Tunable Hydrogels: Smart Materials for Biomedical Applications (Advances in Biochemical Engineering/Biotechnology, 178) 303076768X, 9783030767686

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Advances in Biochemical Engineering/Biotechnology  178 Series Editor: T. Scheper

Antonina Lavrentieva Iliyana Pepelanova Dror Seliktar   Editors

Tunable Hydrogels Smart Materials for Biomedical Applications

178 Advances in Biochemical Engineering/Biotechnology Series Editor Thomas Scheper, Hannover, Germany Editorial Board Members Shimshon Belkin, Jerusalem, Israel Thomas Bley, Dresden, Germany Jörg Bohlmann, Vancouver, Canada Man Bock Gu, Seoul, Korea (Republic of) Wei Shou Hu, Minneapolis, MN, USA Bo Mattiasson, Lund, Sweden Lisbeth Olsson, Göteborg, Sweden Harald Seitz, Potsdam, Germany Ana Catarina Silva, Porto, Portugal Roland Ulber, Kaiserslautern, Germany An-Ping Zeng, Hamburg, Germany Jian-Jiang Zhong, Shanghai, Minhang, China Weichang Zhou, Shanghai, China

Aims and Scope This book series reviews current trends in modern biotechnology and biochemical engineering. Its aim is to cover all aspects of these interdisciplinary disciplines, where knowledge, methods and expertise are required from chemistry, biochemistry, microbiology, molecular biology, chemical engineering and computer science. Volumes are organized topically and provide a comprehensive discussion of developments in the field over the past 3–5 years. The series also discusses new discoveries and applications. Special volumes are dedicated to selected topics which focus on new biotechnological products and new processes for their synthesis and purification. In general, volumes are edited by well-known guest editors. The series editor and publisher will, however, always be pleased to receive suggestions and supplementary information. Manuscripts are accepted in English. In references, Advances in Biochemical Engineering/Biotechnology is abbreviated as Adv. Biochem. Engin./Biotechnol. and cited as a journal.

More information about this series at http://www.springer.com/series/10

Antonina Lavrentieva • Iliyana Pepelanova • Dror Seliktar Editors

Tunable Hydrogels Smart Materials for Biomedical Applications

With contributions by V. Korzhikov-Vlakh
A. Lavrentieva
H. Mohsenin
J. Nie
I. Pepelanova
J. Ren
J. J. Senior
A. M. Smith
F. Sun
T. Tennikova
H. J. Wagner
J.-G. Walter
W. Wang
W. Weber
Z. Yang
A.-P. Zeng
X. Zhang

Editors Antonina Lavrentieva Institute of Technical Chemistry Leibniz University of Hannover Hannover, Germany

Iliyana Pepelanova Institute of Technical Chemistry Leibniz University of Hannover Hannover, Germany

Dror Seliktar Faculty of Biomedical Engineering Technion-Israel Institute of Technology Haifa, Israel

ISSN 0724-6145 ISSN 1616-8542 (electronic) Advances in Biochemical Engineering/Biotechnology ISBN 978-3-030-76768-6 ISBN 978-3-030-76769-3 (eBook) https://doi.org/10.1007/978-3-030-76769-3 © Springer Nature Switzerland AG 2021 This work is subject to copyright. All rights are reserved by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG. The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Preface

A book volume is a cooperative project, which would not be realized without the support and contribution of many people. First, we would like to thank Prof. Dr. Thomas Scheper, the series editor of “Advances in Biochemical Engineering/Biotechnology,” for inviting us to serve as editors of this volume. We would also like to express our gratitude to Sofia Costa and Alamelu Damodharan from Springer Nature Publishing House for their patience, support, and professional guidance. Together with this dedicated team, we were able to complete the following book volume, which explores a versatile class of modern biomaterials: “Tunable Hydrogels: Smart Materials for Biomedical Applications.” Life on our planet would not be possible without water. For this reason, hydrogels, which are polymer networks retaining large amounts of water, represent a natural material of choice in basic science research and applied healthcare. Tunability is defined as the implementation of knowledge-based tools to manipulate material properties in the direction desired for the specific application. Hydrogels offer versatile properties and the possibility to modify the polymer backbone using the toolkits of chemistry and biotechnology results in specialized materials which can be engineered to fulfill almost any imaginable purpose. Moreover, hydrogels can be designed with stimulus-responsive elements, which allow the creation of smart materials, which display adaptive and dynamic responses to environmental or user cues. Tunable and smart hydrogels are currently explored in many venues including drug and cell delivery, 3D cell culture, tissue engineering, biosensors, wound dressings, and soft actuators. This book starts with an overview chapter, which leads the reader to a general introduction to concepts of hydrogel tunability and introduces the mechanisms behind the design of stimulus responsiveness relevant for biomedical applications. In the next chapter, the principles of mechanical and biological tunability are illustrated in-depth on the example of the widely used alginate. Biodegradability is a prized asset of a hydrogel system, especially for materials that are to be implanted or injected in the human body. For this reason, protein hydrogels are an especially attractive platform for biomedical applications. Various protein hydrogels, including v

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those that display thermo and pH-responsiveness are treated in the third chapter “Tunable Protein Hydrogels.” The application of tunable hydrogels as highly specific release systems on the nano-scale is described in the chapter “Nanogels Capable of Triggered Release.” Achieving even more specific responsiveness to biomolecules is possible by using recognition elements like aptamers. The design of aptamer-modified hydrogels is explored in the next chapter. While many of the stimulus-responsive elements are based on chemical principles, nature has also developed a large array of receptors which demonstrate stimulus responsiveness as well. Using genetic engineering, it is possible to design novel stimulus-responsive hydrogels (chapter “Self-Assembly and Genetically Engineered Hydrogels”). The tools of synthetic biology for the creation of intelligent biosensors are presented in “Synthetic Biologically-Empowered Hydrogels for Medical Diagnostics”. Finally, hydrogels are also a promising platform for 3D cell culture and tissue engineering. Tunable hydrogels allow the creation not only of bulk constructs with uniform properties, but also provide researchers with materials which can be designed to exhibit mechanical and biochemical gradients. Readers will find descriptions of fabrication platforms and material selection for this purpose in the last chapter of this book volume “Gradient Hydrogels.” The reader familiar with the fundamentals of biomaterial science will discover the latest frontiers of hydrogel research, while researchers new to the field will be introduced to basic concepts of hydrogel tunability for the most common applications. We are, therefore, convinced that the book will be a valuable contribution to all researchers and may inspire them to pursue some of the open questions and paradigms outlined in these pages. Last but not least, we are grateful to all contributing authors for committing to this book project and for sharing their extensive knowledge in their area of expertise. The Corona pandemic has changed our daily and professional routines, but has not dampened existing scientific enthusiasm. Indeed, the fast development of the Covid-19 vaccines show how knowledge-based tools and scientific cooperation allow researchers to overcome any existing and unexpected global healthcare challenges. Hannover, Germany Hannover, Germany Haifa, Israel March 2021

Antonina Lavrentieva Iliyana Pepelanova Dror Seliktar

Contents

Tunable Hydrogels: Introduction to the World of Smart Materials for Biomedical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Iliyana Pepelanova Alginate Hydrogels with Tuneable Properties . . . . . . . . . . . . . . . . . . . . . Alan M. Smith and Jessica J. Senior Tunable Protein Hydrogels: Present State and Emerging Development . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . J. Nie, X. Zhang, W. Wang, J. Ren, and A.-P. Zeng Nanogels Capable of Triggered Release . . . . . . . . . . . . . . . . . . . . . . . . . Viktor Korzhikov-Vlakh and Tatiana Tennikova

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Aptamer-Modified Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 147 Johanna-Gabriela Walter Self-Assembly and Genetically Engineered Hydrogels . . . . . . . . . . . . . . 169 Zhongguang Yang and Fei Sun Synthetic Biology-Empowered Hydrogels for Medical Diagnostics . . . . . 197 Hanna J. Wagner, Hasti Mohsenin, and Wilfried Weber Gradient Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 227 Antonina Lavrentieva

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Adv Biochem Eng Biotechnol (2021) 178: 1–36 https://doi.org/10.1007/10_2021_168 © Springer Nature Switzerland AG 2021 Published online: 27 April 2021

Tunable Hydrogels: Introduction to the World of Smart Materials for Biomedical Applications Iliyana Pepelanova Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 General Tunability of the Bulk Material . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Control of Mechanical Properties: Stiffness and Mesh Size . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Control of Biological Functionality: With a Focus on Degradation and Adhesion . . 3 Tunable Hydrogels as Smart Materials and Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Thermo-Responsive Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 pH-Responsive Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Photo-Responsive Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Biomolecule-Responsive Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5 Electromagnetic-Responsive Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6 Other Responsive Hydrogel Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.7 Self-Healing Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract Hydrogels are hydrated polymers that are able to mimic many of the properties of living tissues. For this reason, they have become a popular choice of biomaterial in many biomedical applications including tissue engineering, drug delivery, and biosensing. The physical and biological requirements placed on hydrogels in these contexts are numerous and require a tunable material, which can be adapted to meet these demands. Tunability is defined as the use of knowledge-based tools to manipulate material properties in the desired direction. Engineering of suitable mechanical properties and integrating bioactivity are two major aspects of modern hydrogel design. Beyond these basic features, hydrogels can be tuned to respond to specific environmental cues and external stimuli, which I. Pepelanova (*) Institute of Technical Chemistry, Leibniz University of Hannover, Hanover, Germany e-mail: [email protected]

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are provided by surrounding cells or by the end user (patient, clinician, or researcher). This turns tunable hydrogels into stimulus-responsive smart materials, which are able to display adaptable and dynamic properties. In this book chapter, we will first shortly cover the foundation of hydrogel tunability, related to mechanical properties and biological functionality. Then, we will move on to stimulusresponsive hydrogel systems and describe their basic design, as well as give examples of their application in diverse biomedical fields. As both the understanding of underlying biological mechanisms and our engineering capacity mature, even more sophisticated tunable hydrogels addressing specific therapeutic goals will be developed. Graphical Abstract

Keywords Biomedical applications, Biosensors, Drug delivery, Hydrogel, Smart, Soft actuators, Stimulus-responsive, Tissue engineering, Tunable

1 Introduction The biomedical field brings together many experts from diverse disciplines such as clinicians, materials scientists, biologists, and chemists, all working together to understand the intricate functions of human physiology, so as to be able to cure diseases, as well as improve the quality of human life. These fruitful collaborations and the human drive to curiosity and inventiveness have brought forth many biomaterials suitable for application in combination with (seeded) human cells or directly within the human body. In the last 20 years, hydrogels have turned out to be one of the most promising categories of biomaterials due to their unique properties and high biocompatibility [1]. Hydrogels are made up of polymer chains, which are extensively hydrated. In this way, they are able to effectively mimic the extracellular matrix (ECM). The high water content of hydrogels enables them to support

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life, the polymer networks impart the biomaterial with tissue-like elasticity and the porous, open structure allows the diffusion of nutrients and waste substances. Hydrogels undergo a sol-gel transition from a liquid precursor to a gel-like crosslinked network, typically under mild, cytocompatible conditions. The abovementioned properties allow the application of hydrogels for direct cell encapsulation, a property relevant for biomedical areas like tissue engineering, cell-based therapies, cell biology studies, and in vitro 3D modeling [2]. Moreover, the polymer backbone can be modified with different functional groups, allowing design versatility, control of properties, and even the introduction of in-built responsive elements. This makes hydrogels smart materials promising for areas like microfluidics or biosensing. The controlled porosity of the polymer network allows hydrogels to function as a carrier and reservoir of active substances, making them attractive materials for areas like drug delivery. In addition, their ability to store water while displaying gas permeability makes hydrogels superb wound dressings, which find widespread clinical use [3]. In this book chapter, we will explore the concept of tunability of hydrogel systems. In a broad and general sense, “tunability” can be defined as knowledgebased tools used to manipulate hydrogel properties in the direction of the desired outcome. First, we will shortly cover the basics of hydrogel science by venturing into a short discussion of the general tunability of bulk hydrogel systems, focusing on the control of mechanical properties and the introduction of desired biological functionality. Much of this has already been extensively described in the literature, which is why we will just focus on important basic concepts. Afterwards, we will turn our attention to the design of smart hydrogel materials with in-built stimulus responsive elements, and we will discuss in detail how such smart hydrogels are used in biomedical applications. The book chapter will conclude with an outlook of current research frontiers and existing challenges in the field. The use of fabrication techniques for building hydrogels into specific architectures like electrospinning, stereolithography, microfluidics, and bioprinting will not be covered in this chapter, as it is beyond the scope of this book volume. Readers wanting to learn more about existing and emerging technologies for engineering the shape of hydrogel constructs are referred to these comprehensive reviews [4–7].

2 General Tunability of the Bulk Material Hydrogels can be derived from natural biopolymers or from synthetic polymers. Natural hydrogels like collagen, fibrin, or hyaluronic acid are very biocompatible and almost a straightforward choice when desiring to mimic natural tissues. However, due to their sourcing from biological resources, they typically display an undefined structure and batch-to-batch variation in quality. In addition, they frequently suffer from poor mechanical properties. Synthetic hydrogels, on the other hand, like poly(ethylene glycol) (PEG), poly(vinyl alcohol) (PVA), or poly(ε-caprolactone) (PCL) can be created with defined composition and, thus, offer

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control over the resulting mechanical properties. However, the inert structure of synthetic polymers is not cell promoting, it does not provide cells with biological information and cannot instruct cellular physiology. Therefore, researchers typically seek to combine the best of both worlds and use chemical tools to create semisynthetic hydrogels, hydrogels with a synthetic structure for control of stability, and a biological moiety for physiological functionality [8]. Tunability begins with an understanding of the molecular composition of a hydrogel and the crosslinking mechanisms leading to sol-gel formation, for each specific hydrogel system. For a more in-depth discussion of classifications of hydrogels according to their nature of origin (natural, synthetic, and semi-synthetic), as well as the chemical structure of the most widely used hydrogel systems, the reader is referred to these excellent reviews [9–11]. Hydrogel crosslinking mechanisms including physical and ionic bonding, as well as chemical crosslinking, including polymerization mechanisms and click chemistry are treated in-depth here [12–14]. In the current book volume, the tunability of a hydrogel system is explored in great depth and detail on the example of alginate. The authors of the book chapter “Alginate hydrogels and their tunable properties” give a comprehensive analysis of molecular alginate structure and the mechanisms of crosslinking, and how these can be manipulated to alter alginate properties. In addition, the authors discuss how alginate can be further fine-tuned for control of relevant biological parameters like degradation and cell adhesion. In this book chapter, we will turn our attention next to general concepts of hydrogel tunability, applicable in a general sense to all hydrogel systems.

2.1

Control of Mechanical Properties: Stiffness and Mesh Size

Controlling hydrogel stiffness is important in many biomedical applications, as it is a major factor affecting material stability, whether in biosensing or in tissue engineering. Moreover, the mechanical properties of a hydrogel influence many aspects of cellular biology including cell differentiation, migration, and proliferation, in a process called mechanotransduction. As a result, researchers seek to tune hydrogel mechanical properties in the direction desired for their specific application. Apart from chemical composition and mechanism of network formation, the resulting crosslink density and the effective mesh size of the polymer network (see Fig. 1) will influence hydrogel mechanical properties. The crosslink density (ρx) refers to the number of polymer chains in a given volume of material and this property influences in turn the gel shear modulus G (and thus the mechanical properties) and the resulting mesh or pore size of the polymer network (which determines the transport of solutes and biomolecules throughout the network by diffusion). The crosslink density furthermore affects the water-holding

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Fig. 1 Tunability of basic hydrogel properties can be achieved by control of crosslink density, with higher crosslink density ρx related to stronger gel mechanics and shear modulus (G), and smaller mesh size ζ, which in turn affects the swelling ratio (Q) and the diffusivity throughout the hydrogel (D). Inspired from DeForest and Anseth [15]

capacity of the hydrogel, called the mass swelling ratio Qm. The mass swelling ratio is experimentally determined by incubating the crosslinked hydrogel construct to equilibrium (typically 24 h at 37 C in aqueous buffer or media), blotting excess liquid and weighing the swollen gel (Mw). The hydrogel is then dried (freeze-dried) and weighed again to give us its dry weight Md. The ratio between Mw/Md is defined as the mass swelling ratio Qm. The swelling ratio is related to the crosslink density, with more densely crosslinked polymer networks displaying lower swelling capacity. The mesh size is also related to ρx, with stiffer hydrogels having typically higher crosslink density and smaller pores [16]. Mechanical tunability of a given hydrogel system thus starts from a given chemistry of the polymeric backbone and the mechanism used for crosslinking and extends to conditions which influence crosslink density. To illustrate this with an example, gelatin methacryloyl (GelMA) hydrogels are semi-synthetic hydrogels produced by methacryloyl substitution of amine and hydroxyl amino acid residues on the protein backbone of gelatin. These allow the formation of short polymethacryloyl chains between individual gelatin molecules in the presence of light, in a free-radical photo polymerization process (see Fig. 2). The synthetic chains form covalent crosslinks, which stabilize gelatin at a physiological temperature of 37 C [17]. The mechanical properties of GelMA can be controlled by modifying the derivatization reaction to obtain the desired degree of functionalization (DoF), with higher DoF leading to stiffer gels and higher (ρx), due to the presence of more reactive side groups for crosslinking [18]. The mechanical properties of a GelMA of a given DoF can be further fine-tuned by changing the polymerization conditions, such as variation in light dosage or photoinitiator (PI) concentration. Higher PI and/or light dosage leads to a stiffer hydrogel, with higher crosslinking density and resulting smaller pores and lower swelling ratio. Thus, increasing the concentration of a hydrogel or manipulating the crosslink density are ways to enhance the mechanical properties of a given hydrogel system.

