entrapped nanoparticles or quantum dots • To quantify defects in the synthesis of dendrimers • To study the structure and molecular dynamics of outer surface of dendrimers • Dendrimers containing chiral groups • Characterization of amino acid and peptide dendrimers or amino acid and peptide-conjugated dendrimers • Interaction of dendrimers with protein drug molecules or ligands • Chemical

composition, size, and shape of dendrimers (but mostly dendrimers are amorphous solids) • Structure of lower generation of dendrimers • Determination of hydrodynamic radius of dendrimers • Indication on molecular weight of dendrimers

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Table 7.3 

Brief summary of analytical techniques used for characterization of dendrimers [125,126,130,131]—cont’d Method of characterization Dynamic light scattering

Small-angle X-ray scattering (SAXS)

Small-angle neutron scattering

Mass spectrometry (MS)

Microscopy

Chemical ionization MS Fast atom bombardment MS Electron spray ionization MS Matrix-assisted laser desorption ionization time of flight Liquid chromatography–MS Inductively coupled plasma MS Transmission electron microscopy (TEM) Scanning electron microscopy (SEM) Atomic force microscopy (AFM) Confocal laser scanning microscopy (CLSM)

Important applications in dendrimer’s characterization • Average

particle size and particle size distribution • Hydrodynamic radius • Polydispersity index • Zeta potential • Determination of average radius of gyration (Rg) of dendrimers in solution • Structural arrangement of dendrimers • Changes in the shape of dendrimers from lower to higher generations • Determination of average radius of gyration (Rg) with a more detailed information on internal structure of dendrimers than SAXS • Information on molecular weight of dendrimers and density distribution within the dendrimeric architecture • Determination of molecular weight of dendrimers • Synthesis and purity of dendrimers • Detection of imperfections and chemical defects in high-generation dendrimers • Number of generations of dendrimers • Mass of surface engineered dendrimers and number of ligand molecules conjugated to the surface of dendrimers • Diameter and particle size distribution • To visualize a single molecule of dendrimers and determine size, shape, and morphology • Information on surface and internal structure of dendrimers • Change in size on conversion of lower to higher-generation dendrimers and on surface engineering of dendrimers • Arrangement of the atoms and microstructure in their niche could be determined using a highly advanced version of TEM, high-resolution transmission electron microscopy • 2-D and 3-D images of dendrimers with spatial variations could be seen in SEM and AFM, respectively Continued

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Table 7.3 

Brief summary of analytical techniques used for characterization of dendrimers [125,126,130,131]—cont’d Method of characterization Chromatography

Size exclusion chromatography High-performance liquid chromatography

Thermal analysis

Differential scanning calorimetry Differential thermal analysis Thermal gravimetric analysis Electron paramagnetic resonance

Miscellaneous

Electrochemical techniques (coulometry, voltammetry)

Gel electrophoresis

Important applications in dendrimer’s characterization • Separation

of different generations of dendrimers on the basis of size • Purification of dendrimers • Separation of reactants from synthesized dendrimer product • Determination of approximate molecular weight of dendrimers • Polydispersity index • Effect of pH on size of dendrimers • Stability • Interaction with drug molecule • Characterization of surface engineered dendrimers • Physical properties of dendrimers • Compatibility studies • Number of groups substituted or conjugated at surface of dendrimers • Determination of interactions among surface groups of dendrimers • Number of electrically active groups on surface of dendrimers • Quantitative amounts of electrically active groups as a part of internal structure of dendrimers • Possible electrostatic interactions of dendrimers with other electrically active groups • Determination of homogeneity and purity of water-soluble dendrimers • Study of dendriplexes in terms of interactions between dendrimers and DNA

Guo et al. [124] grafted 3.0G PAMAM dendrimers with surface amino groups to graphene oxide nanosheets followed by sequential conjugation of contrast agent Gd-DTPA and targeting moiety PSCA monoclonal antibody with the surface amino groups of dendrimers to develop a cancer cell–targeted diagnostic agent (Fig. 7.5). The developed nanoconjugate showed reduced toxicities, including hemolytic toxicity and cytotoxicity, as observed in in vivo studies carried out in male Kunming mice. The in vivo MRI and anticancer activity studies showed that the developed hybrid nanocarrier increased the contrast effect of the diagnostic agent and targeted the contrast

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agent Gd-DTPA to the prostate cancer cells with selective delivery of loaded anticancer drug doxorubicin hydrochloride to cancer cells, leading to an improved and targeted anticancer activity and diagnostic efficiency paving the way for promising applications of theranostic agents in the treatment and diagnosis of cancer. Hybrids of dendrimers with graphene oxide also showed promising results in codelivery of gene and drug [124,137].

7.6.2   Lipid–dendrimer hybrids Lipid-based carrier systems are extensively being investigated to develop biocompatible and biodegradable drug delivery systems. One of the major problems in the clinical applications of dendrimers is anticipated to be their cationic charge–associated cytotoxicity. Dendrimers are being investigated with liposomes as hybrid carrier system to reduce toxicity of the dendrimer, increase loading efficiency, increase intracellular drug delivery, and enhance targeting efficiency [36,37,138–141]. Nanohybrid conjugates synthesized from PEG, dendrimers, and phospholipids have shown an increase in stability and circulation half-life of liposomes with improved biodistribution and pharmacokinetic profile of the encapsulated drug along with an increase in anticancer efficacy and enhancement of oral absorption of drugs, doxorubicin hydrochloride and probucol, respectively [138,139]. Wang et al. [140] developed the dendrosomes (simultaneously sensitive to two external stimuli, magnet and pH) using magnetic nanoparticles and folate-targeted dendrimers, which were encapsulated into long-circulating pH-sensitive liposomes. These hybrid nanocarriers showed high cellular uptake in folate receptor–positive HeLa cells because of the presence of folic acid on the surface of dendrimers. Furthermore, the developed nanohybrid system was found to be sensitive to acidic pH and showed release of loaded cargo at acidic pH with cytotoxicity and cellular uptake at slightly acidic pH to the extent comparable with free folate–conjugated dendrimers. In vivo tumor xenograft mice model study showed that dendrosomes showed long circulatory time with prompt release stimulated by the slightly acidic environment of the tumor and application of magnetic field [140]. Hybrids of dendrimers with liposomes have also shown superior transfection efficiency with reduced toxicity [36,141].