Fig. 2 (a) GelMA is a tunable semi-synthetic hydrogel, created by methacryloyl substitution of reactive amine and hydroxyl groups on amino acid residues of gelatin, resulting in a material of specific degree of functionalization (DoF) (b) The methacryloyl groups react together in the presence of an initiator and light; depending on the DoF a material of different crosslink density is obtained (c) The different ρx of GelMA translates into variable mechanical strength, here demonstrated for a 5% w/v solution of different DoFs measured by oscillatory rheology (d) Encapsulated cells react to mesh size and crosslink density by displaying spindle-shaped morphology in low DoF materials and remaining round in high DoF materials, Calcein-AM staining, scale bar 1,000 μm

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Fig. 3 Strategies to enhance the mechanical properties of hydrogels include increasing the precursor concentration or enhancing the crosslink density. Further increase in mechanical properties is achieved by the addition of a second polymer by blending (resulting in IPNs or semi-IPNs), co-polymerization or by the addition of nanoparticles

Alternatively, the macromer precursor can be substituted with a material of the same type but of higher molecular weight (for example, by taking a gelatin precursor with a higher Bloom factor (a test related to the gel strength and molecular mass of gelatin) in the case of GelMA). If these interventions do not provide the required mechanical properties, researchers often use further enhancements by using blending, co-polymerization with additional polymers or even nanoparticles to enhance the mechanical properties of the system (see Fig. 3). In the first approach, two or more polymers form an interpenetrating network (IPN), with the entanglement of polymer chains leading to reinforcement of the mechanical properties of the hydrogel network as a whole [19]. It is important to note that in IPN, there is no covalent bonding between the two individual polymers; the increase in stability is due to molecular entanglement. IPNs can be created between natural polymers, synthetic polymers, or a mixture of both. Frequently, a synthetic polymer is combined with a natural one to obtain stability and biological features at the same time. If a linear polymer is blended with another crosslinked network, then the resulting construct is called semi-IPN [20]. This approach has been widely used to boost the mechanical properties of collagen hydrogels to ranges appropriate for tissue engineering applications of elastic and hard tissues. For instance, Munoz-Pinto et al. developed a collagen-poly(ethylene glycol) diacrlylate (PEGDA) IPNs, by dispersing PEGDA in crosslinked collagen gels and then illuminating the constructs with light after 6 h delay for PEGDA polymerization. The resulting IPN displayed improved stiffness, strength, and physical stability, while retaining the ability to support cell spreading [21]. Double network (DN) hydrogels are a special category of IPNs. Here, one polymer provides a rigid “skeleton” structure and the other a ductile, soft secondary

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network. The interpenetration of the macromolecular chains of the “soft-tough” polymer pair results in a non-linear increase of mechanical strength and resistance to fracture, which is able to mimic the properties of load-bearing tissues like muscle and bone [22]. For example, DN gels composed of poly(2-acrylamido-2-methyl propanesulfonic acid) (PAMPS) as the first network and PAAm as the second network [23], have been demonstrated as effective therapies in animal in vivo models of cartilage regeneration [24]. A co-polymerization strategy can also be used to enhance the mechanical properties of hydrogels. Here, the hydrogel is polymerized from two or more different precursor species (monomers or macromers), from which at least one is highly hydrophilic in nature. The individual monomeric constituents can be assembled in a random, alternating, or block fashion and result in a copolymer of improved mechanical stiffness. For example, Buwalda et al created temperature-responsive enantiomeric PEG-polylactic acid (PLA) star block copolymers, which can be used as a drug delivery system. The single enantiomeric hydrogels showed poorer stability and mechanical properties compared to the enantiomeric copolymer at 37 C in phosphate-buffered saline (PBS) [25]. Co-polymerization with another methacrylated macromer is also frequently used to enhance the mechanical properties of photoactive semi-synthetic hydrogels like GelMA. This has been demonstrated for dextran glycidyl methacrylate (DexMA) [26], methacrylated hyaluronic acid (HAMA) [27], and methacrylated oxidized alginate [28] among others. In addition to the above-mentioned strategies, which use bulk polymers for enhancing the mechanical properties of hydrogels, it is also possible to reinforce stability by the addition of nanoparticles. The dispersion of nanoparticles (NPs) in a hydrogel results in a nanocomposite material, which displays properties different from the separate materials themselves. The NPs occupy spaces between the hydrogel chains and can associate with them via chemical or physical bonding. The formation of NP bridges thus reinforces the mechanical strength of the nanocomposite hydrogel. Inorganic materials/clays, carbon-based materials, metals/metal oxides, and polymers have all been used as NPs in nanocomposite hydrogels [29]. In addition to improving the mechanical properties, the NPs imbue the hydrogel with supplementary functionality, which can be tailored to the specific application. For example, inorganic NPs like nano-hydroxyapatite (nHA) are popular in bone tissue engineering applications, as they not only increase the tensile and compressive strength of the hydrogel, but also serve as a mineral reservoir for osteoregeneration [30]. Carbon-based NPs like carbon nanotubes (CNTs) and graphene oxide (GO) reinforce the hydrogel and introduce thermal and electrical conductivity. This makes them potential material candidates for the reconstruction of electrically conductive tissues like cardiac tissue [31], muscles [32], and nerves [33, 34]. Metallic NPs are also frequently used for the creation of nanocomposite hydrogels. To illustrate some examples: silver NPs contribute to antimicrobial activity in hydrogels used as wound dressings [35], while iron/iron oxides NPs impart the composite with magnetic functionality which can be used in biosensing, diagnostics, and drug release [36].

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The reader should not confuse nanocomposite hydrogels with nanogels. The latter term indicates using hydrogels to create nanoparticles. Indeed, there are nanocomposites of nanogel as NP component in a bulk hydrogel. Nanogels are very promising materials for drug/gene delivery applications. They can be made from various stimulus-responsive hydrogels and be designed to unload their cargo of bioactive substances at specific stimuli including temperature, pH, or an external electromagnetic field. The smaller the nanogel, the faster its response, demonstrating the unique advantages of smart hydrogel materials on the nanoscale [37]. The reader who is curious to learn more about nanogels and their special properties, manufacture and applications is referred to the chapter “Nanogels capable of triggered release” in this book volume. The importance of controlling stiffness for biomedical applications directly involving cells cannot be overstated. Cells sense the stiffness of the underlying or surrounding matrix via integrins and membrane-associated proteins [38]. The effects of hydrogel mechanical properties on cell behaviors like cell adhesion, spreading, growth, and differentiation have been thoroughly studied in both 2D and 3D cell culture. For example, it has been shown that cells migrate and proliferate extensively when grown in 2D on hard hydrogel surfaces [39, 40], but may be hindered in adhesion, migration, and proliferation, when 3D encapsulated in dense networks of high crosslink density [41, 42]. Moreover, it should be considered that human tissues and organs including skin, bone, and blood vessels are not uniform in mechanical properties, but exhibit directional stiffness gradients. These gradients regulate cell behaviors, including cell death and differentiation, which lead to the generation and maintenance of complex, anisotropic tissue structures. Indeed, hydrogels are the materials of choice used to model the existence of stiffness gradients, as well as other gradient types (oxygen, biomolecules, etc.), in cell culture. Interested readers are referred to the chapter “Gradient hydrogels” in this book, which describes in detail the rationale behind working with gradient substrates, as well as methods for their manufacture and analysis.

2.2

Control of Biological Functionality: With a Focus on Degradation and Adhesion

The tuning of hydrogel degradation is strongly dependent on the goal of the experiment. In tissue engineering, for example, it is often desirable that the hydrogel matrix supports cells as long as they can gradually produce their own ECM. In drug releasing applications, the degradation of the hydrogel is a factor controlling the release kinetics of the loaded bioactive substances. The rate of hydrogel degradation is related to its mechanical properties – with higher polymer concentration, higher molecular weight (longer chains) derivatives or a higher crosslink density usually leading to slower degradation. Degradation proceeds through the hydrolytic cleavage of bonds or through the action of enzymes [43].

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Fig. 4 Hydrogel degradation proceeds through hydrolysis or via enzymatic attack. It might be necessary to modify the hydrogel with such degradation sites to control the rate of polymer degradation. The graph shows a typical experimental degradation profile of a hydrogel, with sudden mass loss at the point of reverse gelation

This means that the degradation reaction itself gradually transforms the hydrogel’s mechanical properties, as cleaved bonds lead to lower crosslink density, bigger mesh size, and increased swelling. As the degradation proceeds, so many bonds within the crosslinked hydrogel are broken, that the 3D network becomes loose and gradually disintegrates, dissolving in the medium. This process is called “reverse gelation” and is accompanied experimentally by a sharp and sudden mass loss from the hydrogel (see Fig. 4). Protein hydrogels like collagen, elastin, or fibrin are derived from the ECM and contain inherent degradation sequences recognized by matrix metalloproteases (MMPs), secreted by cells. Other natural hydrogels like HA can also be degraded by cells with hyaluronidase enzymes. Hydrogels with ester linkages degrade with various rates as soon as these are exposed to aqueous environments. The degradation of synthetic polymers like PEG, without hydrolytically labile bonds, does not occur on time-scales relevant for most biological experiments, but can be tuned by two strategies. In the first approach, the hydrogel is turned into a composite with another polymer, which possesses hydrolytically labile bonds. For example, PEG can be rendered biodegradable by co-polymerization with PLA, as in the block copolymers of PEG-b-PLA [44]. Even natural hydrogels like HA have been engineered with bonds, which are susceptible to hydrolysis, in order to further fine-tune their degradation via a dual hyaluronidase and hydrolytic pathway [45]. In the second approach, synthetic polymers are modified using artificial peptide linkers, bearing the recognition sequences for MMPs or other physiological enzymes like plasmin or elastase, which render the synthetic hydrogel enzymatically degradable [46]. Such peptide linkers are synthesized with the functional groups of the crosslinking reaction and can be added to the hydrogel precursor mix prior crosslinking, thus being integrated throughout the hydrogel network during the sol-gel transition.

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The protease-mediated degradation of hydrogels is an interesting application for site-specific drug delivery. Some pathological conditions like cancer or arthritis are associated with higher tissue levels of proteases. For example, Tauro and Gemeinhart engineered PEGDA hydrogels with MMP linkages to achieve targeted delivery of a chemotherapeutic agent. The model system was used effectively in glioblastoma multiforme, which displays elevated levels of MMP-2 and MMP-9, associated with its invasive and angiogenic nature [47]. Since the degradation of hydrogels proceeds through an enzymatic or hydrolytic cleavage of bonds, it is possible to fine-tune hydrogel degradation even further by controlling the cleavage site. For instance, by modifying the amino acid sequence of the peptide linker it is possible to exploit different affinities of an enzyme for a cleavage site [48]. In another example of tuning enzyme-mediated hydrolysis, Feng et al. sulfated HA leading to slower enzymatic degradation by hyaluronidase, which strongly promoted chondrogenesis and improved the stability of the hydrogel in vivo [49]. In the case of ester bond hydrolysis, it is possible to slow down degradation by building caprolactone esters, which degrade slower than lactic acid esters, for example [50]. Besides the type or length of the hydrolytically labile linker, it is important to consider the environment of the hydrogel, as the presence of cells, the media used or true in vivo conditions will all influence the rate of degradation. Some biomedical applications require cell adhesion to the hydrogel network, if anchorage-dependent cells are to be grown or studied. ECM-derived protein hydrogels like collagen or GelMA offer natural integrin-binding sequences mediating cell adhesion. Synthetic polymers or polysaccharide-based hydrogels do not contain these sequences, with encapsulated cells remaining in round morphology [10]. Such non cell-promoting hydrogels can be modified with arginine-glycineaspartate (RGD) peptides or other integrin-binding sequences for achieving cell adhesion. Functionalization with RGD peptides to create bioactive hydrogels has been demonstrated for a diverse class of hydrogel materials including alginate [51], HA [52], chitosan [53], and PEG [54]. The resulting materials display superior biological properties over their unmodified counterparts, significantly affecting cell adhesion, spreading, proliferation, and differentiation (see Fig. 5). While adhesion is one of the most important biological functionalities to consider when engineering a hydrogel for anchorage-dependent cell growth, the simulation of other biological interactions may be just as important for the specific biomedical application. In the area of tissue engineering, it is often of interest to load growth factors into a hydrogel, instead of adding these to the cell culture medium. In this way, it is possible to use lower effective concentration of growth factors and to administer them in a targeted and local manner to the cells that require them. Moreover, the physiological ECM also contains sequestered growth factors, which exhibit prolonged action and are protected from degradation in this way [55]. Growth factors can be encapsulated into a hydrogel network, in which case mesh size and hydrogel degradation will control their rate of release [56]. An elegant approach includes modifying hydrogels with heparin, which binds many growth factors through non-covalent affinity interactions [57]. In order to slow down the rate of

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Fig. 5 Most hydrogels are non-cell-promoting, representing a blank slate for cells. In order to engineer cell-permissive properties for specific applications like tissue engineering, it is necessary to add sites enhancing the biological functionality and instructing cells to adhere, spread, and migrate throughout the hydrogel network

release of growth factors even further and thus decrease their effective concentration, it is also possible to directly use covalent linkages for anchoring growth factors to hydrogels. This is especially relevant for the delivery of vascular endothelial growth factor (VEGF), which displays adverse effects when administered in high doses. Zisch et al. demonstrated a highly effective VEGF-modified hydrogel matrix, which was remodeled into vascularized tissue in rat experiments. VEGF was connected to the PEG network via Michael-type addition and released by cells on demand through remodeling of the PEG hydrogel, which was additionally engineered with RGD and MMP-sensitive peptides [58]. When a covalent modification for growth factor immobilization is chosen, care must be taken that the chosen reaction does not diminish the bioactivity of the growth factor. Building even more specific biological functionality is especially relevant in the area of biosensing, drug delivery, or bioimaging. Here the hydrogel may be engineered with a specific recognition structure like an aptamer, DNA sequence, enzyme, etc., which binds specifically to the target of interest. More information about biomolecule-responsive hydrogel systems is presented in this book chapter in Sect. 3.4.

3 Tunable Hydrogels as Smart Materials and Applications So far, we have considered the general principles in understanding hydrogel tunability, especially in the general context how to modify mechanical properties and biological functionality. Most bulk hydrogels display a static response to their environment, for example swelling or shrinking in relation to the water availability in its milieu. Living organisms, on the other hand, have evolved mechanisms to react or even adapt to environmental changes and their highly dynamic surroundings. Organisms are able to do so by their structural components, which recognize and show specific responses to environmental stimuli. It is therefore of great interest, to design biomimetic materials which are able to elicit some form of tailored or

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“intelligent” response to external stimuli. A smart hydrogel exhibits a change in property, most usually swelling behavior, but it can also be a change in stiffness, degradation, sol-gel transition, permeability, etc. in response to an external stimulus. The type of stimuli, which are most relevant for the biomedical context, include parameters like pH and temperature but also specific biomolecules, which can serve as recognition events (see Fig. 6). Moreover, physical cues like light, electric and magnetic fields can also be very interesting to incorporate in hydrogel tunability, as they can be easily applied externally to a construct. Smart hydrogels have therefore been investigated in a very wide-ranging context including (1) biosensors, (2) drug delivery and controlled release, (3) tissue engineering, and (4) soft actuators, which convert a molecular reaction into motion (e.g., smart valves in microfluidic devices). In the next part of this manuscript, we will explore the most frequently used types of stimulus-responsive hydrogels and give examples of their applications.

3.1

Thermo-Responsive Hydrogels

Thermo-responsive polymers possess the ability to change their volume in response to ambient changes in temperature. For biomedical application, a temperature range around physiological temperature represents an attractive stimulus for this transition. The volume change occurs due to an interplay of enthalpy and entropy between the solvent water and the polymeric chains. In these hydrogels, the polymer chains show a different and reversible molecular organization, depending on their critical solution temperature. In low critical solution temperature (LCST) polymers, the molecular chains interact with the aqueous medium, display extended coil configuration, and exist in a hydrated, swelled state. As the temperature exceeds the LCST, water is expelled leading to a collapsed, shrinking state of the hydrogel. Typical examples of LCST polymers which undergo de-swelling at the near-physiological range of 31–35 C include poly (N-isopropylacrylamide) (PNIPAAm) and poly (oligo (ethyleneglycol) methacrylate) (POEGMA). There is also upper critical solution temperature (UCST) thermo-responsive hydrogels (see Fig. 7). Famous examples include polyacrylic acid (pAAc) and polyacrylamide (pAAm). Here, above the UCST, the water solvent obtains enough energy to displace the strong intermolecular interactions between the polymer chains, thus effectively swelling the hydrogel. These polymers change into their sol state above their UCST [59]. The examples mentioned above represent the most frequently studied thermoresponsive hydrogels, which are synthetic. Therefore, research has focused on improving their biocompatibility by combining these systems with polysaccharides or proteins (by, e.g., co-polymerization, IPNs) or by the addition of bonds susceptible to enzymatic or hydrolytic attack. For example, synthetic block-copolymers of polyesters, such as poly(ε-caprolactone-co-lactide)-poly(ethylene glycol)-poly(ε-caprolactone-co-lactide) (PCLA-PEG-PCLA) have been applied as biodegradable thermo-responsive hydrogels [60]. Many polysaccharides including chitosan, chondroitin sulfate, and dextran have been combined with thermo-responsive hydrogels

Fig. 6 A smart hydrogel exhibits a change in property, in response to a stimulus. The triggers may arise from the endogenous environment (cells, tissue, body) or be controlled by the user. The hydrogel responds to the stimulus typically by a change in its swelling properties, with an associated volume/shape transformation. Other changes in material properties like sol-gel/gel-sol transitions are also possible. The volume/shape changes and forces arising in the hydrogel can be used to flush drug and cell payloads out, to stimulate encapsulated cells, to exert forces on other objects as soft actuators, to construct a biosensor device and many more related purposes in tissue engineering, soft machines, biosensing, and drug delivery

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Fig. 7 Thermo-responsive hydrogels display a different swell state according to their lower (LCST) or upper critical solution temperature (UCST). These shape and volume changes can be used in drug delivery, cell sheet technology and to create soft actuators

to engineer biocompatible injectable hydrogels [61]. Thermo-reversible peptides, like elastin-like polypeptides (ELPs), undergo a conformational change in response to temperature and have also been explored as biodegradable injectable systems [62]. Interested readers wanting to learn more about ELPs and other tunable protein hydrogels are referred to the book chapter “Tunable protein hydrogels: present state and emerging developments” in this book volume. Thermo-responsive hydrogels have been widely studied for biomedical applications. LCST polymers like PNIPAAm and its derivatives have been used to develop temperature-responsive cell culture plate surfaces which swell when the temperature is lowered, leading to non-enzymatic detachment of cell sheets. This technique is advantageous, when it is not desirable to use enzymes for the passaging of cells, as it may damage cell surface proteins and disrupt cell–cell junctions. In addition, this technique has been used to manufacture cell sheets, stacks, and rolls for cardiac and vascular tissue engineering [63]. Another widely explored venue of application for thermo-sensitive polymers uses their in situ gel forming properties for injectable cell or drug delivery. The low-viscosity hydrogel in its sol state is mixed with cells or a bioactive substance at room temperature. The mixture then forms a gel at the physiological temperature of the body, when injected through a needle or a catheter. For a comprehensive description of the applications of thermo-responsive hydrogels in drug delivery and tissue engineering, the reader is referred to these detailed reviews [64, 65]. Beyond the above-mentioned examples, thermo-sensitive hydrogels have been used for the creation of wound dressings, which improve wound healing by controlling moisture and allowing a more pain-free removal of the dressing from wounds [66, 67].

3.2

pH-Responsive Hydrogels

A lot of research has focused on smart hydrogels, which swell and change volume in response to the pH of the milieu. Various physiological niches have their own distinct pH environments (like acidic in the stomach, mildly acidic in the vagina,

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or slightly alkaline in the blood and colon). Pathogenic conditions like tumors lead to local pH decrease, due to growth under hypoxic conditions and fast lactate production. Bacterial infections are also associated with lowering of pH. For these reasons, pH-responsive smart hydrogels have been developed especially for drug loading and targeted therapy, serving as the earliest example of hydrogels used as responsive biomaterials [68]. pH-responsive hydrogels are essentially polyelectrolytes, containing multiple weak acidic or basic groups which can be ionized depending on the pH of solution. When charge builds up along the polymeric backbone, the hydrogel effectively functions as a semi-permeable membrane, only allowing the transport of oppositely charged ions. As a result of the differential ionic localization, an osmotic pressure difference builds up between the inner hydrogel and its surroundings, leading to swelling or deformational shape change of the structure. There are both basic and acidic pH-responsive hydrogels, depending on the charge of the pendant side groups. Examples of basic hydrogels include those bearing amine or amide groups like pAAm or poly dimethylaminoethyl methacrylate (PDMAEMA), while acidic hydrogels include those with carboxylic, phosphonic, or sulphonic side groups, e.g. poly(methacrylic acid) (PMAA), poly vinylphosphonic acid (PVPA). Basic hydrogels undergo swelling at pH lower than their pKa, while acidic hydrogels swell when the pH is higher than their pKa. This property can be used to tailor the swelling and drug release of pH-responsive hydrogels to areas of the body with a specific pH milieu or to target sites with local acidosis (e.g., wounds, tumors, heart disease) [69]. Similar to the discussion above with thermo-reversible polymers, it might be necessary to improve the biodegradability of synthetic pH-responsive polymers by combination with natural hydrogels or by the engineering of hydrolytic [70] or enzyme-cleaved bonds [71]. Alternatively, many natural polymers contain ionizable side groups and show natural pH-responsive behavior. Examples include proteins like gelatin [72] or charged polysaccharides like chitosan [73] and alginate [74]. These hydrogels have also been widely used as pH-responsive systems. A drawback of pH-sensitive hydrogels is that they often lack the sensitivity or specificity for fine-tuning more complex applications. For this reason, pH-responsive polymers are often combined with enzymes, which perform a specific bioconversion, leading to a lowering or rising of pH, which is then used to locally change the swelling behavior of a pH-responsive material. The most famous example of such systems are glucose-sensitive insulin-releasing hydrogel depots for treatment of diabetes mellitus. In this application, the enzyme glucose oxidase (GOD) converts glucose to gluconic acid, leading to a drop in pH and release of insulin loaded in a pH-responsive hydrogel like chitosan [75]. Beyond the release of loaded active substances, pH-responsive hydrogels have been used in biosensors and as actuators in biomedical devices. In biosensors, the volume change occurring due to the pH-dependent swelling of a thin hydrogel film is used to deflect a microcantilever and transduce the biochemical signal to an, e.g., optical signal [76]. In microfluidic devices, pH-responsive hydrogels were used to

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create micro valves for controlling flow without any electrical or mechanical components [77].

3.3

Photo-Responsive Hydrogels

Light tunability is an attractive property to engineer into a hydrogel, because it allows the remote manipulation of material properties, in the most precise spatial and temporal manner. For this reason, photo-tunable hydrogels have been explored as actuators, as drug delivery systems and in basic research to uncover the secrets of the dynamic 3D cellular microenvironment [78]. In a photo-responsive hydrogel system, light can be used to break or create bonds, leading to controlled and local gel degradation or stiffening. Alternatively, light can induce a conformational change of the polymer chains, which modifies their interaction with the solvent, thus altering the swelling behavior of the hydrogel and leading to a volume change. To create a light-responsive hydrogel, it is necessary to functionalize the hydrogel with a photoactive moiety in the hydrogel structure, which captures the light signal leading to a photochemical reaction (see Fig. 8). Depending on the mechanism triggering the light-induced change, the photoresponsive reaction can be reversible or irreversible. The following discussion differentiates photo-responsiveness from photo-crosslinking. In photo-responsiveness, we investigate the use of light to change a property of a smart hydrogel network. In photo-curable materials, light is used to initiate crosslinking of the precursor, leading to hydrogel formation. Typical examples of photosensitive moieties for reversible swelling changes include azobenzene (with its trans-cis configuration change) or spiropyran (and its ring opening and closing reaction). These photo-isomerizations lead to conformational changes in the associated polymer chains, thus affecting interactions with the solvent water and leading to swelling and hydration. The light-induced change in hydrogel volume is especially useful for designing soft actuators, which are triggered remotely via light. In this field of application, full reversibility is important to create truly useful materials. For example, Schiphorst and colleagues co-polymerized a hydrogel with spiropyran chromophores and demonstrated how the resulting soft valves can be used for flow control in a microfluidic device with non-contact operation via LED-illumination [79]. Since UV light is short wavelength and possesses high-energy intensity, it is frequently applied for light-tunability. However, UV light is undesirable in some biological and medical applications, which is why there have been approaches developed for visible or infrared light. Even with near infrared (NIR) light lasers, the penetration into tissues is limited to about 6 mm, which limits some in vivo applications, but can be used to build effective light-responsive drug delivery hydrogels [80]. In such systems, using the photo-thermal effect, thermo-responsive hydrogels are combined with nanoparticles that absorb light and convert it to localized heat dissipation, thus driving swelling changes in the polymer. Many

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Fig. 8 Common photoactive moieties used to modify hydrogel chains to create photo-tunable materials. Light-reversible isomerizations change the conformation of associated hydrogel chains and lead to light-induced swelling and de-swelling, while light-activated bond cleavage or addition contributes to local bond breakage or formation, leading to controlled gel softening or stiffening

nanoparticles and organic compounds have been studied for their photo-thermal effects in such applications. Examples include gold nanoparticles [81], carbon nanotubes [82], and organic photo-thermal agents such as copper chlorophyllin [83]. For instance, a biodegradable drug delivery system composed of low-melting point agarose and black phosphorous nanosheets has been used effectively in animal models to treat subcutaneous cancers. Here, NIR light was used to precisely heat the nanocomposite hydrogel at the therapy site, which led to gel melting and unloading of the drug. It was possible to control the drug release rate through manipulating light intensity and exposure duration [84].

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Light-responsive hydrogels are also ingeniously used to mimic the dynamic nature of the ECM and to study how matrix and cells interact and influence each other. For these experiments, cells are first encapsulated in a light-responsive hydrogel of defined and stable stiffness and crosslink density. Then, light signals can be used to locally break bonds or create additional crosslinking, leading to controlled local gel degradation or stiffening. To engineer such responsiveness, it is necessary to covalently crosslink the hydrogel in the presence of photo labile protective groups. Examples of photo labile groups include o-nitrobenzyl esters and coumarin derivatives. Upon light illumination, these bonds are photo-cleaved, thus decreasing the crosslink density of the hydrogel. To achieve the opposite effect, bond formation and crosslink density increase, other photoactive compounds, e.g. cinnamic acid, which shows reversible light dimerization, can be used. The group of Anseth modified HA hydrogels with RGD peptide, o-nitrobenzyl acrylates, and methacrylate groups. The resulting hydrogels could be sequentially softened and then stiffened with cytocompatible photoreactions. Reversible cellular mechanosensing was demonstrated in human mesenchymal stem cells (hMSCs) encapsulated in these hydrogels by changes in cell area and the nuclear YAP/TAZ ratio [85]. Beyond mechanical changes in the hydrogel, it is possible to engineer lightresponsive materials with light-activated ligand presentation or light-controlled growth factor release, thus further mimicking the dynamic biophysical and biochemical nature of the ECM. In addition, light can be used for establishing functional group presentation with defined spatial orientation within the hydrogel (photopatterning), and even for the temporal control of functional groups presentation in situ. For example, DeForest and Tirell created a biomimetic synthetic hydrogel, in which the ECM protein vitronectin was patterned and presented to cells at specific locations in the gel in a reversible manner. A photoprotection-oxime ligation sequence allowed user-defined presentation of the full-length protein to cells and an o-nitrobenzyl ester photo scission site allowed consequent removal of the protein. MSCs were shown to reversibly react to these local matrix fluctuations by changing their differentiation at the respective hydrogel sites [86]. Most photo-responsive hydrogels described above are based on synthetic hydrogels modified with photoactive moieties. Another novel strategy uses the diversity of nature’s own photoreceptor proteins and genetic engineering (GE) to build photoactive proteins into light-responsive hydrogel systems. The group of Fei Sun thus designed GE protein hydrogels for the release of stem cells and model globular proteins upon illumination with light [87]. The recombinant hydrogel is based on the bacterial photoreceptor CarHC and SpyTag-SpyCatcher chemistry. CarHC forms tetramers in the dark in the presence of adenosylcobalamin (AdoB12) and rapidly dissociates back to the monomeric form through cleavage of the C-Co bond upon light exposure. This is but one example of using GE to create smart biomaterials. Readers curious to learn more about assembling genetically engineered proteins into modular molecular hydrogels with tunable and stimulus-responsive properties are referred to the chapter “Self-assembly and genetically engineered hydrogels” in this book volume.

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Biomolecule-Responsive Hydrogels

Living organisms coordinate their responses to their internal and external milieu via a cascade of specific biological recognition events, based on the exact and reproducible interaction between affinity partners like enzyme-substrate, ligand-receptor, protein/protein complex, promoter-transcription factor pair, or nucleic acid base pair hybridization. These biological mechanisms have also been applied for the engineering of specific biomolecule-responsive hydrogels. Biomolecule-responsive hydrogels show great promise as smart materials in the area of biosensors and targeted drug delivery. Biomolecule-responsive hydrogels are usually biohybrids, composed of a synthetic polymer and a biological recognition element. A prominent example of such hydrogels are DNA hydrogels. Here, the hydrogel is composed either entirely from DNA or is a biohybrid structure made up of a synthetic or natural hydrogel which incorporates a DNA-responsive element (e.g., an aptamer, a single-stranded DNA (ssDNA), or a promoter sequence) by entrapment or covalent crosslinking. The DNA element performs a precise recognition of its target and eventually leads to changes in swelling, sol-gel transition, or other property of the gel matrix [88, 89]. Aptamers are an especially attractive option for the creation of such biomoleculeresponsive hydrogels, because they can be selected against almost any target (small molecules, proteins or even cells). Aptamers are short single-stranded oligonucleotides that adopt a specific folding and bind with high specificity to their target. Examples of aptamer-modified hydrogels, as well as detailed description of their design, applications and the challenges associated with their use are described in the book chapter “Aptamer-Modified Hydrogels” in this book volume. Beyond using oligonucleotides as sensing elements, it is possible to employ promoter-transcription factor pairs to create highly specific sensing elements. In an elegant example of such system, the group of Weber created a biohybrid hydrogel, which detects elevated uric acid concentrations and responds by dissolution and release of the loaded enzyme urate oxidase, which degrades the uric acid [90]. Elevated urate levels are indicated in arthritic gout and in tumor lysis syndrome. The hydrogel was created by grafting the uric acid sensitive transcription factor HucR into polyacrylamide. The protein-grafted polymer was crosslinked into a hydrogel by the operator DNA sequence hucO, which binds HucR specifically. Elevated uric acid levels lead to urate binding to HucR and its dissociation from the DNA crosslinker sequence. As a result, the hydrogel dissociates, releasing its uric aciddigesting cargo. Readers wanting to learn more about how the tools of synthetic biology are used to create similar hydrogel systems as highly specific devices in biosensing are referred to the book chapter “Synthetic Biological-Empowered Hydrogels for Medical Diagnostics” in this book volume. Biomolecule-specificity is difficult to engineer in a completely synthetic polymer. Approaches to accomplish this in synthetic hydrogels involve molecular imprinting. In molecular imprinting, the synthetic hydrogel is crosslinked in the presence of a template molecule. After polymerization, the template is removed, but the polymer

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retains a specific affinity architecture to its template [91]. Completely synthetic molecular imprinted hydrogels have been described for the biosensing of glucose [92] and for drug delivery [93]. The technology is mostly effective for small biomolecules (< 1,500 Da) and challenges remain for the engineering of the imprinting of proteins and larger structures [94].