7.6.3   Quantum dots–dendrimers Hybrid systems of quantum dots with dendrimers and other nanomaterials have been designed by scientists for various biomedical applications, mainly for theranostic applications particularly in cancer, to develop nano- and biosensors and to design nanomaterial-based antimicrobial agents [142–146]. Silicon quantum dots coated by PAMAM dendrimers having surface hydroxyl groups have shown the ability to detect Cr(VI) as nanosensor [142]. An aptasensor for the detection of platelet-derived growth factor BB was developed by functionalizing a hybrid nanocarrier made up of PAMAM dendrimers and CdS quantum dots with aptamer [20,143]. A multiple hybrid system of PAMAM

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dendrimers, multiwalled carbon nanotubes (MWCNTs), and quantum dots (CdS and Ag2S) has been developed, and in antimicrobial activity assay (tested against E. coli, Pseudomonas aeruginosa, and S. aureus), it was observed that the PAMAM dendrimers grafted on MWCNTs enhanced the germicidal action of CNTs. The CdS and Ag2S immobilized in PAMAM dendrimer-grafted MWCNTs further enhanced the germicidal activity [146].

7.6.4  Nanoparticles–dendrimers Hybrids of nanoparticles and dendrimers have been developed for various biomedical applications, including controlled and targeted drug delivery, diagnostic and theranostic applications, gene delivery, and stimuli-controlled delivery [7,37,110,147–149]. Dendrimer-entrapped gold nanoparticle hybrid carriers conjugated with α-tocopheryl succinate have shown promising results for targeted imaging and therapy of cancer [147]. Conjugates of DNA, gold nanoparticles, and PAMAM dendrimers have shown promising applications in cancer diagnosis in preliminary investigations [148]. A hybrid nanocarrier made up of magnetic nanoparticles functionalized with temperature-sensitive hyperbranched poly(epsilon-lysine) peptide dendrons showed controlled release of vascular endothelial growth factor in response to mild hyperthermic pulses generated by the application of external magnetic field [7]. Magnetic nanoparticles with dendrimers have also shown promising results in preclinical investigations for targeted delivery mediated by external stimuli [37,150].

7.6.5  Miscellaneous nanohybrids based on dendrimers Microemulsions and dendrimers are being investigated for their possible applications as hybrid nanocarriers. Lidich et al. [62] speculated that solubilization of 2.0G PPI dendrimer and a neutraceutical, triglyceride of docosahexaenoic acid (TG-DHA), into microemulsion can potentiate membrane permeation and drug delivery efficiency. The results of this research showed that microemulsion can solubilize 2.0G PPI and TG-DHA with retention of microviscosity, micropolarity, structure, and stability and can be explored positively in the delivery of drugs and neutraceuticals [62]. Torres et al. [151] functionalized TiO2 nanotubes with PAMAM dendrimers using (3-glycidoxypropyl)methyldiethoxysilane as the coupling agent. These PAMAMfunctionalized TiO2 nanotubes were then evaluated for drug delivery applications using three model drugs, curcumin, methotrexate, and silibinin. The result of this study showed that these hybrid nanocarriers enhanced the drug-loading capacity as well as prolonged the duration for drug release [151]. Campos et al. [152] grafted 3.0G PAMAM dendrimers to the surface of alumina tubes and evaluated them for drug delivery applications. From various analytical studies, investigators observed that 3% by weight dendrimers were conjugated to the surface of the tube. Grafting of PAMAM dendrimers over alumina tubes increased the loading efficiency of the system with decrease in cytotoxicity, suggesting that the PAMAM dendrimers–Alumina tube is a prospective drug delivery vehicle [152].

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7.7  Biomedical applications of dendrimers Engineering of dendrimeric architecture provides opportunity to develop a promising biomaterial having unique properties of nanomaterials with high drug-loading capacity. The dendritic polymers, including dendrimers and hyperbranched polymers, could emerge as promising drug delivery vehicles in future because of the presence of large number of tailor-made functional groups with low viscosity and exceptional intrinsic stability [32,67,91,153–159] (Table 7.4).

7.7.1  Combined delivery of drugs As discussed earlier, one of the important advantages of dendrimers as nanocarrier is the presence of multiple surface functional groups that enable conjugation of a variety of moieties at the surface of dendrimers, including targeting ligands, bioactives, diagnostic agents, and PEGylation. Furthermore, the presence of multiple hydrophobic cavities in the dendrimers enables the encapsulation of different molecules into these cavities (Fig. 7.4). PAMAM dendrimers have been explored for the combined delivery of a diuretic drug, hydrochlorothiazide, and an antihypertensive drug, ramipril [45] (Fig. 7.6). The complexes of dendrimers and drugs were formed by the phase equilibration method. Results of this research showed that the solubility of drugs was influenced by the concentration of dendrimers and the pH of dendrimer solution. A 0.8% w/v solution of dendrimers showed a 4.91and 3.72-fold increase in solubility of ramipril and hydrochlorothiazide, respectively, amine- and carboxy-terminated dendrimers, respectively. In the in vitro dissolution studies, dendrimeric formulations of ramipril and hydrochlorothiazide showed rapid and complete dissolution as also observed with a hybrid formulation containing both the drugs. The formulation was also found to be stable in dark and refrigerated conditions. All these results suggested that dendrimers could be developed as a versatile nanoscaffold for combined delivery of multiple drugs for therapeutic applications [45].

7.7.2   Brain delivery Designing of a homing device for efficient brain-targeted delivery of chemotherapeutic agents across the BBB is the prime requirement for the treatment of the most common as well as aggressive disorder of brain, glioblastoma multiforme. The major hurdle posed in the therapy of glioma is the BBB, which restricts the entry of chemotherapeutic agents, and poor retention in the tumor tissue results in negligible or subtherapeutic concentration of drug at the target site. Dendrimers have been explored to sort this problem, and the glioma homing peptide (Pep-1)–conjugated PEGylated PAMAM dendrimers have been explored to overcome these two hurdles in brain-targeted delivery. The developed conjugate showed efficient delivery across the blood–brain tumor barrier by IL-13 receptor α2–mediated endocytosis as well as because of the small size of dendrimers. Furthermore, suitable biological interactions of dendrimers for penetration into tumor were also observed with U87MG cells. The results of this

Potential biomedical applications of dendrimers

194

Table 7.4 

Ligand/ nanocarrier

Bioactive

Objective

Results

References

Mannose dendrimers



-–

To designate the structure of amphiphilic Janus dendrimers

• Study

[194]

PPI dendrimers (3.0 and 4.0G)

Maltose

Cytidine5′-triphosphate (CTP)

5.0G PPI dendrimers

Sialic acid, glucosamine and concanavalin A

Paclitaxel

To investigate developed glycodendrimers as prospective drug delivery vehicle Compared sialic acid, glucosamine, and concanavalin A as ligand for brain-targeted delivery

showed that these amphiphilic Janus dendrimers self-assemble to form onionlike vesicles denoted as dendrimersomes • The assembly as dendrimersomes was obtained by injecting the solution of these amphiphilic Janus dendrimers in aqueous solvents, including water, water-miscible solvents, and buffers • Furthermore, dendrimersomes were found to mimic the biological membranes and the number of bilayers in dendrimersomes was found to be influenced by concentration of Janus dendrimers in water Complexes between dendrimers and CTP were developed with high efficiency

Brain-targeting potential of ligands was found to be in the following order: sialic acid > glucosamine > concanavalin A

[119]

[174]

Biopolymer-Based Composites

Dendrimer

Folic acid

3,4-Difluorobenzylidene diferuloylmethane (CDF)

• To

enhance the aqueous solubility of hydrophobic bioactive • Specific delivery to folate receptor overexpressing cancer cells

1.0 and 3.0G PAMAM



Calf-thymus DNA

To design effective nonviral systems for DNA delivery

• Dendrimers

were found to solubilize a highly hydrophobic anticancer bioactive, which is a flavonoid analog, CDF • Folate-conjugated dendrimers showed higher accumulation of drug in folate receptor– overexpressing cervical cancer cells (HeLa) and ovarian cancer cells (SKOV3) cells • Properties of dendriplex were found to be dependent on cationic polymers under investigation • Authors suggested that the more relaxed state complexation in the case of PEGylated cationic block copolymers could result in superior transfection efficiency as compared with lower-generation dendrimers • Authors also stated that higher generations (4.0–7.0G) of PAMAM dendrimers showed better transfection efficiency as compared with lower-generation dendrimers

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

4.0G PAMAM dendrimers

[44]

[56]

Continued 195

Potential biomedical applications of dendrimers—cont’d

Dendrimer

Ligand/ nanocarrier

196

Table 7.4 

Objective

Results

References

Examining the effect of dendrimer physicochemical properties as well as pathophysiology associated with neuroinflammation disease on brain uptake, diffusion, and cell-specific localization of dendrimers To study fetal–maternal transport and neuroinflammation targeting of intraamniotically administrated dendrimers with neutral/anionic surface functional groups

• Neutral

dendrimers showed rapid localization into glial cells in regions of injury • The degree of BBB breakdown, glial cell activation, and severity of disease influenced the extent of uptake of dendrimer as an imaging biomarker as measured using rabbit model of cerebral palsy

[161]

Both neutral and anionic PAMAM dendrimers were absorbed by fetuses and demonstrated bidirectional transport between fetuses and mother, but dendrimers were found more efficient in crossing through fetal blood– brain barrier, and targeting activated microglia

[21]

Hydroxyl-terminated 4.0G PAMAM dendrimer, amine-terminated dendrimer (G4-NH2), 3.5G carboxylate dendrimer

Cy5, a near-infrared imaging agent

-–

Dendrimers

Hydroxyl (neutral) and carboxyl (anionic) functional groups



Biopolymer-Based Composites

Bioactive

Amine and acetyl surface groups

Nateglinide

To investigate the effect of structural properties of dendrimers on its drug-loading efficiency using molecular dynamic simulation studies

• Acetylated

and amine-terminated PAMAM dendrimers encapsulated five and six molecules of drug, respectively, at neutral pH • At low pH, acetylated and amine-terminated PAMAM dendrimers encapsulated 13 and 12 molecules of drug, respectively. • In the case of amine-terminated PAMAM dendrimers it was observed that most of the drug molecules were located in the interior hydrophobic pockets of dendrimers at both the pH as observed in centerof-mass distance analysis • In the case of acetylated PAMAM dendrimers, most of the drug molecules were found to be distributed near the surface at neutral pH and in the interior hydrophobic pockets at low pH

[122]

Continued

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

4.0G PAMAM

197

Potential biomedical applications of dendrimers—cont’d

Dendrimer

Ligand/ nanocarrier

Bioactive

Objective

198

Table 7.4 

Results • Authors

Polyglycerol core multishell hyperbranched dendritic polymers

Finasteride, dexamethasone

To investigate as prospective dermal drug delivery system

[91]

Biopolymer-Based Composites

Octadecen1-yl succinic anhydride and methyl poly(ethylene-glycol) (mPEG500) shell

also reported that at low pH, encapsulated drug molecules in the amine-terminated PAMAM dendrimers were present in cluster form, whereas in the case of nontoxic acetylated PAMAM dendrimers, they were uniformly distributed inside the dendritic cavities, suggesting acetylated PAMAM dendrimers as suitable delivery vehicle for nateglinide • Increase in the hydrophobicity of the amphiphilic shell resulted in significant increase in the loading efficiency of these systems • A 6.2% and 10.3% (by weight) loading capacity was observed for drug cargos, dexamethasone and finasteride, respectively.

References

4.0G PPI dendrimers

Maltotriose

Fludarabine

To investigate unique gene expression signature in the B-cell receptor (BCR) signaling pathway during treatment of chronic lymphocytic leukemia (CLL) with maltotriose-modified dendrimers

PPI dendrimers

Maltose

Cytarabine

To facilitate the delivery of activated cytarabine to cancer cells to overcome metabolic limitations of the drug as well as to overcome the problem of resistance

[195]

[162]

199

cytotoxicity studies with human keratinocyte cells HaCaT, the developed hyperbranched core–shell dendritic system displayed no cytotoxicity up to concentration of 0.05 mg/mL. But the system was found to be toxic at the concentration of 0.5 mg/mL for HaCaT cells • Maltotriose-modified dendrimers were found to influence BCR pathway gene expression effectively with possibility of inhibition of clonal expansion • Authors also reported that these maltotriose-modified PPI dendrimers could also be helpful in treatment of patients of CLL with chromosomal alterations Enhanced uptake by human leukemia cell lines 1301 with enhanced cytotoxicity was observed for active drug and glycodendrimer complex with blocked nucleoside transporter hENT1, suggesting efficacy of this complex against resistant acute lymphoblastic leukemia cells with lower expression of hENT1