3.5

Electromagnetic-Responsive Hydrogels

Other examples of smart tunable hydrogels include those which can swell or change volume in the presence of external magnetic fields or when an electric current is applied. Such systems are interesting from a biomedical perspective, in that they offer external and user-controlled input of the required response [95]. For magneticresponsive hydrogels, it is essential to add magnetic particles (e.g., iron oxide, nickel nanorods) to the polymer network by physical entrapment or, preferably, by covalent or coordination bonds. When an external magnetic field (EMF) is applied, the magnetic particles can exert torque forces on the polymer network, effectively bending, stretching, or changing the shape of the hydrogel. The rapid response time of such systems makes them attractive as soft actuators [96] and in the delivery of therapeutic agents (Fig. 9). The release of the drug is frequently associated with convection movements arising from the rapid hydrogel volume changes in the magnetic field, effectively flushing the drug out at the required space and time [97].

Fig. 9 The addition of magnetic nanoparticles makes it possible to induce shape changes of hydrogels under the influence of an external magnetic field (a) Drug release from macroporous hydrogels as a result of convection movements arising from EMF stimulation (b) Induction heating causes de-swelling and drug release from a thermo-responsive hydrogel (c) Physical stimulation of encapsulated cells beneficial for tissue engineering (d) Magnetic templating to produce aligned pore structure in a scaffold and (e) shape change can be used to design a soft actuator

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Especially if a thermo-sensitive hydrogel is used as a carrier for the magnetic particles and is subjected to a rapidly alternating magnetic field (> 10 kHz), electromagnetic induction heating leads to temperature change and swelling of the thermo-responsive hydrogel network. This approach has been used for targeted drug delivery and in cancer therapy to induce local hyperthermia to selectively kill cancer cells. Many researchers boost the efficacy of such systems by combining both hyperthermia and chemotherapeutic drug delivery in a single magneto-responsive hydrogel, which has been shown to reduce tumor sizes in cell culture and in animal models in a more effective manner than when a single therapy mode was used [98, 99]. Magneto-responsive hydrogels are also being increasingly explored in tissue engineering applications, especially for the regenerative therapy of bone, cartilage, and nerve tissue [100]. Here, the magnetic nanocomposites offer enhanced mechanical strengths and the ability to exert torque on EMF allows to actually mechanically stimulate cells within the scaffold, which may induce differentiation into required lineages [101], enhance proliferation [102], or provide topological clues for directed growth [103]. Magneto-responsive hydrogels are also used to create structured scaffolds for tissue engineering. By performing magnetic templating, structures with aligned pore microarchitecture [104] or with multiscale anisotropic topographic features are generated [105], which constitutes an elegant bottom-up approach for the fabrication of micro-structured scaffolding materials. Electrically responsive hydrogels are inspired from the biology of muscle. These hydrogels swell or shrink when subjected to an electric signal (e.g., voltage, current). Electrically active hydrogels are usually polyelectrolytes, which have ionizable side groups. When an external electrical field is applied, an equilibrium is reached between the motion of external ions in the solvent medium and the ionizable groups with fixed charges on the polymer backbone. This local distribution of charge between the external and internal compartment of the hydrogel leads to an osmotic pressure difference that drives swelling or shrinkage. The volume response of an electrically responsive hydrogel can be tuned by varying the network characteristics (e.g., macromer/monomer concentration and mesh size), by modifying the ionic strength of the solvent, as well as by varying the relative position of electrodes and the hydrogel. Using electric signals as stimulus has great advantages, because electric fields can be applied externally with high precision and magnitude control [106]. Synthetic electrically responsive hydrogels include: pVA, pAAc, poly (2-acrylamide-2-methylpropanesulfonic acid) (PAMPS), sulfonated polystyrene, and poly (4-hydroxylbutyl acrylate) (poly (4-HBA)) among others [107]. In addition, many natural hydrogels can exhibit electric responsiveness, as they are natural polyelectrolytes. Examples include proteins or charged polysaccharides like chitosan, hyaluronic acid, and chondroitin sulfate. In biomedical applications, natural electro-responsive hydrogels are frequently combined with synthetic polymers to create semi-synthetic electroactive hydrogels [108]. The propensity of electrically responsive hydrogels to change shape in electrical field has been used to design soft actuators [109] and artificial muscles [110]. The electrically induced de-swelling or

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other shape changes, which lead to solute movement within the gel, are used for drug release applications [106]. The vast majority of drug release studies with electro-responsive systems have been performed in vitro. Most systems involve the use of a transdermal hydrogel patch, which can release drugs on demand after electric stimulation on the skin [111]. In an interesting alternative example, Ying et al. developed nanogels from electro-responsive hydrogels, which are meant to cross the blood–brain barrier and deliver anti-epileptic drugs under electroencephalograph epileptiform abnormalities, showing a strategy which can be used to manage epilepsy and reduce side effects for existing medications [112]. The reader should note that in the literature electro-responsive and electroconductive hydrogels are sometimes used as synonyms, although the terms denote different type of materials. Electro-conductive hydrogels are composite systems, which display high electrical conductivity and electrochemical redox properties, which can be used to perform electrical stimulation of tissues, as well as to detect electric signals generated in biological systems. Electro-conductive hydrogels are obtained by introducing conductive components into conventional hydrogels, such as by grafting conductive polymers like polypyrrole into the hydrogel backbone or by the addition of conductive nanoparticles like carbon nanotubes (CNT). Readers interested in the synthesis and biomedical applications of electro-conductive hydrogels are referred to these informative reviews [113, 114]. Electro-conductive hydrogels can be used to create electro-responsive systems. For example, the Zare group created an electro-conductive hydrogel by dispersing conductive polypyrrole polymer nanoparticles in a thermo-sensitive hydrogel. The authors could demonstrate the dosage-controlled release of drugs in a mouse model, when a weak external direct current was applied [115].

3.6

Other Responsive Hydrogel Systems

Beyond the major systems covered so far, there are other responsive hydrogels, which are relevant for biomedical applications and show the great ingenuity of material researchers in designing smart hydrogels for diverse biomedical purposes. In this section, we will shortly cover several more exotic, but relevant examples. These include ultrasound-responsive and redox-responsive hydrogels, and because of a large body of research devoted to managing diabetes, glucose-responsive hydrogels based on phenylboronic acid. The use of ultrasound for imaging is a widely established, cost-effective, and mature medical technology. Ultrasound (US) consists of pressure waves at frequencies of 20 kHz or greater and displays deep penetration into tissues. For these reasons, there have been efforts to design hydrogels, which are ultrasoundresponsive, releasing their drug payload at specific times and locations, after US stimulation [116]. The first developed systems of this kind used ultrasound contrast agents, which is gas microbubbles encapsulated in a biodegradable shell. The US waves cause oscillation of the contrast agent, leading to their oscillatory movement

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and rupture, which is a mechanism that can be used to release drugs, encapsulated into the microbubbles. The hydrogels used in these systems usually serve as carriers for the microbubbles [117, 118] or make up the biodegradable shell of a hollow nanogel serving as a microbubble [119] and allow greater spatiotemporal control of the drug release. Recently, ultrasound-responsive hydrogels for drug delivery have been reported, which do not carry any microbubbles, but respond to the US directly due to disruption of reversible host–guest interactions [120] or dynamic covalent bonds [121]. In other elegant applications in tissue engineering, ultrasoundresponsive hydrogels with loaded microbubbles have been used as porogens for the creation of porous scaffolds [122], for the remote mechanical stimulation of seeded cells [123], as well as for the in vitro and in vivo transfection of hydrogelencapsulated cells [124]. Another material category relevant especially for drug/gene delivery include redox-responsive hydrogels. One approach to create such systems involves the use of disulfide crosslinking for formation of the hydrogel. Upon entering a cell, the hydrogels dissociate due to reduction of the disulfide bridges by cytosolic glutathione, thus delivering their drug or gene payload. This strategy has been especially studied for the creation of nanogels for targeted cancer therapy, as tumor cells display a more reducing microenvironment than healthy cells [125]. For example, Pan et al. created dual pH/redox-responsive nanogels from PMAA and N,N-bis (acryloyl)cystamine as a disulfide-functionalized crosslinker. They could show in in vitro experiments that the nanogels were nontoxic to normal cells, but could effectively deliver doxorubicin to glioma tumor cells [126]. Studies about the localization and exact internalization mechanism of redox-sensitive nanogels have been performed, but further improvements in specificity and targeting are necessary [127]. One way to achieve this is to use target-specific aptamers on, e.g., DNA hydrogels. Biocompatible and redox-responsive DNA nanogels with disulfide links have been described for gene/drug delivery and cancer therapy [128]. In the area of cell therapy, Kar et al. could show in animal experiments that disulfide-modified PEG hydrogels are effective for stem cell delivery, and that the rate of degradation can be tuned by the type and number of encapsulated cells, as well as by tuning the fraction of disulfide moieties in the hydrogel backbone [129]. Worldwide, diabetes is one of the most common chronic conditions afflicting humans and absolute numbers of patients are projected to keep rising. For this reason, research into the creation of smart sensors for glucose monitoring and insulin delivery is a very active field [130]. We shortly discussed an example of a smart hydrogel for diabetes management in the section devoted to pH-sensitive tunable hydrogels. In this sensor, the sensitive moiety was the enzyme GOD, which converts glucose with an associated pH change of the milieu, causing shrinkage of the pH-sensitive hydrogel, and release of the encapsulated insulin cargo [131]. Other sensor strategies utilize the carbohydrate-binding lectin protein Concanavalin A as the glucose-binding partner [132]. Both designs are based on the affinity of specific proteins to glucose. However, building sensors with proteins brings its own set of disadvantages, as proteins are prone to denaturation and may cause immunogenic effects. This makes such hydrogel biosensors unsuitable for long-term and

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Fig. 10 The formation of a charged PBA-glucose complex is swift and reversible. The charged complex swells the hydrogel network in relation to the glucose concentration, making it possible to design synthetic insulin management systems

implantable use. Alternative approaches to develop stable devices for diabetes management include glucose-responsive hydrogels based on synthetic phenylboronic acid (PBA) copolymers [133]. PBA binds reversibly to 1,2- and 1,3-diols (including glucose) leading to the formation of a charged phenylboronate complex (see Fig. 10). The formation of the charged complex results in osmotic pressure changes in the hydrogel and a corresponding volumetric change, which can be used to release insulin payloads in physiological conditions [134]. Current research is focused on engineering injectable and biocompatible PBA-based hydrogels [135]. Recent studies have shown that a single injection of insulin-loaded PBA-based hydrogel led to a two-week steady-state glucose level in a diabetes mouse model [136]. These advances bring closer the hope of a clinically viable, glucose-responsive insulin delivery system, which may one day be able to mimic the pancreatic activity of healthy individuals.

3.7

Self-Healing Hydrogels

Self-healing hydrogels are a special category of tunable materials, which exhibit the ability to restore their mechanical properties and/or shape after being subjected to deformational forces. Inspired by native tissues, which can repair themselves after injury, self-healing hydrogels should ideally be able to recover themselves repeatedly and rapidly after stress. In order to display self-healing properties, the hydrogel network has to be able to allow flexible movement of the polymer chains within the bulk hydrogel and to display the ability to restore broken links. These properties originate from the presence of reversible and dynamic linkages throughout the

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polymer network, which are able to reconnect the polymer chains to one another. The mechanisms behind self-healing can involve both covalent bonds like dynamic covalent bonding and coordination bonds, as well as a diverse array of physical interactions like hydrogen bonding, ionic, hydrophobic, and host–guest interactions (Fig. 11). These reviews offer detailed descriptions of the reversible crosslink mechanisms leading to self-healing functionality [137–139] and the methods used to measure self-healing efficiency [140]. Self-healing properties can be tuned by careful selection of the hydrophilic hydrogel backbone and the specific reversible crosslinking linkages/motifs

Fig. 11 The different mechanisms behind the design of self-healing hydrogels, figure reproduced from Talebian et al., under the terms of the Creative Commons Attribution License 4.0 [139], © The Authors

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throughout the polymer. In addition, it is possible to combine self-healing properties with additional stimulus-responsiveness, relevant for physiological environments like pH [141] or redox-responsiveness [142]. Both synthetic and natural hydrogels, as well as semi-synthetic materials have been used to engineer self-healing hydrogels [139]. Due to dynamic nature of the reversible crosslinks, self-healing hydrogels also typically exhibit shear-thinning behavior, which means that their viscosity decreases when subjected to shear stress [143]. As a result, they have been widely investigated as injectable hydrogels, for cell and drug delivery, as well as bioinks for 3D bioprinting [144]. In these cases, the hydrogel can be loaded with therapeutic substances or cells and be injected by application of shear stress through a needle at the desired site or location. After removal of the shear stress, the shear-thinning hydrogel reassembles its structure, leading to in situ gelation at the target site, or in the case of bioprinting, to the formation of a spatially defined 3D construct. The ability to withstand repeated mechanical stresses has made self-healing hydrogels attractive materials for the tissue engineering of load-bearing tissues like bone and cartilage [145, 146]. In cardiac tissue engineering, self-healable hydrogels protect sensitive regenerative cells, which are subjected to high pulsatile stresses when transplanted at the infarction site for the repair of damaged myocardium [147]. In a similar vein, self-healing hydrogels have been used for the tissue engineering of muscle [148], neural [149], and vascularized tissues [150]. Other biomedical applications include the creation of durable adhesive wound dressings, which withstand skin movement stress and promote healing [151], as well as injectable tissue adhesives to seal stomach wounds under highly acidic conditions [152]. Finally, self-healing hydrogel materials, especially those with additional electrical conductivity, are increasingly being explored as soft actuators opening up possibilities of designing wearable soft electronics, soft robots, and biosensors [107].

4 Outlook The first hydrogels in healthcare were introduced already in the 1950s and 1960s and were mainly prized as bioinert materials that minimized protein/cell interactions. Their primary use was intended as optical lenses, blood-contacting materials, and anti-fouling coatings, although other applications like drug delivery and tissue engineering were already being explored [153]. Since then, rapid advances in material science and bioengineering have turned hydrogels into one of the most promising classes of materials for biomedical applications. Hydrogels have been transformed from inert, static materials to dynamic structures, which display specific bioactivity and responsiveness to environmental cues. These stimuli can either derive from the user (e.g., light, electromagnetic field, ultrasound) or can arise autonomously from the cells/tissues/organism itself (e.g., pH, glucose concentration, enzyme secretion, redox potential change). Hydrogels have now become smart

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materials and our deepening understanding of structure-function relationships, as well as the interaction between the material-biology interface, allows the design of ever more sophisticated and tunable hydrogels. In the last decade, research has concentrated on enhancing the biomimetic properties of hydrogels, bringing together spatiotemporal control of stiffness and presentation of bioactive ligands, through the introduction of multiple and dynamic orthogonal chemistries [154]. The mechanical and load-bearing properties of hydrogels have also been greatly improved via IPN, NC, and DN designs [155]. These advances, together with the development of fabrication technologies, which allow the creation of more complex hydrogel architectures on different hierarchies of scale, have brought about great progress in the area of tissue engineering and 3D cell culture. In the area of drug delivery and biosensing, the specificity of hydrogels has been further enhanced by engineering multiple responsiveness. We already covered examples like dual pH and redox-responsiveness for improved drug delivery to cancer cells, but the development of multi-responsive hydrogels with several sensing and response pathways is a key step in the successful integration of smart hydrogels within physiological systems. Researchers are already developing hydrogels with complex gate architectures, which perform controlled operations like drug release according to in-built Boolean logic gates (AND or OR logic-gate response), in response to the presence of multiple specific biochemicals [156, 157]. Despite these advances, it must be said that many studies remain proof-of-concept reports, few systems reach clinical trials, and many challenges still need to be overcome until smart hydrogel materials are translated into clinical practice [158, 159]. For clinical applications, it is necessary to produce hydrogel materials on a large scale under standardized conditions. The chemical synthesis processes may need to be simplified and potentially toxic reagents must be substituted with nontoxic alternatives. Scalable production of protein hydrogels by recombinant synthesis could be an option for sourcing standardized protein hydrogel materials [160]. Moreover, the fate of materials which are implanted or injected in the human body still needs to be thoroughly studied – especially in regard to the safety of unreacted monomers/macromers and residual initiators, as well as induced foreign body responses [161]. More research is also required into the degradation profiles of hydrogels alongside the biocompatibility of all breakdown products (e.g., with nanocomposites). Additional investigations with in situ analytical methods are needed, which are able to monitor degradation, structure, and function changes directly in implanted/injected hydrogels. These should be complemented with suitable in vitro assays, which are better able to predict the in vivo performance of hydrogels. For example, magnetic resonance elastography (MRE) has been used to study the mechanical properties of engineered tissues both in vitro and in vivo, and could be one technology, which speeds up transition to more cost-effective clinical trials [162]. Despite the existing challenges, interdisciplinary researcher teams will keep on learning how to create intelligent and tunable hydrogel biomaterials for improving healthcare and quality of life well into the twenty-first century.