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

• In

Continued

Potential biomedical applications of dendrimers—cont’d

200

Table 7.4 

Ligand/ nanocarrier

Bioactive

Objective

Results

References

Amphiphilic dendrimers based on PAMAM dendrimers





To design an ideal gene transfection system by combining the physical and chemical properties of polymers and cationic lipids

• Amphiphilic

[196]

1.0 and 1.5G PAMAM dendrimers

Tris(2,2′bipyridine) ruthenium(II)

5-Fluorouracil (5-FU)

Exploration of PAMAM sensors in the presence of a suitable sensitizer

Dendrimer

[33] Biopolymer-Based Composites

dendrimers were developed using a low-generation hydrophilic dendron attached to hydrophobic tail • Among the developed amphiphilic dendrimers, dendrimers with sufficiently long-chain hydrophobic tail (18 carbons) and second generation (among 0.0, 1.0 and 2.0G) showed efficiency as siRNA transfecting agent. • The dendrimers showed self-assembling ability to form aggregates • 1.0 and 1.5G PAMAM dendrimers were reacted with tris(2,2′-bipyridine)ruthenium(II) via electrogenerated chemiluminescence (ECL) reactions at pH 6.1 and 10.0 in an aqueous medium in the presence and absence of 5-FU • The efficiency of ECL was found to be dependent on binding between dendrimer and 5-FU

PAMAM dendrimers

To study interaction of dendrimer and biological membrane interactions to elucidate the possibility of prospective biological application or problems of disruption of biomembrane and cytotoxicity associated with the interaction

Pluronic F127

Doxorubicin

To modify dendrimers using Pluronic F127 in such a way that it will work as unimolecular micelles to reverse multidrug resistance

• Interactions

of positively charged [3.0G amino (dNH2) terminated] and negatively charged [2.5G sodium carboxylate (COOd Na+) terminated] dendrimers with dipalmitoylphosphatidylcholine liposomal bilayers were studied • Interactions between dendrimers and liposomal model bilayers showed increase in zeta potential of liposomal surface and hydrophobic regions of bilayers • Dendrimer penetration into liposomal bilayers produced the perturbation of the hydrophobic alkyl chains of the bilayers • Increase in anticancer efficiency of doxorubicin (DOX) was observed when delivered as a conjugate with dendrimers toward multidrug-resistant breast cancer cells (MCF-7/ADR).

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

PAMAM dendrimers

[34]

[42]

Continued 201

Potential biomedical applications of dendrimers—cont’d

Dendrimer

Ligand/ nanocarrier

Bioactive

Objective

202

Table 7.4 

Results • A

TiO2 nanotube

Curcumin, methotrexate, and silibinin

To devise a prospective drug delivery vehicle

[151]

Biopolymer-Based Composites

3.0G PAMAM dendrimers

fivefold increase in uptake of DOX was observed with conjugate in comparison to free DOX • Increase in apoptosis with more accumulation and distribution in the nuclei was observed after treatment with DOX-loaded dendrimeric conjugates • Finally it was concluded by the investigator that DOXloaded Pluronic F127PAMAM conjugates could result in a superior alternative for treatment of resistant cancer • Improved drug-loading capacity • Prolonged drug release, which was diffusion controlled • Developed hybrid nanocarrier showed negligible toxicity with a nearly 100% viability of cell culture

References

Magnetic nanoparticles and liposomes

Rhodamine B isothiocyanate

To develop a pH- and magnetic-responsive targeted drug delivery system, dendrosome, to deliver anticancer drug safely and selectively to cancer cells

3.0G PAMAM dendrimer

Graphene oxide

DOX

To treat the multidrug resistance of tumor cells in drug and gene delivery

• Vehicles sensitive to two exter-

nal stimuli, magnet and pH, were developed by encapsulating magnetic nanoparticles and folate-targeted dendrimers into long-circulating pH-sensitive liposomes • Higher cellular uptake into folate receptor–positive HeLa cells • Cytotoxicity and cellular uptake at slightly acidic pH • Prompt release of loaded cargo at acidic pH and presence of magnetic field • Prolong residence time in tumor xenograft mice model Graphene oxide–PAMAM hybrid exhibited lower cytotoxicity with good in vivo biocompatibility showing promising potential of this graphene and dendrimeric hybrid nanocarrier in drug and gene codelivery

[140]

[137]

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

Folate-conjugated 4.0G PAMAM dendrimers

BBB, blood–brain barrier; PAMAM, polyamidoamine; PPI, poly(propyleneimine); siRNA, small interfering RNA.

203

204

Biopolymer-Based Composites

RAPL

Step-I PAMAM dendrimers

RAPL-dendrimer complex

Step-II

HCTZ

Hybrid drug-dendrimer complex

Step-I PAMAM dendrimers

HCTZ-dendrimer complex

Figure 7.6  Schematic diagram showing strategy for the development of combination therapy of ramipril (RAPL) and hydrochlorothiazide (HCTZ) using polyamidoamine (PAMAM) dendrimers. Reprinted from Singh MK, Pooja D, Kulhari H, Jain SK, Sistla R, Chauhan AS. Poly (amidoamine) dendrimer-mediated hybrid formulation for combination therapy of ramipril and hydrochlorothiazide. European Journal of Pharmaceutical Sciences 2017;96:84–92. Copyright (2017) with permission from Elsevier.

study suggested the promising opportunities in the development of Pep-1 engineered dendrimers for brain tumor–targeted delivery [160]. Nance et al. [161] observed that PAMAM dendrimers with surface hydroxyl groups showed preferential localization into the regions of BBB impairment (Fig. 7.7; Table 7.4).

7.7.3  Ability to overcome drug resistance Szulc et al. [162] investigated maltose-conjugated 4.0G PPI glucodendrimers as delivery vehicle for anticancer drug cytarabine to treat leukemia and overcome drug resistance associated with cytarabine. The efficiency of this anticancer drug is also limited by its inability to accumulate inside the cancer cells where it is converted to its active analog, cytosine arabinoside triphosphate. Szulc and coworkers have exploited the possibility of interactions between cationic surface charges of PPI dendrimers and cytarabine to improve the uptake and accumulation of drug in cancer cells as well as to overcome the problem of resistance. The complexes of drug with dendrimers showed improvement in anticancer activity against 1301 human leukemia cells with enhanced uptake. The delivery of drug with dendrimers also showed blockade of human equilibrative nucleoside transporter, hENT1, suggesting the possibility of efficacy of this complex to overcome the problem of resistance [162].

Dendrimers: smart nanoengineered polymers for bioinspired applications in drug delivery

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Figure 7.7 Images showing localization of hydroxyl-terminated 4.0G polyamidoamine dendrimers in regions of blood–brain barrier impairment and glial cell activation (where SVZ and PVR stands for subventricular zone and periventricular region, respectively). Reprinted from Nance E, Zhang F, Mishra MK, Zhang Z, Kambhampati SP, Kannan RM, et al. Nanoscale effects in dendrimer-mediated targeting of neuroinflammation. Biomaterials 2016;101:96–107. Copyright (2016) with permission from Elsevier.

7.7.4  Topical delivery system For treatment of skin ailment, local delivery seems to be a fascinating approach, and different delivery systems, including dendrimers, have been evaluated for this purpose. Polyglycerol-based core–shell dendritic nanovehicles are one of the transporters investigated for topical delivery because of their ability to increase penetration of lipophilic substances into skin without cytotoxicity toward keratinocytes [91]. PAMAM dendrimers showed promising results as topical delivery vehicle for antisense oligonucleotides in skin cancer mouse model [163]. Dendrimers have also been evaluated for topical and systemic delivery of photosensitizers [164]. Abnormal cell proliferation of skin layers have also been observed with PAMAM dendrimers in a report [165]. Surface engineering strategies or masking of toxic surface groups with biocompatible ligands could sort out this problem [27,166]. Dendrimers in combination with iontophoresis and sonophoresis have also shown an increase in skin permeation of drugs [153,163].

7.7.5   Stimuli-responsive dendrimers Wang et al. [54] modified the 4.0G PAMAM dendrimers via covalent conjugation of Ac-arg-ala-ala-asp-tyr-cys (RAADyC), PEG, and Arg-gly-asp-cys (RGDC) with surface groups of the dendrimers. Then these dendrimers were converted to enzyme-sensitive nanogel using sodium periodate oxidizing agent to start cross-linking of functional groups present on the surface of the dendrimers. The developed conjugate of dendrimers and the nanogel showed spherical shape with size in the nanometric range (20 and 50 nm, respectively, for conjugated dendrimers and nanogel). The nanogel showed a greater loading

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capacity than conjugated dendrimers. Furthermore, the shrinkage in size (about 15 nm) of doxorubicin-loaded nanogel was observed in the presence of enzyme, elastase, suggesting enzyme-triggered degradation of nanogel and sustained release of drug. The nanogel carrier was also found to be biocompatible and noncytotoxic, and the released doxorubicin showed a cytotoxicity comparable with that of free doxorubicin [54]. Li et al. [109] designed pH-responsive nanocarrier to overcome the problems associated with the delivery of anticancer drug, including restricted accumulation at tumor site, using PAMAM dendrimers. These nanocarriers were also designed in such a way that they can switch their size in the acidic environment of the tumor to improve tumor-selective penetration. The change in size was highly significant and sensitive to the acidic pH microenvironment of the tumor and a reduction in size from approximately 80 nm to 10 nm) is much thicker than that of the liposome (3–4 nm), which makes it hard to be oxidated and stable [98]. (2) The particle size, membrane thickness, membrane permeability, drug-loading capability, surface modification, and even in vivo behavior of polymersomes can be easily controlled by tuning the molecular weight, the block ratio, and the type of amphiphilic polymers [99]. Thus polymersomes are promising drug carriers for wide application in the field of drug delivery system. Polymersomes can be classified into biodegradable polymersomes and unbiodegradable polymersomes based on the biodegradability of amphiphilic polymers. The most typical biodegradable polymers to prepare biodegradable polymersomes are PEGylated polyesters, including PEG-PLA, PEG-PCL, PEG-PLGA. Unbiodegradable

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Biopolymer-Based Composites

polymers including PEO-poly(ethyl ethylene) (PEE), PEO-polybutadiene (PBD), PAApolystyrene (PS), and PBD-poly(glutamic acid) (PG) are used to prepare unbiodegradable polymersomes. Some polymersomes made of amphiphilic polypeptide such as PBD-PG, PEG-polylysine (PL) are called peptosomes. Polymersomes with unique physicochemical properties (particle size, membrane thickness, permeability, drug-loading capabilities, and surface modification) can be made by choosing amphiphilic polymers with different chemical properties (category, molecular weight, block ratio, and copolymer structure). Take PEO-PBD for instance. Polymersomes made by PEO-PBD cannot be degraded in organisms. As the molecular weight of the PBD block increases, the circulation time of the PEO-PBD polymersomes in blood prolongs and the thickness and stability of the membrane of PEO-PBD polymersomes also increase [100]. Inserting biodegradable polymers PEG-PLA or PEG-PCL into the bilayers of PEO-PBD polymersomes can help regulate the permeability of the PEO-PBD polymersome [101]. Theoretically, almost all the methods for preparing liposomes can be used to prepare polymersomes. Common methods preparing polymersomes include the injection method, the film hydration method, and the solvent precipitation method. For the injection method, the polymers were dissolved in the organic solvent and then injected into the aqueous phase, which can be deionized water or phosphate buffered saline (PBS), under vigorous stirring. It takes approximately 3 min for the formation of polymersomes. Stirring for 1–2 h is necessary to prepare polymersomes [102]. The organic solvent can be removed by repeated dialysis against water. The film hydration method is most widely used to prepare polymersomes. Similar to liposomes, polymers are first dissolved in dichloromethane and evaporated under vacuum to form a thin polymer film with the help of rotary evaporation. Afterward, PBS, sucrose solution, or water is added to hydrate the film for polymersome formation. The size and size distribution of polymersomes can be improved by ultrasonication, homogenization, and extruding after freezing and thawing. For the solvent precipitation method, the polymer is first dissolved in a suitable solvent and the aqueous phase (such as water) is added to form polymersomes via precipitation [103].