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147. Gaffey AC et al (2015) Injectable shear-thinning hydrogels used to deliver endothelial progenitor cells, enhance cell engraftment, and improve ischemic myocardium. J Thorac Cardiovasc Surg 150(5):1268–1277 148. McKinnon DD et al (2014) Biophysically defined and cytocompatible covalently adaptable networks as viscoelastic 3D cell culture systems. Adv Mater 26(6):865–872 149. Tseng TC et al (2015) An injectable, self-healing hydrogel to repair the central nervous system. Adv Mater 27(23):3518–3524 150. Tseng T-C et al (2017) Glucose-sensitive self-healing hydrogel as sacrificial materials to fabricate vascularized constructs. Biomaterials 133:20–28 151. Zhao X et al (2017) Antibacterial anti-oxidant electroactive injectable hydrogel as self-healing wound dressing with hemostasis and adhesiveness for cutaneous wound healing. Biomaterials 122:34–47 152. Phadke A et al (2012) Rapid self-healing hydrogels. Proc Natl Acad Sci 109(12):4383–4388 153. Kirschner CM, Anseth KS (2013) Hydrogels in healthcare: from static to dynamic material microenvironments. Acta Mater 61(3):931–944 154. Arkenberg MR, Nguyen HD, Lin C-C (2020) Recent advances in bio-orthogonal and dynamic crosslinking of biomimetic hydrogels. J Mater Chem B 8(35):7835–7855 155. Vedadghavami A et al (2017) Manufacturing of hydrogel biomaterials with controlled mechanical properties for tissue engineering applications. Acta Biomater 62:42–63 156. Ikeda M et al (2014) Installing logic-gate responses to a variety of biological substances in supramolecular hydrogel–enzyme hybrids. Nat Chem 6(6):511–518 157. Badeau BA et al (2018) Engineered modular biomaterial logic gates for environmentally triggered therapeutic delivery. Nat Chem 10(3):251 158. Williams DF (2019) Challenges with the development of biomaterials for sustainable tissue engineering. Front Bioeng Biotechnol 7:127 159. Li J, Mooney DJ (2016) Designing hydrogels for controlled drug delivery. Nat Rev Mater 1 (12):1–17 160. Sonmez C, Nagy KJ, Schneider JP (2015) Design of self-assembling peptide hydrogelators amenable to bacterial expression. Biomaterials 37:62–72 161. Morais JM, Papadimitrakopoulos F, Burgess DJ (2010) Biomaterials/tissue interactions: possible solutions to overcome foreign body response. AAPS J 12(2):188–196 162. Othman SF, Xu H, Mao JJ (2015) Future role of MR elastography in tissue engineering and regenerative medicine. J Tissue Eng Regen Med 9(5):481–487

Adv Biochem Eng Biotechnol (2021) 178: 37–62 https://doi.org/10.1007/10_2020_161 © Springer Nature Switzerland AG 2021 Published online: 6 February 2021

Alginate Hydrogels with Tuneable Properties Alan M. Smith and Jessica J. Senior

Contents 1 2 3 4 5 6 7

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Origin and Chemical Structure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Ionotropic Gelation Mechanism . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Methods of Ionotropic Gelation: Internal and External Gelation . . . . . . . . . . . . . . . . . . . . . . . . . . . . Enzymatically Controlled GM Sequences . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alginate Degradation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Modified Alginates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1 Covalently Crosslinked Alginate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.2 Amphiphilic Alginate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.3 Oxidised Alginate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.4 Peptide Modified Alginate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Outlook and Future Directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

38 39 41 44 47 48 49 49 51 53 54 55 56

Abstract Alginate is a material that has many biomedical applications due to its low toxicity and a variety of favourable physical properties. In particular, the ease in which hydrogels are formed from alginate and the variety of mechanical behaviours that can be imparted on the hydrogels, by understanding alginate chemistry and intuitive design, has made alginate the most widely investigated polysaccharide used for tissue engineering. This chapter provides an overview of alginate, from how the source and natural variations in composition can influence mechanical properties of alginate hydrogels, through to some innovative techniques used to modify and functionalise the hydrogels designed specifically for cell-based therapies. The main focus is on how these strategies of understanding and controlling the chemistry of alginates have resulted in the development of hydrogels that can be tuned to

A. M. Smith (*) and J. J. Senior Department of Pharmacy, School of Applied Sciences, University of Huddersfield, Huddersfield, UK e-mail: [email protected]

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A. M. Smith and J. J. Senior

deliver the physical behaviours required for successful application. This will also highlight how research on the physicochemical properties has helped alginate evolve from a structural polysaccharide in brown seaweed into a highly tuneable, multifunctional, smart biomaterial, which is likely to find further biomedical applications in the future. Graphical Abstract

Keywords Alginate, Hydrogel, Tissue engineering, Tuneable

1 Introduction Alginate was first isolated and described in 1881 by E.C.C Stanford [1] and has since been widely investigated and subsequently applied in various industries. Alginate is often exploited in tissue engineering research, but the majority of commercial use is within the food and pharmaceutical industry. Indeed, as with many other polysaccharides, the original research on the chemical and physical properties of alginate was focused towards the food industry and has since been applied to more “high value” pharmaceutical and biomedical applications, which have helped to drive the development and understanding of alginate further. Probably the most useful property of alginate, with regard to tissue engineering and cell culture, is the ability to form hydrogels that are non-toxic, have excellent mass transport properties and that are formed via a mild gelation mechanism. Consequently, this has driven research on biomedical applications of alginate which has led to the process of hydrogel formation being well understood [2]. Hydrogel-forming biopolymers such as alginate are highly advantageous for the creation of three-dimensional (3D) cell culture matrices. The high water content of

Alginate Hydrogels with Tuneable Properties

39

these materials facilitates the diffusion of nutrients, waste products, and signalling molecules both to and from the cells seeded within a hydrogel matrix. The culture of cells in 3D matrices is an important precursor to tissue creation and is necessary because cells do not spontaneously assemble into 3D structures when simply seeded onto an unmodified flat surface. Surfaces modified with an ultra-low attachment coating may allow aggregation of cells in 3D yet with poor control over specific cell placement. Indeed, cells need to be able to organise and interact with other cells in a 3D space that reflects the tissue that is to be grown. This has been demonstrated by Hishikawa et al. [3] who reported that the spatial relationship between cells seeded within a 3D matrix has an influence on determining gene expression. Furthermore, the physical properties of the 3D cell substrates, such as the stiffness and elasticity, are key to the successful engineering of tissues. This has been highlighted in mesenchymal stem cells (MSC) which have been shown to differentiate towards an osteogenic linage when cultured in stiff materials and will differentiate along phenotypic lineages of softer tissue cell types, such as myocytes or chondrocytes, when cultured in more compliant materials [4, 5]. In addition to the physical properties of the substrates, the chemical properties of 3D cell culture materials also have a great impact on cell behaviour and tissue development. Alginate was originally thought to be an inert material for cell encapsulation, which relates to its biological function as a structural component in seaweed, however, its polyanionic nature promotes interactions with positively charged species and repulsive interactions with other anionic species [6]. This is an important consideration when one is culturing cells, as cell signalling molecules and cell recognition proteins will often contain charged domains, which can ultimately influence cell behaviour. In addition, subtle changes in the chemistry of alginate can dramatically change the physical properties and therefore have an impact on cells cultured via a change in mechanical behaviour of the substrate. Rather than considering alginate as an inert scaffold, alginate is now rightly thought of as a cell culture material that has the ability to interact directly and indirectly with cells both physically and chemically. Taking this into account, the ability to control or tune the physical and chemical properties of alginate opens up the opportunity to intelligently design alginates that are optimised for a particular application. This chapter intends to provide an overview of the tuneable nature of alginate hydrogels and review how the molecular structure (and as a consequence, physical properties) has been modified and controlled for use in cell culture and tissue engineering.

2 Origin and Chemical Structure Alginate is a structural polysaccharide found in brown seaweed (Phaeophyceae), providing both mechanical strength and flexibility to seaweed structure. Of the many different species of brown seaweed that contain alginates, the most commercially exploited for alginate production are Laminaria hyperborea, Macrocystis pyrifera,

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Fig. 1 Molecular structure of alginate showing the conformation of guluronate and mannuronate blocks and the block-wise structure along the polymer chain

and Ascophyllum nodosum [7]. Generally, these commercial alginates are extracted with aqueous alkali (sodium hydroxide or sodium carbonate), then filtered and precipitated with calcium before being acidified to form alginic acid. The alginic acid can then be converted to the required salt form by mixing with the appropriate metal carbonate. For detailed methods of extraction of alginate from brown seaweed, the authors recommend consulting the following articles [8–10]. Alginates can also be produced by fermentation of several species of bacteria, which include Azotobacter vinelandii [11] and Pseudomonas aeruginosa [12, 13]. These bacterial strains have helped in the understanding of the biosynthetic pathway of alginate, which has led to the production of high-quality alginates with defined material properties, produced under a controlled environment. Structurally, alginate is described as a linear polysaccharide that consists of two uronic acid residues, 1–4 linked β-D-mannuronic acid and 1–4 linked α-L-guluronic acid, which are present along the polymer chain as blocks of mannuronate (M), blocks of guluronate (G) and heteropolymeric blocks of both monomers (MG) (Fig. 1). During biosynthesis, alginate consists solely of β-D-mannuronate and some of these residues are then enzymatically epimerised at C(5) and converted to α-Lguluronate. This epimerisation results in converting the inter-residue linkages from 1–4 equatorial to 1–4 axial, having a dramatic effect on the molecular geometry of the polymer. The 1–4 equatorial bond of the polymannuronate blocks give a molecular geometry that is described as a flat extended ribbon, whereas the 1–4 axial configuration of polyguluronate blocks gives rise to a buckled ribbon

Alginate Hydrogels with Tuneable Properties

41

molecular geometry. Heteropolymeric regions result in an irregular geometry which results in sections of the polymer chain that do not pack together uniformly [14]. The physical properties, therefore, are directly related to subtle changes in MG sequence. Before discussing how the mechanical behaviour of hydrogels can be tuned, it is important to explain the mechanism of gelation.

3 Ionotropic Gelation Mechanism Alginate has a great affinity for divalent cations, in particular calcium ions, which are known to bind blocks of guluronate residues in adjacent chains. First described by Grant et al. [15], this mechanism is known as the “egg box model” and is widely accepted as the ionotropic gelation mechanism for alginate (and low methoxyl pectin) (Fig. 2). Since then, it has been shown that MG blocks can also bind calcium and contribute to gelation by binding with adjacent MG blocks and GG blocks [16] Crosslinking is therefore dependent on the M:G ratio. As a general rule of thumb, alginates with a high proportion of G residues (high G alginates) produce stronger gels when crosslinked with calcium than those with a higher proportion of M residues (high M alginates) which produce softer, more elastic gels that are less stable [17]. This allows alginate hydrogels to be tailored with predicted mechanical properties by simply selecting the composition of the alginate. Alginate is present in seaweed as a heterogeneous mixture in terms of molecular weight and MG composition, which varies depending on the area of the seaweed the alginate is extracted from [18] and with respect to different growth conditions and season [19, 20]. This natural variation results in the production of alginates with compositions that range from 10% G to 75% G, which ultimately impacts upon their physical properties. Alginates from bacterial sources produce alginates with a larger variation in compositions, with G contents varying from 1 ¼ ðFG  FMGM Þ=FGGM The value calculated from this expression has been shown to correlate with mechanical properties of resulting hydrogels (Fig. 3). Information regarding the G content and block structure of alginates from various seaweed sources is shown in Table 1. Molecular weight, as with most polymers, also has an influence on the mechanical properties of alginate. In the solution state, low molecular weight fractions show almost Newtonian flow behaviour and as the degree of polymerisation (DP) increases, they begin to exhibit pseudoplastic flow. When gelled, the strength is also dependent upon molecular weight. Martinsen et al. [17] showed, using a high G alginate that is beneath a molecular weight of 2.4  105, gel strength increased with increasing molecular weight. Once this molecular weight value was exceeded however, the gel strength became independent of molecular weight.

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Table 1 Chemical composition of alginate from various seaweed sources (adapted from [27]) Seaweed source Laminaria hyperborea (stipe) Laminaria hyperborea (leaf) Macrocystis pyrifera Laminaria digitata Lessonia nigrescens Ascophyllum nodosum Laminaria japonica Durvillea antarctica

FGM FG 0.63

FM 0.37

0.49

FGGM

Average length of G blocks G>1 15

0.11

FGG 0.52

FMM 0.26

FGGG 0.48

FMGM 0.07

0.51

0.19

0.31

0.32

0.25

0.13

0.05

8

0.42

0.58

0.21

0.20

0.37

0.16

0.17

0.04

6

0.41

0.59

0.16

0.25

0.43

0.20

0.11

0.05

6

0.41

0.59

0.19

0.22

0.40

0.17

0.14

0.05

6

0.39

0.61

0.21

0.22

0.38

0.13

0.14

0.07

5

0.35

0.65

0.17

0.18

0.48

0.32

0.68

0.17

0.16

0.51

0.12

0.12

0.05

4

(MG)

(MGG)

0.05

One of the simplest routes to controlling mechanical properties of alginates is to exploit the highly selective nature towards differing divalent cations. The affinity of alginate to different species is generally in the order of Pb2+ > Cu2+ > Ba2+ > Sr2+ > Ca2+ > Mn2+ > Mg2+ [28]. The selectivity for these metals, however, is dependent on the G content of the alginate, increasing markedly with increasing G content. M blocks and alternating MG blocks show very little selectivity. In biomedical applications, divalent calcium is most frequently used for crosslinking alginate due to minimal toxicity and low cost, whereas lead and copper have been shown to elicit adverse effects on cell viability [29–31]. FG denotes the fraction of the alginate containing guluronic acid and FM the fraction of mannuronic acid. FGG, FMM, and FGM(MG) are the fractions of dimeric blocks of guluronic acid, mannuronic acid, and mixed sequences, respectively. FGGG is the fraction of trimeric blocks of guluronic acid. FMGM is the fraction of block that contains a guluronic acid between two mannuronic acid residues and FGGM(MGG) is the fraction that begins or ends with a G block. It has also been shown that highly purified G-blocks (with degree of polymerisation of approximately 20) removed from alginate polymer chains by acid hydrolysis, refined and fractionated, can modify both the gelling kinetics and apparent equilibrium properties when mixed with an ordinary alginate [32]. The addition of these G blocks results in greater gel strength at high Ca2+ concentrations (Fig. 4), reduced sol-gel kinetics and a reduced extent of syneresis. Furthermore, in the solution state, the refined G blocks do not contribute to the viscosity of the alginate solution prior to hydrogel formation due to the low DP and therefore, chain length of

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Fig. 4 Variation in Young’s modulus of alginate hydrogels as a function of calcium concentration and the addition of 0.5% and 1.0% short chain G blocks to alginate compared with alginate alone (adapted from [33])

these G blocks. It is thought that these short chain G blocks behave as adhesive domains between regions of the gelling alginate chains that are relatively deficient in G blocks.