8.3.2.5  Solid lipid nanoparticles Solid lipid nanoparticles refer to nanoparticles mainly made of biocompatible highmelting-point lipids. As the dispersion properties of the lipids change during the preparation process, the lipids in the lipid nanoparticles are unlikely to remain in a solid form (stable crystal) but they usually are in the form of subfrozen and metastable crystals. As a result, in the latest papers, solid lipid nanoparticles are called lipid nanospheres as well. Because solid lipid nanoparticles are solid at room temperature, they have the advantages of polymer nanoparticles, such as high physical stability, less drug leakage, slow release, passive targeting, and ease of sterilization. At the same time, solid lipid nanoparticles also have some characteristics of liposomes, such as low toxicity, that make them apt for large-scale production [104]. Materials for preparing solid lipid nanoparticles include solid lipids, emulsifiers, and water. Solid lipids are commonly high-melting-point lipids, including glycerol esters of saturated fatty acids (such as stearic acid, capric acid, palmitic acid, and

Nanoparticles for tumor targeting

235

eicosanoic acid), steroids, and waxes. A wide variety of emulsifiers can also be used, including various phospholipids, sodium glycocholate, and synthetic emulsifiers [105]. The distribution of drugs in solid lipid nanoparticles is mainly determined by the nature of the drugs (melting point, polarity, etc.), lipid material properties, surfactant concentration, and other conditions. Generally, there are three main types of drug distribution in solid lipid nanoparticles. (1) Solid solution state: the drug is dispersed in the lipid material in the form of a single molecule. (2) Lipid core state: the drug is dispersed in the outer layer of solid lipid nanoparticles. (3) Drug core state: the drug is dispersed in the core. The preparation of solid lipid nanoparticles is not very simple. Especially when the optimization of the overall processes is poor, solid lipid microcapsules and solid lipid microspheres are actually what we get. There are four methods to obtain solid lipid nanoparticles, including the melt homogenization method, the cold homogenization method, the emulsification solvent evaporation method, and the nanoemulsification method. For the melt homogenization method, the amount of input lipids used is typically between 5% and 10%. The drugs are mixed with melted lipids, phospholipid, and surfactant and heated to above 70°C to melt the mixture. Through high-speed homogenizer, the molten mixture is made into colostrum. Through high-pressure homogenizer, colostrum is homogenized under the pressure of 50–150 MPa for three to five times. After cooling, the solid lipid nanoparticles are obtained. For the cold homogenization method, the drugs are dissolved and dispersed in the molten solid lipid. Liquid nitrogen or dry ice is used to freeze it into a fragile solid, while the drug is evenly dispersed in the solid lipid skeleton. Using the grinding method, solid microspheres are obtained with a particle diameter of 50–100 μm. Solid microspheres are then dispersed in a cooled emulsifier solution and homogenized several times at a temperature lower than the melting point of the lipid to obtain the solid lipid nanoparticles [106]. For the emulsification solvent evaporation method, the solid lipids are first dissolved in a water-insoluble organic solvent (such as hexyl benzene, chloroform), mixed with the aqueous phase containing emulsifier, and subjected to high-pressure homogenization to form oil-in-water–type nanoemulsions. Afterward, the organic solvent in the internal phase is evaporated by vacuum, and solid lipid nanoparticles with small particle size are solidified [107].

8.3.2.6   Dendrimer nanoparticles Dendrimers are hyperbranched polymers, which form monodisperse spherical molecules by continuously adding repeating units to the central core [111]. A huge number of papers on dendritic architectures, such as dendrimers, dendronized polymers, hyperbranched polymers, and brush polymers, have generated a vast variety of inconsistent terms [16]. The dendritic molecules can be simply divided into the low-­ molecular-weight and the high-molecular-weight species. The first category includes dendrimers and dendrons, whereas the second category encompasses dendronized polymers, hyperbranched polymers, and brush polymers (also called bottle brushes). Dendrimers consist of three parts: a core, an inner shell, and an outer shell. The dendrimer properties are mainly determined by the functional groups on the molecular

236

Biopolymer-Based Composites

surface [112–114]. Poly(amidoamine), or PAMAM, a kind of spherical molecule with a lot of cavity inside and a lot of functional groups outside, is perhaps the most wellknown dendrimer. Dendrimers have a precise molecular structure, and the molecular weight, size, and number of branches on them can be easily controlled. Dendrimers are typically symmetric around the core and available for a large number of functional groups presented on the surface [115]. Because the dendrimers have a variety of terminal groups and diverse chemical structures, they allow the conjugation of a number of drugs, especially, antitumor drugs using various cleavable linkers. The multifunctional dendrimer surfaces carry the antitumor drugs and can also be modified with solubilizing functions and tumor-targeting ligands [108–110]. The first dendrimers were synthesized divergently by Fritz Vögtle in 1978 [116]. Dendrimers can be synthesized by convergent methods or via click chemistry. There are three methods for using dendrimers in tumor drug delivery: first, the covalent conjugate drugs are conjugated to the periphery of the dendrimers to form dendrimer prodrugs; second, the drug is physically encapsulated into the void spaces of the dendrimers to assemble dendrimer–drug complexes; third, the chemical groups of the drugs are coordinated to the outer functional groups via electrostatic interactions.

8.3.2.7  Inorganic nanoparticles Inorganic nanoparticles have good application prospects in the field of reaction catalysis, medicine, electromagnetic function, material modification, and so on [76]. Several typical inorganic nanoparticles, including GNPs and iron oxide nanoparticles, are discussed in the following sections. GNPs refer to gold particles with diameter from 5 to 300 nm and with the shape of sphere, shell, cage, and rod [117–119]. The GNP suspensions are usually presented in different colors from red to blue/purple because of the tunable size and the surface plasmon resonance phenomena caused by absorption of light. For small (∼30 nm) monodisperse GNPs, the absorption of light is located in the blue-green region of the spectrum (∼450 nm) while red light (∼700 nm) is reflected, resulting in solutions with a rich red color. As particle size increases, the wavelength absorption shifts to longer, redder wavelengths, and blue light is reflected, resulting in solutions with a pale blue or purple color [120,121]. Uncoated GNPs are susceptible to aggregation in solution and can melt under laser irradiation. The GNPs are unstable in salt solutions. When excess salt is added to the GNP solution, the surface charge of GNPs becomes neutral, causing nanoparticle aggregation. As a result, the solution color changes from red to blue. GNPs exhibit unique physicochemical properties including their ability to bind amine and thiol groups, allowing easy surface functionalization and bioconjugation [122]. For example, thiolated PEG can bind with GNPs via Au–S bonds, which forms thiolate protecting GNPs, increasing the stability of GNPs, and producing a hydrated PEG barrier to sterically hinder the attachment of phagocytes and reduce reticuloendothelial system uptake [123]. Tumor-specific ligands can be grafted onto GNPs to increase the specificity and likelihood of drug delivery. In the 19th century, Michael Faraday published the first scientific paper on GNP synthesis [124] describing the production of GNP by the reduction of aurochloric acid

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by phosphorous. Owing to their easy synthesis and high stability, various GNPs have been explored in antitumor application. A direct method of accessing and destroying tumor cells is photothermal therapy. These GNPs have high absorption cross sections requiring only a minimal input of irradiation energy. Near infrared (NIR)-absorbing GNPs produce heat when excited by light at various wavelengths. When light is applied to a tumor containing GNPs, the particles rapidly heat up, causing tumor cell ablation [125]. GNPs are also used as carriers for chemotherapy drugs (such as paclitaxel) because of their high drug-loading ability [126]. GNPs can also be used to optimize the biodistribution of drugs to diseased organs, tissues, or cells, to improve tumor drug delivery [127]. Synergistic antitumor effects can be obtained by a combination of photothermal therapy and chemotherapy drugs in GNPs. Iron oxide nanoparticles are iron oxide particles with diameters between 1 and 100 nm. Three most common forms of iron oxides in nature are magnetite (Fe3O4) and its oxidized forms maghemite (γ-Fe2O3) and hematite (α-Fe2O3) [128,129]. Magnetite has an inverse spinel structure, with oxygen forming a face-centered cubic crystal system. There are various shapes of iron oxides, such as nanorods, porous spheres, nanohusks, nanocubes, distorted cubes, and self-oriented flowers. Iron oxide nanoparticles with the size 150 nm are highly susceptible to filtration at interendothelial cell slits of venous sinuses (size 200–250 nm) in the spleen, an area rich in macrophages. Shape also affects the biological responses to nanoparticles. The shape of the particles can be spherical, discoidal, rod-like, or filamentous. Mathematical modeling suggests that the uptake of discoidal particles is more than that of spherical particles regardless of their density and size [174,175]. Specifically, cylindrical nanoparticles of 500 nm to 1 μm were internalized at a rate of 75% in comparison with cubic-shaped 2-μm nanoparticle at a rate of 45%. Rod-shaped and high-aspect-ratio nanoparticles (450 nm long) were even more readily internalized in comparison with more symmetrical and low-aspect-ratio particles, possibly because of the multivalent cationic interactions between the high-aspect-ratio particles and the larger surface area of the cell membrane [5]. The charge of the nanoparticles can impact on their membrane translocation ability, intratumoral processes, and systemic circulation times [176,177]. Positively charged, faceted rice-shaped nanoparticles may experience reorientation driven by electrostatic charges at the vicinity of the cell membrane. Negatively charged nanoparticles are repelled electrostatically by the membrane. A negative charge may increase, decrease, or have no effect on blood clearance, whereas positive charges are believed to have a negative impact [178]. Particle stiffness also plays an important role in the circulation, cellular uptake and tissue distribution of nanoparticles [179–181]. It is reported that soft discoidal polymeric nanoparticles can resist macrophage uptake and enhance vascular targeting in tumors [182]. Huang et al. quantified the relationship between nanoparticle stiffness and cellular uptake by bovine aortic endothelial cells. They found that stiffer nanoparticles resulted in a higher total cellular uptake on a per cell basis but a lower uptake per unit membrane area [179]. Other studies demonstrated that softer nanogels passed through physiological barriers, especially the splenic filtration, more easily than their stiffer counterparts, consequently leading to longer circulation half-life and lower splenic accumulation [180].

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8.3.3.2  Active targeting nanoparticles The theory of active targeting was first proposed by Paul Ehrlich in the 19th century, which he described as the “magic bullet,” an idealized delivery system that would target drugs to specific areas in the body [183]. Active targeting, also called ligand-mediated targeting, involves the use of affinity ligands on the surface of nanoparticles for specific retention and uptake by the targeted disease cells while eliminating off-target adverse effects in normal tissue. Ligands are selected to bind surface molecules or receptors differentially overexpressed in diseased organs, tissues, cells, or subcellular sites [67,184], a process that has considerable implications for targeting delivery. Targeting delivery is ensured by the high specificity of the ligand for its cognate receptor. A wide range of ligands have been used to actively target nanoparticles including antibodies [Herceptin (trastuzumab)] [185], antibody fragments, aptamers (pegaptanib) [186], peptides and proteins (transferrin) [187,188], and small molecules (folic acid) [189]. The first report of targeting nanoparticles dated back to 1980s and involved the surface modification of liposomes with monoclonal antibodies that recognized antigens on the target cells [190,191]. For efficient targeting, nanocarriers should be stable enough to avoid premature release and degradation of the drug in the circulation. The physicochemical properties of nanoparticles must be taken into consideration early in the design of active targeting nanoparticles. Smaller size represents higher curvatures, which can be problematic for ligand functionalization and affect cellular uptake. Additionally, the size can also affect the intracellular deposition of active targeting nanoparticles [192]. The shape of nanoparticles influences the cell uptake kinetics and internalization pathways by modulating the interactions between the nanoparticles and the cell surface [193]. From a synthetic perspective, the charge of the native nanoparticles can affect the conjugation yield of the ligand with nanoparticles and the spatial display of the ligand on the surface [194]. A chemical spacer with reasonable length, such as PEG, can be helpful to reduce the effect but may simultaneously complicate synthesis and increase the final particle size [195]. Besides surface charge, hydrophobicity can also affect the architecture of the ligand display [196], which can have serious effects because most polymeric nanoparticles have hydrophobic cores [197]. Targeting delivery to the tumor cells is the most common strategy for tumor treatment (Fig. 8.3). Utilizing ligand–receptor interactions, facilitated tumor cell uptake of active targeting nanoparticles is achieved [198]. A lot of receptors overexpressed specially on tumor cells, such as folate receptor, transferrin (Tf) receptor, and low-density lipoprotein receptor, can be used for tumor cell drug delivery. Some active targeting nanoparticles based on this strategy have entered clinical trials. The first targeting nanoparticle delivery system to feature siRNA was CALAA-01, which consisted of a cyclodextrin-containing polymer, a PEG corona, and human Tf as a targeting ligand [199]. The Tf on the nanoparticle surface binds to overexpressed Tf receptors on tumor cells, and the nanoparticles are then internalized via receptor-mediated endocytosis. When these siRNA-containing targeted nanoparticles were administered intravenously to patients with melanoma, they circulated in the body and localized in