4 Methods of Ionotropic Gelation: Internal and External Gelation Divalent cations, such as the most frequently used Ca2+, can be introduced into alginate using different methods that can produce gels of various levels of heterogeneity. The most commonly used process for encapsulating cells is by addition of a viscous sodium or potassium alginate solution to a solution of calcium chloride. When the soluble Ca2+ comes into contact with the alginate, localised instantaneous gelation occurs at the surface of the alginate due to the rapid gelation kinetics, which makes it extremely difficult to produce homogeneous gels. Using this technique, the Ca2+ must diffuse through the alginate structure. In addition, the un-crosslinked alginate will also migrate towards the calcium at the surface, ultimately producing a hydrogel that has a higher concentration of alginate (and therefore crosslinking density) at the surface than at the centre of the structure [34] (Fig. 5). Large alginate concentration gradients have been reported that have up to five times the alginate at the surface (from the concentration in the alginate solution prior to gelation) to virtually zero concentration in the centre of the hydrogel [35, 36]. This non-uniform crosslinking is dependent on alginate concentration, molecular weight, and the concentration of the calcium ions in the alginate. The homogeneity therefore can be controlled to some extent. Maximum hydrogel heterogeneity is achieved by using a low molecular weight alginate with a low concentration of the gelling ion and maximum homogeneity is reached by gelling a high molecular weight alginate with high concentrations of crosslinking ions [36].

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Fig. 5 Schematic diagram of the external gelation of alginate using a soluble calcium source

Currently, multiple research groups are investigating alternative ways of externally introducing cation sources and subsequently tailoring the crosslinking behaviour within alginate hydrogels. Bajpai et al. [37] devised a novel system to deliver Ca2+ and Ba2+ by diffusion through dialysis tube and was able to reduce degradation and drug release rates in alginates by increasing concentrations in Ba2+ and reducing Ca2+ concentration, respectively. It was deduced that the greater ionic radius of the Ba2+ ions were better positioned within the alginate egg box structure, thus increasing stability. Some groups have adopted the use of a supporting gelatin or agarose slurry containing calcium ions, in order to 3D bioprint, support and crosslink low viscosity alginate precursor solutions [38, 39]. Other groups have 3D bioprinted alginate structures within a supporting agarose particulate bed, followed by injection of Ca2+ ions around the extruded structure to preserve shape fidelity in a more controlled manor [40–42]. Using a calcium chloride nebulizer is another approach toward bioprinting alginate and is conducive to using different alginate concentrations, yielding viable cells and stable hydrogels after 24 h cultivation [43]. Moreover, it has also been shown that the hydrogel properties may depend on the original alginate salt used. Sodium alginates have been shown to undergo slower sol-gel transitions with reduced elastic moduli compared with the potassium alginates [44]. Time exposed to the crosslinking ion source will also affect hydrogel strength and homogeneity. The longer the structure is exposed to the source of divalent cations, the more time there is for diffusion to occur and increase the crosslinking density of the alginate structure. Although often thought of as disadvantageous due to the difficulty to achieve homogeneous gels using the external gelation method, it does provide scope for tailored mechanical properties and textures throughout the hydrogel, especially when careful consideration is taken with regard to the properties of the alginate discussed in this section. To overcome the problem of heterogeneous alginate gels formed by external gelation, Draget et al. [45] developed an internal gelation mechanism that delivers calcium ions from within the alginate bulk. Calcium in an insoluble form (such as

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Fig. 6 Schematic diagram of the internal gelation of alginate using an insoluble calcium source (adapted from [46])

calcium carbonate) can be dispersed throughout a sodium (or potassium) alginate solution. The Ca2+ is then released by controlled dissociation of the CaCO3 initiated by addition of glucono-δ-lactone (GDL), which hydrolyses over a period of 40–60 min in aqueous solutions producing gluconic acid and subsequently reducing the pH. This in situ release of the Ca2+ is known as internal gelation (Fig. 6). The gelation time, gel strength, and acidity of these internally crosslinked gels can be easily controlled by careful adjustment of CaCO3 to GDL ratio. As with the external gelation, the gelation kinetics and the resulting mechanical properties are also affected by alginate concentration, molecular weight, and MG sequence [45]. One way to control rate of gelation is to select an appropriate particle size of the insoluble calcium salt, as it has been shown that smaller particle sizes favour a more rapid gelation and conversely, larger particles favour a slower gelation [16, 33]. This is likely the result of the difference in surface area affecting dissolution rate of the calcium and ultimately does not affect final gel strength once all the calcium is in solution. Other external factors, such as introducing a shear environment during gelation, can also be manipulated to markedly change the mechanical properties of the hydrogel without altering the chemical structure, the concentration of alginate or quantity of crosslinking ions. Gelation is typically induced by preparing alginate solutions and introducing divalent cations under static conditions and these “quiescent gels” comprise a continuous bulk gel network. However, when the crosslinking solution is added dropwise to alginate solutions under shear, the resulting alginate gel consists of gelled microparticles and is known as a “sheared gel” or “fluid gel” that has shear thinning flow properties rather than being a self-supporting solid hydrogel, as would be formed if gelled quiescently [47]. Interestingly, this differing mechanical behaviour has been shown to have a biological impact on cells encapsulated and then released from such systems. In particular, chondrocytes formerly entrapped and released from quiescent alginate gels show recovery of the native phenotype, whereas chondrocytes formerly suspended and released from alginate fluid gels are unable to maintain recovered phenotype, indicating that geometric entrapment is crucial for the preservation of chondrocyte phenotype despite being exposed to identical chemical environments [48].

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Other methods to control the delivery of calcium ions to alginate have included using sequestering agents to bind the calcium ions when mixed with the alginate [49] or by using a calcium source encapsulated in liposomes, then triggering the release using an external stimulus such as temperature or light [50–52].

5 Enzymatically Controlled GM Sequences It is possible to engineer alginates with specific MG sequences to tailor physical properties. The precursor that cells produce during the biosynthesis of alginate is poly-β-D-mannuronate, from which some of the residues are epimerised at C (5) enzymatically by mannuronan C(5) epimerases. This results in the conversion of β-D-mannuronate to α-L-guluronate and ultimately produces alginates with the characteristic block co-polymer structure containing various MG sequences. Genome studies performed on the alginate producing bacterium Azotobacter vinelandii has shown that this species produces seven different mannuronan epimerases that function during the biosynthesis of alginate. Using recombinant DNA, these epimerases known as AlgE1 to AlgE7 can be produced from Escherichia coli on a large scale. All these epimerases (AlgE1-AlgE7) basically perform the same function, which is epimerisation of mannuronate at C(5) to produce guluronate. Used collectively, however, these enzymes can produce alginates in vitro with vastly different sequences and consequently, different physical properties. For instance, AlgE4 converts mannuronate to guluronate in alternate residues producing MGM blocks (Fig. 7) [53]. Alginates treated with this enzyme have been shown to exhibit slower gelation kinetics and improved solubility at pH < 3 compared with native alginates. This is thought to occur due to the alternating axial–equatorial/equatorial– axial glycosidic linkages, which causes a delay in the sol-gel transition compared to the di-equatorial (poly-M) and di-axial (poly-G) sequences [54]. In another study, Ca-saturated hydrogel cylinders of poly MG produced using AlgE4 had extremely high syneresis and a rupture strength threefold that of native alginate from Laminaria hyperborea [55]. The epimerases AlgE1 and AlgE6 produce polyguluronate blocks. In particular, AlgE1 forms long homopolymeric G-blocks in excess of 50 residues, while AlgE6

Fig. 7 Conversion of mannuronate to guluronate in alternate residues by C5 epimerase and Alg4 producing MGM blocks (adapted from [53])

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A. M. Smith and J. J. Senior

produces shorter G-blocks with a broad block size distribution [56]. This combination of enzymes has been used to produce alginates with very high G content (>97% G) [56] which form highly stable calcium crosslinked gels with a high gel strength that is less elastic due to a tighter, more brittle network [55]. The ability to tailor alginate chemistry and subsequently manipulate the mechanical behaviour of alginate hydrogels, by using epimerases to precisely control the monomer sequence, provides scope for the development of 3D cell substrates that are intelligently designed for the desired tissue type or application.

6 Alginate Degradation A particular physical property of alginate hydrogels, which needs to be considered prior to use, is the degradation behaviour, which can be advantageous or disadvantageous depending on the application. For example, rapid hydrogel degradation would be an attractive property when applied as a cell delivery system, whereas rapid hydrogel degradation would be problematic in applications where mechanical strength is required over a long period of time. Alginate hydrogels are known to degrade over time in physiological conditions and therefore have limited long-term physiological stability. This degradation occurs via dissipation of the divalent cation crosslinkers as a result of ion exchange when exposed to monovalent cations (such as sodium and potassium). Chelation of divalent cations by phosphate ions can also occur, all of which are present in cell culture media and in vivo environments [27, 57–60]. Studies measuring the changes in mechanical properties of alginate hydrogels in cell culture conditions have been performed [61, 62], which showed that degradation of the hydrogel was greatest within the first 7 days of exposure to cell culture conditions. A sharp release of calcium ions from the hydrogel to the surrounding cell culture media was also detected in this period. When degradation rates were compared with other similar biopolymer hydrogels (in this case gellan gum and low methoxyl pectin), the alginate gels showed the most dramatic reduction in mechanical properties [62]. When implanted in the body, alginate polymer chains are non-biodegradable due to the absence of alginase. The gels however will dissolve into the surrounding milieu via ion exchange as they do within in vitro cell culture conditions, leaving the non-biodegradable alginate polymers. Perhaps the most eloquently designed method to retard degradation was reported by Birdi et al. [63] who illustrated how alginate hydrogel degradation can be tailored through the addition of orthosilicic acid (OSA), even when immersed in a potent calcium chelator. This mechanism was attributed to a molecular interplay between the OSA, the alginate and the crosslinking agent Ca2+ (Fig. 8). Other methods to overcome these problems have generally focused on developing covalently crosslinked alginate hydrogels and chemical modification of the alginate chains.

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49

Fig. 8 Schematic diagram to demonstrate how the addition of OSA to alginate hydrogels strengthens the interaction between the calcium and alginate, holding it in the egg–box junction, even during incubation in EDTA (reproduced from [63] with permission)

7 Modified Alginates Many alginate derivatives have been developed with a view to modify the physical properties for different applications with varying levels of success. What is important to realise however is that chemical modification may indeed modify the physical behaviour of alginate, but in some cases, this modification may introduce toxicity to the polymer; whether that is from the modified alginate itself or from residual chemicals used in the synthesis. This is of particular consideration when modifying alginate for cell culture and tissue engineering purposes.

7.1

Covalently Crosslinked Alginate

Covalently crosslinked alginate hydrogels have been widely investigated as a method to overcome some of the disadvantages associated with ionotropic alginate hydrogels and to improve their physical properties in an effort to widen and improve performance in cell culture and tissue engineering applications. The major difference in mechanical behaviour between ionically crosslinked gels and covalently

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Fig. 9 Covalent crosslinking of alginate using adipic acid dihydrazide (AAD) (adapted from [65])

crosslinked gels is that when stress is imparted upon ionic gels, the stress causes the dissociation of the crosslinks and loss of water, resulting in plastic deformation. In covalently crosslinked gels, the stress is unable to dissociate the covalent bonds and although water is also lost causing stress relaxation, elastic deformation occurs. This is elegantly demonstrated by Zhao et al. [64] who used adipic acid dihydrazide (AAD) to produce covalently crosslinked alginate gels (Fig. 9) designed to have the same modulus as a calcium crosslinked gel prepared by internal gelation and showed that ionic crosslinks relax stress much more rapidly than a gel with covalent crosslinks. It has also been demonstrated that the physical and swelling properties of alginate hydrogels can be tailored by using various crosslinking molecules, and importantly, by controlling the crosslinking densities [66]. Alginates covalently crosslinked using poly (ethylene glycol)-diamines of various molecular weights have been used to create gels with wide ranges of mechanical behaviour. It was shown that the elastic modulus initially increased with an increased crosslinking density, measured by the weight fraction of poly (ethylene glycol) in the gel. Interestingly, the increase in elastic modulus reduced once the crosslinking density increased beyond 27% [67], further demonstrating the potential control over the mechanical properties using covalently crosslinked alginate. The use of chemical crosslinkers with multiple attachment points has been reported to impart greater control over mechanical properties and degradation rates compared with bi-functional crosslinkers. This has been shown in isolated α-Lguluronate residues that were oxidised to prepare poly(aldehyde guluronate) (PAG). The mechanical stiffness and degradation behaviour of PAG gels when crosslinked using a multifunctional crosslinker poly (acrylamide-co-hydrazide) (PAH) was compared with a bi-functional crosslinker – adipic acid dihydrazide (AAD). The results revealed that the PAG gels crosslinked with PAH had a greater mechanical stiffness and degraded much more slowly than PAG crosslinked with AAD. It is thought that these differences in physical properties are a direct result of the multifunctional crosslinker having a greater number of attachment points in the gel [68, 69]. The ability to crosslink alginate on demand, as a response to external stimuli, is an attractive proposition for applications such as in situ gelation and rapid prototyping. One way to achieve this is to use a photosensitive covalent crosslinker. Despite the

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toxicity due to free radical production of some photo initiating agents, photo crosslinking can be performed using relatively mild reaction conditions and has been shown to be a safe procedure even in the presence of cells, as long as the appropriate chemical initiators are chosen. Mu et al. [70] prepared alginate modified with methacrylate and crosslinked by irradiation for 60 s using a halogen light source (400–520 nm; 1,000 mW/cm2) in the presence of 0.5% eosin (in 1-vinyl-2pyrrolidinone) and triethanolamine. The authors also developed a method to produce cell-loaded microcapsules from the methacrylated alginate that were monodisperse in size when crosslinked by visible light. This was achieved by suspending cells in the pre-crosslinked mixture, which was then extruded through a 26-gauge needle into a co-flowing stream of immiscible liquid paraffin extruded through a 22-gauge needle. The micro-particulate polymer solution was then collected and irradiated for 5 min using an LED light source (524 nm), which crosslinked the alginate, encapsulating the cells. The encapsulated cells exhibited no apparent toxic effects and remained viable for up to 2 weeks. Reports of using a dual crosslinking technique by combining ionic and covalent gelation methods have been described for the delivery of islets as a treatment for type I diabetes [71]. Modification of sodium alginate with 2-aminoethyl methacrylate hydrochloride (AEMA) introduced photoactivatable groups for the generation of covalent bonds. Solutions of this material were ionically crosslinked into alginate beads and then exposed to UV light to form dual crosslinked alginate structures. Mouse insulinoma cells (MIN6) were encapsulated within these dual crosslinked alginate beads and implanted in vivo within the greater omentum of rats. The dual crosslinked structures were shown to be stable under inflammatory challenge up to 3 weeks and capsule breakdown was prevented, whereas alginate/cell implants that were only ionically crosslinked failed within 1 week. Although there are many reported uses of methacrylate photo crosslinking in the presence of cells, which do not appear to cause any dramatic cell death, the free radical-producing photo crosslinking reactions could still elicit potentially harmful effects. Another photo crosslinker, polyallylamine, is an alternative approach to methacrylate photo crosslinking. This has been used to produce photo-crosslinkable alginate that provides an adequate environment for cell growth [72]. One disadvantage of this system is that it uses UV light exposure at about 330 nm, which is not ideal as light at this wavelength has the potential to damage the cultured cell’s DNA.

7.2

Amphiphilic Alginate

Several alginate derivatives have been investigated in a range of biomedical applications and amphiphilic alginates have shown particular promise. Amphiphilic alginate derivatives have been produced synthetically by the addition of hydrophobic moieties (e.g. alkyl groups) to the alginate chain. The addition of these hydrophobic regions enables the alginate to form self-assembled structures that include aqueous gels and have potential for use in biomedical applications.

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Amphiphilic sodium alginate was investigated by Pelletier et al. [73] and was prepared by esterification with alkyl chains of different lengths (dodecyl, octadecyl) to the alginate backbone. Rheological properties of these derivatives in aqueous solutions were shown to be typical of physically crosslinked, weak gel networks, which could have applications in cartilage regeneration. Leonard et al. [74] used these long alkyl chain derivatised alginates to form microparticles as a method to encapsulate and control the release of proteins. In addition to high encapsulation yields, no release of model proteins was observed in water for several days. This was in contrast with protein-loaded calcium alginate microparticles, which exhibited a significant release of protein within only a few hours. Release of the proteins could be triggered however, by the addition of either surfactants or lipases that disrupt intermolecular hydrophobic junctions and cleave the alkyl chains from the alginate backbone respectively, resulting in dissociation of the gel in both cases. To achieve longer-term stability, dodecylamine has also been conjugated to the alginate backbone through an amide linkage. This alginate derivative produced hydrogels that had long-term stability, showing a reduction in degradation in vitro at 37  C over a two-month period when compared with alginate ester derivatives. This increase in hydrogel stability was explained by the greater susceptibility to hydrolysis of the ester derivatives compared with the amide-linked derivatives [75]. Butyl esters of alginate have also been produced from sodium alginate through esterification with butanol in the presence of concentrated sulphuric acid as a catalyst. This butyl alginate ester was shown to encapsulate both hydrophilic and hydrophobic molecules without altering the gelling and non-toxic properties of native alginate [76]. Amphiphilic alginates containing cholesterol have also been synthesised via a reaction between the carboxylic acid groups of the alginate and the hydroxyl of the cholesterol, using N,N0 -dicyclohexylcarbodiimide as a coupling agent and 4-(N,N0 -dimethylamino) pyridine as a catalyst at room temperature. This cholesterol esterified alginate could self-assemble into nano-aggregates in NaCl solutions with a mean diameter of 136 nm, which was in contrast to the native sodium alginate, which formed aggregates in the region of 30 μm in diameter [77]. Other biomaterials have also been used to graft upon alginate. Colinet et al. [78, 79] investigated grafting poly (ε-caprolactone) (PCL) of different lengths to the carboxyl group of the alginate. This resulted in associative behaviour of alginate in aqueous sodium chloride solution that varied according to the length of PCL chains. This was also reflected in bulk rheological properties that were dependent on PCL chain length, ionic strength of the media, and % of PCL substitution [79]. Although not extensively used in tissue engineering, propylene glycol alginate (PGA), first prepared by Steiner [80], is the most widely used commercially available modified alginate. It is produced by partial esterification of the carboxyl groups on the uronic acid residues in a reaction with propylene oxide [81]. The major difference in material properties of PGA compared with native alginate is observed at low pH where it is soluble. Moreover, PGA also has a low sensitivity to divalent cations [82]. This has led to the industrial application of PGA alginate as a thickener and stabiliser; however, the hydrophobic propylene glycol groups facilitate its use as an emulsifier [83]. Polypropylene glycol chains have recently been investigated to

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Fig. 10 Synthesis of hydrophobically-modified alginate using dodecyl glycidyl ether (DGE) (adapted from [85])

control hydrophobicity in methacrylic alginate in a study of the effects of mineralisation in hydrogels. These findings revealed that an increase in concentration of hydrophobic propylene glycol moieties caused the hydrogel to be more kinetically favourable to mineral deposition [84]. More recently, amphiphilic alginates modified with dodecyl glycidyl ether (DGE) in a nucleophilic substitution reaction have been employed as a drug delivery system for water-insoluble drugs (Fig. 10) [85]. The solubility of model drugs clofazimine and amphotericin B increased up to 1,100 and 160 times, respectively, whereas the presence of sodium chloride in polymer solution decreased drug solubility due to charge screening and weakening the drug-polymer electrostatic interactions.

7.3

Oxidised Alginate

The use of oxidised alginate has recently become an attractive material for biomedical applications as properties such as swelling behaviour, degradation profile, and stiffness of the resultant gel can be controlled depending on the degree of oxidation. In particular, oxidised alginate has been used in developing wound healing dressings [86–88]. Moreover, it has also shown to be compatible for the culture of a range of cell types that include corneal endothelial cells [89, 90], hepatocytes [91], fibroblasts [87, 88], mesenchymal stem cells [92], adipose stem cells [93], and human osteoprogenitors [94]. The partial oxidation of alginate is a relatively simple approach to introduce a predictable degradation profile in physiological conditions, which can be useful for the controlled release of cells, growth factors, and drugs. The most common method to oxidise alginate is by using sodium periodate (Fig. 11). The periodate oxidation reaction occurs on the OH groups at C-2 and C-3 positions of the uronic acid residues which cleaves the carbon–carbon bond opening

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Fig. 11 Reaction scheme for the oxidisation of alginate by sodium periodate (adapted from [90])

the ring structure and facilitates degradation of the alginate backbone. In addition, this C-2 and C-3 cleavage results in the formation of two aldehyde groups giving the alginate chain two new reactive groups for each oxidised monomer. Partial oxidation of alginate does not prevent the ionotropic gel forming capability [95], as there are no changes to the carboxylates that are involved in the formation of calciummediated junction zones. However, Gomez et al. [96] have shown that this is only the case when the oxidation is below 10 mol% even in excess calcium. Oxidised alginates have also been used in conjunction with other chemical modifications to enable controlled degradation and additional functionality. In particular, methacrylated alginate has been of interest as a photo-crosslinkable alginate, which can have a tailored degradation rate by subjecting the material to partial oxidation [97, 98].

7.4

Peptide Modified Alginate

As is the case with other anionic polysaccharides, the inherent non-cell-adhesive nature of alginate has limited its use in cell culture and tissue engineering applications [99]. Indeed, it has been recently shown that cells can be immobilised within alginate gels for time periods in excess of 150 days [100], however, they enter a state of inhibited mitotic activity [61]. This is likely due to the lack of cell interactions with the alginate, yet cell recognition peptides are abundant in vivo in native extracellular matrix. This intrinsic non-cell-adhesive behaviour can be addressed by covalently bonding the fibronectin and laminin derived Arg-Gly-Asp (RGD) integrin recognition peptide to the alginate chains to promote better cell attachment (Fig. 12). The preparation of these alginate derivatives involves chemically bonding peptides as side-chains, coupled via the carboxylic acid groups on the individual sugar residues. In particular, peptides with this RGD sequence have been extensively used as adhesion ligands in tissue engineering. This is due to the relative abundance of integrin receptors for this peptide sequence on various cell types [102]. RGD modification of alginate was pioneered by Rowley et al. [103], who showed that RGD peptides can be chemically grafted to the carboxylate groups of the alginate by using aqueous carbodiimide chemistry. To achieve cell adhesion to peptide modified alginate and subsequent cell proliferation, a minimum concentration of peptides is needed on the alginate chains, along with cellular affinity to the RGD peptide. For

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Fig. 12 Comparison of human umbilical vein endothelial cells (HUVECs) cells seeded within (a) unmodified and (b) RGD-modified alginate hydrogels (adapted from Bidarra et al. [101])

instance, for MC3T3 osteoblast-like cells, substantial cell attachment occurs at ~12.6 μM of RGD per gram alginate [68, 69, 104]. This concentration, however, was not sufficient to stimulate proliferation in human bone marrow stromal cells [105]. Cyclic RGD peptides, however, have demonstrated the ability to enhance cell attachment and subsequent cell activity compared with linear peptides [106, 107]. When alginate hydrogels were used that contained cyclic RGD rather than linear RGD, hBMSC showed a threefold increase in proliferation. Furthermore, the cyclic RGD peptides appeared to enhance osteogenic differentiation when compared with linear RGD-alginate [105]. Following incorporation of RGD complexes within alginate solutions, cells suspended within the modified matrix are capable of effectively crosslinking the alginate network, assuming a uniform dispersion. This cell-initiated crosslinking is shear-reversible, ascribed by the weak and reversible ligand–receptor interactions within the system, making it an attractive characteristic for use in injectable cell delivery therapeutics [108]. Several other peptide sequences derived from other extracellular matrix proteins have also been investigated for particular cell types. For example, the peptide Tyr-Ile-Gly-Ser-Arg (YIGSR) has been grafted onto alginate to facilitate neurite outgrowth [109]. Cell recognition peptides have also been attached to alginates that have been methacrylated and oxidised to functionalise the alginate hydrogels to be cell adherent, photo-crosslinkable, and with a tailored degradation profile [110].

8 Outlook and Future Directions Alginate has continued to prove to be one of the most versatile biopolymers used in biomedical applications due to the greater understanding of alginate chemistry, which has progressed over the last century. From being originally thought of as an inert thickener or gel forming agent used in foods, alginate is now regarded as a

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smart biomaterial with high value functional properties that can be tailored either by simply understanding the properties of native alginates or by modifications to its chemistry to produce alginate derivatives. With the developments in the tunability of alginates discussed in this chapter, it is now becoming possible to intelligently design the microstructure of alginate hydrogels imparting regional variations within their structure, subsequently enabling an increasing level of control on the resulting physical properties of alginate hydrogels. It is likely that alginate will continue to be a useful tool in biomedical research, not only in the native form, but also as chemically modified alginate and in combination with other biopolymers/biomaterials as the need for tissue replacement and regeneration increases over the next decade.

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Adv Biochem Eng Biotechnol (2021) 178: 63–98 https://doi.org/10.1007/10_2021_167 © Springer Nature Switzerland AG 2021 Published online: 16 April 2021

Tunable Protein Hydrogels: Present State and Emerging Development J. Nie, X. Zhang, W. Wang, J. Ren, and A.-P. Zeng

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Building Blocks of Protein Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Natural Structural Proteins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 The Emerging of Catalytic Protein Hydrogel . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Engineered Building Blocks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Responsive Blocks of Protein Hydrogels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Temperature Responders . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Light Responders . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 pH Responders . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

J. Nie and X. Zhang contributed equally to this work. J. Nie and X. Zhang Beijing Advanced Innovation Center for Soft Matter Science and Engineering, Beijing University of Chemical Technology, Beijing, China W. Wang Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology, Hamburg, Germany J. Ren Beijing Advanced Innovation Center for Soft Matter Science and Engineering, Beijing University of Chemical Technology, Beijing, China State Key Laboratory for Biology of Plant Diseases and Insect Pests/Key Laboratory of Control of Biological Hazard Factors (Plant Origin) for Agri-product Quality and Safety, Ministry of Agriculture, Institute of Plant Protection, Chinese Academy of Agricultural Sciences, Beijing, China A.-P. Zeng (*) Beijing Advanced Innovation Center for Soft Matter Science and Engineering, Beijing University of Chemical Technology, Beijing, China Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology, Hamburg, Germany e-mail: [email protected]

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4 Summary and Outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 86 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 87

Abstract In recent years, protein and peptide-based hydrogels have received great attention for applications in biomedicine. Compared to hydrogels based on synthetic materials, they have the decisive advantages of being biological origin, providing cells with a more in vivo-like microenvironment and possessing potential biological activity. Empowered by the steadily deepened understanding of the sequencestructure-function relationship of natural proteins and the rapid development of molecular-biological tools for accurate protein sequence editing, researchers have developed a series of recombinant proteins as building blocks and responsive blocks to design novel functional hydrogels. The use of multi-block design further expands the customizability of protein hydrogels. With the improvement of standardization of preparation and testing methods, protein hydrogels are expected to be widely used in medical treatment, skin care, artificial organs and wearable electronic devices. More recently, the emergence of catalytically active protein hydrogel brings new opportunities for applications of protein hydrogels. It is believed that through integrated approaches of engineering biology and materials sciences novel and hereto unthinkable protein hydrogels and properties may be generated for applications in areas beyond medicine and health, including biotechnology, food and agriculture, and even energy.

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Graphical Abstract

Keywords Biomedical materials, Catalytic hydrogels, Protein hydrogels, Tunable properties

1 Introduction Hydrogels are water-containing soft materials comprised of three-dimensional (3D) networks of polymers, which are hydrophilic but insoluble because of the cross-linking interactions among the constituents [1]. Hydrogels cross-linked by non-covalent interactions such as hydrogen bonding, hydrophobic interaction, and electrostatic interaction are called physical hydrogels, which generally form reversible and dynamic cross-linking networks. In contrast, chemical hydrogels cross-

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linked by covalent bonds build up irreversible network structures showing higher mechanical strength. Because of their unique properties hydrogels are widely used in various fields, such as food [2], agriculture [3, 4], and medicine [5, 6]. Most biomolecules and cells perform their roles in aqueous solutions. Hydrophilic network in a hydrogel can maintain and support an aqueous environment that mimics the natural cytoplasmic or extracellular matrix required for biomolecules and cells to perform their functions. For example, the porous microenvironment in hydrogels enables controlled diffusion of encapsulated molecules, which is the basis of using hydrogels for drug delivery [7]. Protein hydrogels are hydrogels based on using proteins as building blocks. They have been an important topic in biomedical material research for many years (Table 1). Especially in recent years, the impressive advances in the application of protein hydrogels in regenerative medicine [33–35], 3D-printed tissues [36, 37], and flexible electronics [38, 39] have received considerable attention. Most of the proteins used in hydrogels are structural proteins. People have long extracted collagen and gelatin from animal skins for food and medicine applications. The modern gelatin industry has gradually established and matured in the past century. Other structural proteins such as elastin and silk protein have also been used in hydrogels in the past decades. Recently, non-structural proteins such as enzymes [40, 41] and receptors [34, 42] are explored to endow hydrogels with diverse biological functions and responsiveness. Protein hydrogels have shown many distinctive advantages which is unmatched by classical synthetic polymers. First, proteins as building blocks are renewable “green” materials that can be obtained from biological resources without relying on petroleum resources. Second, proteins are amino acid polymers with well-defined amino acid sequences. Genetic tools make it possible to fine-tune the properties of protein hydrogels by making changes at the primary sequence level. With the popularization of solid-phase peptide synthesis technology and the progress in synthetic biology, it becomes convenient to customize and produce engineered proteins for various applications [43, 44]. Last but not least, most protein hydrogels are inherently non-cytotoxic and biodegradable, which are prerequisites for medical use. Several commercial protein hydrogels have already been approved for medical applications [45]. Despite promising success in tissue engineering [46] and drug delivery [47], applications of protein-based hydrogels still face some challenges. The protein hydrogel used in vivo must be able to adapt to the highly dynamic physiological environment [48] which may significantly affect its functionality. More and more studies have emphasized the importance of the tunability of the various properties of hydrogels, such as mechanical properties, degradation rate, critical temperature, for their use in vivo [10, 49]. Furthermore, it is observed that the requirements on biocompatible materials have gradually shifted from non-toxic and inert to being also biologically active [50]. In this context, discovering protein hydrogels with catalytic activity or for biological sensing receives increasing interest in protein hydrogel research [41]. Based on improved understanding of protein sequence-structure-function relationships, a variety of natural and recombinant proteins with tunable properties have

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Table 1 Tunable hydrogels developed based on natural structural proteins Hydrogel types Collagen hydrogel Collagen hydrogel

Photo-crosslinked collagen hydrogel Imine-crosslinked collagen hydrogel Photo-crosslinked silk fibroin hydrogel Silk fibroin hydrogel Silk fibroin hydrogel

Tunable properties Strain stiffening Compression modulus, pore and fiber diameter, and diffusivity Mechanical strength Stress relaxation Mechanical strength Mechanical strength

Photocurable silk fibroin hydrogel

Mechanical strength and transmittance Pore diameter and compressive strength

Silk fibroin/collagen hydrogel

Mechanical strength

Elastin-like polypeptides hydrogel Metal ion coordinationcrosslinked elastin-like polypeptides hydrogel Cysteine-crosslinked elastinlike polypeptides hydrogel Enzymatically cross-linked silk-elastin-like proteins hydrogel Elastin-like polypeptides hydrogel Graphene/resilin hydrogel

Mechanical strength and degradation rate in vivo Mechanical strength

Resilin-like polypeptides hydrogel Resilin-like polypeptides hydrogel Photo-crosslinked resilin/silk fibroin hydrogel Resilin-silk-collagen copolymers hydrogel Keratin hydrogel Keratin hydrogel

Viscoelastic properties Mechanical strength

Mechanical strength Mechanical strength, stiffness, and electroconductive Mechanical strength Viscoelastic properties Elasticity and strength Viscoelastic properties Crosslink density and degradation rate in vivo Viscoelastic properties

Potential applications Tissue engineering Tissue engineering

Reference [8] [9]

Regenerative medicine Regenerative medicine Regenerative medicine Artificial organs/ tissues and catalyst carriers Artificial organs/ tissues Extracellular matrix for fibroblasts Cell delivery and tissue engineering Tissue engineering

[17]

Tissue engineering

[18, 19]

Injectable biomaterial Tissue engineering

[20]

3D cell culture and tissue engineering Wearable sensors

[22]

Regenerative medicine Injectable biomaterial Tissue engineering Drug delivery and tissue engineering Drug delivery and tissue engineering Tissue engineering

[10] [11] [12] [13]

[14] [15]

[16]

[21]

[23] [24] [25] [26] [27] [28, 29] [30] (continued)

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Table 1 (continued) Hydrogel types Keratin hydrogel Alkylated keratin hydrogel

Tunable properties Mechanical strength and transmittance Degradation rate and diffusivity

Potential applications 3D cell culture and tissue engineering Drug delivery

Reference [31] [32]

been developed for use in the design of hydrogels. In this review, we classify these proteins into building blocks and responsive blocks according to their basic roles in the formation and functionality of cross-linked networks of stimuli-responsive protein hydrogels. From an engineering perspective, we highlight the latest developments in tunable protein hydrogels, and how the two major components – building blocks and response blocks – in protein hydrogels address critical needs through their tunability. New development trends are also discussed.

2 Building Blocks of Protein Hydrogels 2.1

Natural Structural Proteins

Hydrogel materials based on natural structural proteins have been extensively studied and experienced rapid development in the past two decades (Fig. 1), thanks to the profound understanding of the sequence and structure of these proteins. In addition to the well-studied collagen hydrogel, silk, elastin and keratin hydrogels have received more and more attention. A large number of biomedical experiments and applications have proved the outstanding biocompatibility and degradability of these natural structural proteins as building blocks of hydrogels. Another attractive feature is their tunable mechanical properties, which are critical in the design of scaffolds for tissue engineering drug delivery systems and artificial extracellular matrices. Advances in synthetic biology have enabled recombinant structural proteins to show great potential as smart biomaterials and provide a powerful reference for the de novo design of protein hydrogels. An overview of some tunable hydrogels made of natural structural proteins and their engineered variants is given in Table 1. It can be seen that so far primarily mechanical and viscoelastic properties have been the objectives of the investigation. Biological functions of the hydrogels have been seldom addressed except for the degradation rate of some of the modified hydrogels.

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Fig. 1 Annual publications on hydrogels based on natural structural proteins from 2000 to 2020. The search engine used is “Web of Science,” and searching is done by using the terms “collagen hydrogel,” “silk hydrogel,” “elastin hydrogel,” “resilin hydrogel,” and “keratin hydrogel,” respectively

2.1.1

Collagen

As the most common natural structural proteins, collagens are widely used in food [51], medicine [52, 53], and cosmetic [54] products for decades. Collagens can be obtained economically in large quantities from wastes (such as skin, bone, cartilage, and scales) generated in the processing of fishery and animal husbandry products [55]. So far, there are at least 29 types of collagen uncovered and classified according to their homology and functions [56]. All of them have a characteristic right-handed triple helix structure assembled by three α-chains [57, 58]. Each α-chain is composed of the amino acid repeating sequence [G-X-Y]n, and the X and Y positions are usually occupied by proline and its hydroxylated form, hydroxyproline, respectively [59]. Intramolecular hydrogen bonds stabilize the triple helix structure of collagen, and the small steric hindrance of glycine ensures the tightness of assembly (Fig. 2a). Type I collagen is the most abundant and best studied collagen type, which comprises 90% of the protein in human connective tissues [61]. With excellent tensile stiffness, type I collagen is mainly responsible for providing maintenance and support for tissues and organs and plays a key role in bone development [62]. Since collagen is widely present in tissues, collagen-based hydrogels are considered to be the most ideal biocompatible materials. In recent years, they have significant applications especially in the researches of disease models [63], regenerative medicine [64, 65], and artificial tissues/organs [36, 37]. There are various sources of collagens and different cross-linking methods employed [56] which can lead to huge differences in the various properties of the collagen hydrogels, so the hydrogel preparation method should be rationally designed according to the purpose [66]. By adjusting collagen concentration, polymerization pH and polymerization temperature, physical collagen hydrogels can be obtained with properties different in compression modulus, pore and fiber diameter,

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Fig. 2 Natural structural proteins used in protein hydrogels. (a) Side view and axial view of the characteristic triple helix of collagen. Glycine is marked in green, and hydrogen bonds are marked with dashed lines. The protein structure is from a collagen-like peptide (Protein databank ID: 1K6F). (b) A model for the silk fibroin architecture in spider silk. Hydrophobic β-sheet crystalline regions are linked by small hydrophilic linker segments. Modified from Gosline et al. [60]. (c) Elastin widely exists in animal connective tissues such as lungs, blood vessels, and tendons. Elastinlike peptides with a characteristic sequence of pentapeptides have a lower critical solution temperature (LCST). The entropy-driven transformation between the disordered conformation and the β-spiral conformation occurs near the LCST. Resilin is an elastic protein found in insects crosslinked by intermolecular di-tyrosine bridges in vivo. (d) Keratin is the main component of animal hair and has a characteristic cysteine-rich sequence. They can be divided into α- and β-keratin according to their arrangement structure. The structure of α-keratin is from Protein databank (PDB) 6EC0

or diffusion coefficient [9]. Cells or bioactive components can be incorporated directly into the hydrogel during the fabrication process. [8]. However, the fast biodegradation and low mechanical strength of such “untreated” collagen hydrogels

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cannot match the demand of in vivo applications in many cases, and this has been one of the crucial factors that limit their further use [67, 68]. Therefore, chemically cross-linked [10, 11, 69] and hybrid cross-linked [27, 70, 71] collagen hydrogels have become research trends and shown advantages in the tunability of mechanical properties. Although conventional chemical cross-linking agents (such as glutaraldehyde, isocyanates, and carbodiimides) have been widely used to improve the stability of collagen hydrogels and regulate the degradation rate, even low-concentration residues of these cross-linking agents can still cause cytotoxicity, calcification, or foreign body response [72, 73]. The development of non-toxic but efficient chemical cross-linking agents/methods remains challenging.

2.1.2

Silk Fibroin

Silks from silkworms and spiders are recognized as high-strength protein fibers, comparable to synthetic fibers manufactured by modern technology [74, 75]. Silk fibroin is the main component of silk, accounting for about 75% of the total mass [76]. Structural biology and genetic studies have uncovered that silk fibroin is a block copolymer rich in hydrophobic β-sheet crystalline regions linked by small hydrophilic linker segments or spacers (Fig. 2b) [77–79]. These crystalline regions are primarily composed of repeating sequence motifs, where short side-chain amino acids (such as glycine, alanine, serine, threonine, and valine) play a leading role in β-sheet formation [60, 80–82]. The highly oriented crystalline phase along the fiber axis imparts high strength to silk, and the amorphous chains of silk can absorb most of the energy under stress, thus, leading to high toughness [83]. Till now, reconstituted silk fibroin polymers have insufficient mechanical properties compared with native fibers produced by insect silk-spinning organ [75]. As the water environment in silk fibroin hydrogel is very different from that in natural silk, optimizing the silk fibroin assembly process and developing better cross-linking method to obtain hydrogels with ideal performance are of considerable interest [15, 84]. Li et al. [13] developed a simple method for preparing silk fibroin hydrogels with adjustable mechanical strength by adding different concentrations of surfactants to the silk fibroin aqueous solution and incubating at 60
C. The hydrogel prepared by this method has excellent tensile and compressive modulus and can be used as a catalyst carrier. Binary-solvent-induced conformation transition (BSICT) strategy has been tested for the preparation of silk fibroin hydrogel [14]. The introduction of hexafluoroisopropanol (HFIP) as a solvent promotes the formation of β-sheet structure, and the elastic modulus of the hydrogel after replacing the solvent with water can reach 6.5 MPa. Some chemical methods, such as ruthenium catalyzed photocrosslinking and active oxygen-induced crosslinking, were also used to prepare silk fibroin hydrogels with tunable mechanical properties [12, 15]. Silk fibroin has demonstrated excellent biocompatibility and degradability in decades of biomedical applications [85–88]. Some of the silk fibroin-based products have already passed biocompatibility and safety tests and are commercially available [89]. In recent

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years, silk fibroin gel has also attracted much attention as a substrate material for soft electronics [38, 90].

2.1.3

Elastin and Resilin

Elastic fibers are an essential component of the extracellular matrix responsible for the elasticity and resilience of tissues and organs that requires repeated distension and relaxation throughout a lifetime such as lungs, skin, ligaments, and other connective tissues [91]. Elastin is the main component of elastic fibers, and the primary structure of its monomeric precursor tropoelastin is a block copolymer composed of a hydrophobic domain and a hydrophilic domain responsible for cross-linking [92]. It has been reported that tropoelastin is the most elastic and distensible monomer protein known, extending to eight times its length and recoiling back without hysteresis [93]. This impressive elasticity is mainly contributed by the repeating sequence motifs -VPGXG- (where X can be any amino acid except proline) in the hydrophobic domain [94]. In the study of the polypeptides [VPGVG]n, a molecular mechanism of the transition between disordered conformation and ordered β-spiral has been revealed, which is regarded as a “entropy-driven molecular spring” (Fig. 2c) [95, 96]. Beyond its impressive mechanical properties, elastin-like peptides (ELPs) comprising a repeating pentapeptide sequence -VPGXG- have also been found to have lower critical solution temperatures (LCST) [95–97]. This means, ELPs exist in disordered conformations below their transition temperatures and form more ordered β-turns above their transition temperatures. This characteristic has led to the rapid development of LCST-type ELP hydrogels, which are widely used in tissue engineering and drug delivery applications [98–103]. Regarding the adjustable response temperature of ELPs, we will introduce in detail in Section 3. Enhancing the mechanical strength of elastin is critical for its application in tissue engineering. ELPs with cysteine as the fourth amino acid of the pentapeptide repeat -VPGXG- can form a hydrogel with tunable viscoelasticity, and the covalent crosslinking through cysteine sulfur bonds is very convenient and minimizes the use of toxic crosslinking agents [20]. Ghoorchian et al. [18] designed a multi-block peptide containing a hydrophilic sequence end that can be coordinated and cross-linked by metal ions. The ELP assembly can be further cross-linked after adding zinc ions to form a hydrogel with adjustable stiffness, self-healing capability, and fatigue resistance. The double network ELP hydrogel composed of ionic coordination crosslinking and covalent crosslinking exhibits extremely strong toughness, and its strength can be adjusted by the concentration of zinc ions [19]. Resilin is another member of the elastic protein family and is found in specialized regions of the cuticle of most insects [104, 105], which usually undergo long-term and high-frequency mechanical stretching and bending (such as the elastic wing hinges in locusts and sound-producing tymbal in cicada) [106]. The excellent toughness, extensibility, and resilience make resilin a promising biomaterial candidate [104, 107]. Since the first resilin sequence was identified in Drosophila

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melanogaster, there have been reports of putative resilin-like proteins in many other insects, and the repeating motifs rich in proline and glycine are a common feature found in resilin from different species (Fig. 2c) [108, 109]. A study on resilin structure and function identified a reversible β-turn transition in resilin-like peptides (RLPs), suggesting a molecular mechanism similar to that of elastin [110]. Based on the understanding of natural resilin, more and more biocompatible and external stimuli-responsive RLPs hydrogels have been developed and are expected to replace some traditional materials in tissue engineering and flexible electronic materials [23, 25, 111, 112]. Hydrogels with elastic moduli in the range of 1–25 kPa can be produced from recombinantly engineered RLPs [24]. Crosslinking agent such as hydroxymethyl phosphine (THP) can improve the strength of the RLP hydrogel; however, such crosslinking agents must be avoided in medical use because of their cytotoxicity [25]. RLP hydrogel prepared by ruthenium-mediated photo-crosslinking displays impressively enhanced modulus (about 78 MPa) and is shown to be non-cytotoxic when cultured with a mouse pre-chondrocyte cell line [113]. Bracalello et al. [27] designed an engineered polypeptide comprising resilin-, elastin-, and collagen-like sequence that can assemble to form a hydrogel with excellent elasticity. This model embodies the high flexibility and tunability of multi-block design and its potential as a general strategy in the design and engineering of protein hydrogels.

2.1.4

Keratin

Keratin, the major component of wool, hair, hooves, feathers, and horns, is one of the most abundant protein sources. Cysteine accounts for 7–20% of the total amino acids in keratin and functions in cross-linking each other within and between molecules, which is a feature distinguishing keratin from other proteins (Fig. 2d) [114]. The molecular structure of keratin from various sources is quite different and can be roughly divided into α-keratin and β-keratin according to their secondary structure [115]. The most commonly used keratin in biomaterials are α-keratins from human hair and wool, whose molecular units can self-assemble into coiled coils and higher-ordered nanostructures [116]. In recent years, keratin has been widely used in drug delivery applications. The process of cross-linking and in vivo degradation of keratin molecules mediated by the formation and reduction of disulfide bonds is convenient to regulate. This makes the drug release profile of the carrier adjustable to meet the specific drug release requirements [28–30, 32, 117]. Wang et al. [31] developed and optimized the procedure of producing human hair keratin-based hydrogels. The mechanical strength and light transmittance can be adjusted by the concentration of the keratin solution and the pH of inducing gelation. In addition, there are reports showing that the degradation rate can be adjusted by changing the ratio of oxidized and reduced keratin in the hydrogels [28, 29].

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The Emerging of Catalytic Protein Hydrogel

Proteins which have been found to form hydrogels as described above are all structural proteins. The properties which can be tuned are merely mechanical, viscoelastic, or physiochemical ones as summarized in Table 1. Recently, our group has discovered that the enzyme lipoate-protein ligase A (LplA) of E. coli can self-assemble to form a reversible hydrogel under non-denaturing conditions (Fig. 3) [41]. LpLA is responsible for the lipoylation of many important proteins (enzymes) such as the E2 subunit of pyruvate dehydrogenase and 2-ketoglutarate dehydrogenase complexes and the H-protein of the glycine cleavage system (GCS). GCS is involved in the one-carbon metabolism of all types of cells [118–122] and disease development including aging, obesity, and cancers [123–125]. According to our knowledge, this is the first report of a catalytic globular protein functioning as a building block of hydrogel. The self-assembly of LplA is sensitive to the solution environment (buffer, pH, and temperature) and has a tunable LCST in the range of 10–37
C. Tests of some mutants with mutations at key sites on the protein surface showed that intermolecular hydrogen bonding, electrostatic and hydrophobic interactions drive the selfassembly of LplA molecules. Preliminary studies showed that LplA remains its catalytic function to H-protein lipoylation of the glycine cleavage system at gel state. Thus, the unique catalytic activity of LplA hydrogel makes it an outstanding candidate as tissue regeneration material and biosensor in vivo. Considering that E. coli LplA is a powerful enzyme with broad substrate specificity [126], further researches will be carried out with focus on its modification and applications, such as the development of in vivo fluorescent labeling for in situ

Fig. 3 LplA from E. coli self-assembles to form a reversible hydrogel. The hydrogel exhibits tunable thermo- and pH-responsiveness, whose response ranges are close to physiological conditions. The protein structure is from PDB 3A7A

Tunable Protein Hydrogels: Present State and Emerging Development

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cellular protein imaging. LplA-based hydrogels offer the decisive advantages of being composed of a catalytic enzyme compared to all other known hydrogels and may therefore open up hitherto inaccessible applications. In tissue engineering, hydrogel scaffolds provide cells with a basic frame for attachment and microenvironment of cellular functions. Different from conventional inert biological materials, LplA hydrogel with one-carbon and energy metabolism-related catalytic activities is expected to have the ability to more effectively regulate cell proliferation and differentiation. In this regard, further research is needed to quantitatively study the effects of LplA on the growth and biochemical behavior of various tissue cells.

2.3

Engineered Building Blocks

Protein–protein and protein–ligand interactions are very common in nature, providing abundant building blocks for the design of engineered protein hydrogels. These proteins can be combined and cross-linked in specific covalent or non-covalent ways to drive the formation of gel networks which, compared with classic polymerization techniques, avoids the use of cytotoxic cross-linking agents and is more controllable, and therefore becoming more and more popular. From an engineering perspective, the ideal protein crosslinking component should be as small as possible, with good stability and sufficiently high binding affinity. Since the coiled-coil motif, the first universal cross-linking module, was used in the design of protein hydrogels, the number of available protein cross