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Figure 8.3  Active targeting delivery strategies of nanoparticles include tumor cell targeting strategy, tumor stroma cell targeting strategy, tumor extracellular matrix (ECM) targeting strategy, and multitargeting strategy of targeting both tumor cells and different types of tumor stromal cells.

tumors. Tumor biopsies showed a correlation between the dose administered and the amount of intracellularly localized nanoparticles. Furthermore, levels of both the specific mRNA and the protein were lower after injection of the targeted nanoparticles. It is now increasingly recognized that tumor stroma contains distinct cell types, including TAFs, tumor-associated macrophages (TAMs), and neovascular cells, which interact with each other, exert variable roles in enabling tumor progression, and function as the “soil” of tumor growth [200]. Therefore targeting drug delivery to the tumor stroma cells to deplete them is another common strategy for tumor treatment (Fig. 8.3). For example, navitoclax-loaded nanoparticles have been established to selectively eradicate TAFs to produce an effective therapeutic response [201]. Targeting delivery to TAFs or neovascular cells by active targeting nanoparticles can significantly increase the accumulation of drugs in TAMs or neovascular cells, depletion of TAMs and neovascular cells, resulting in improved tumor therapy [202]. The approaches targeting tumor stroma cells deliver therapeutics only to a limited number of cell types in tumors and cannot eradicate the tumor “soil” totally, and multitargeting strategy of targeting both tumor cells and different types of tumor stromal cells and destroying tumor “seeds” and “soil” simultaneously may be a more promising approach for tumor treatment (Fig. 8.3). Tissue factor (TF) was shown to be extensively located in a tumor and was abundantly sited in both tumor cells and stromal cells, including neovascular cells, TAFs, and TAMs, indicating it may function as a favorable target for drug delivery to multiple cell types simultaneously [200]. Using enhanced green fluorescent protein-epidermal growth factor-like domain-1

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(EGFP-EGF), a fusion protein derived from factor VII, the natural ligand of TF as the targeting moiety, a multitargeting nanoparticle drug delivery system was established and offered promising therapeutic outcomes. However, both targeting moieties ­capable of targeting multiple cell types in tumors simultaneously and the targetable receptors widely and homogeneously distributed in tumors are scarce. The abundant ECM is a dynamic and hierarchically organized nanocomposite that not only provides mechanical support for embedded tumor cells, but also interacts with tumor cells and promotes tumor cell migration, proliferation, and differentiation [203]. Targeting tumor ECM offers an optional strategy of active targeting nanoparticles for tumor treatment (Fig. 8.3) [204,205]. Active targeting nanoparticles have yielded promising findings in preclinical studies and, in some cases, early clinical trials of tumor therapy. However, some studies involving active targeting nanoparticles have been inconclusive and therapeutic efficacy in humans has not yet been convincingly demonstrated overall. The interactions of the protein corona around a nanoparticle with physiological proteins in the body, as well as factors that interrupt the orientation and proper display of the targeting ligands, highlight the need for further studies of the clinical relevance of actively targeting nanoparticles.

8.3.3.3  Smart nanoparticles for tumor targeting Smart nanoparticles mainly refer to nanosized particles working in stimuli-responsive systems in response to physical, chemical, or biological triggers that promote the release of drugs by interfering with the phase, structure, or conformation of the nanoparticles [206,207]. In this stimuli-triggered strategy, nanoparticles are passively delivered and accumulated in tumor tissues via the EPR effect at first. When the nanoparticles reach the target site, the nanoparticles are then activated and release the payload [208]. The advantage of smart nanoparticles is obvious: the drug is released through a trigger presented in the specific tissue region, thus minimizing systemic exposure to the drugs. Triggers can be divided into internal and external stimuli (Table 8.2) [18]. Internal stimuli are the pathophysiological/pathochemical conditions that are differentially presented in normal tissues and tumor tissues, including changes in hypoxia, pH, redox, ATP, enzyme expression, ionic strength, and shear stress [207,209]. Low blood perfusion and dense tumor growth lead to hypoxia at the tumor site. Nitroaromatic derivatives that can be converted to hydrophilic 2-aminoimidazoles under hypoxic conditions with a relatively high sensitivity are among the most widely exploited functional motifs for hypoxia imaging and the design of bioreductive prodrugs [210,211]. The pH within the tumor microenvironment is much more acidic (pH 6.8–7.2), especially their endosomes and lysosomes (pH 5.0–5.5), than blood (pH 7.35–7.45) and other normal tissues [233]. Smart nanoparticles that are destabilized in a slightly acidic environment [213] release the loaded drug into the extracellular medium triggered by tumor microenvironment, acting as tumor targeting [208]. The hypoxic area of tumors exhibits low oxygen pressure and poor nutrient levels [234]. Differences in redox potential exist at both the tissue and cellular level.

Common smart nanoparticles used in targeting drug delivery to tumors

Stimuli

Principles

246

Table 8.2 

Examples of nanoparticles

Function

Citations

Nitroimidazole-conjugated polymer nanoparticles

Accelerating drug release

[210–212]

PEGMnCaP nanoparticles, DOXloaded calcium carbonate hybrid nanoparticles, zwitterionic polymer nanoparticles Nanoparticles containing disulfide linkage

Accelerating drug release or increasing tumor cell uptake of nanoparticles

[213–215]

Accelerating drug release or increasing tumor cell uptake of nanoparticles Accelerating drug release

[216,217]

Accelerating drug release

[220,221]

Improving tumor penetration or tumor cell uptake of nanoparticles

[222,223]

Internal stimuli Hypoxia

pH

Redox

Normal critical oxygen partial pressure is 8–10 mmHg on a global tissue level; solid tumor regions with low oxygen partial pressures (down to zero) Normal pH range: 7.38–7.42; tumor extracellular pH 7.2–6.5

Glutathione: intracellular, 10 mM; extracellular fluids, 2–10 μM Inflammation and injury tissue: oxygen species levels reactively elevated

Enzyme

Intracellular environment: 1–10 mM; extracellular environment: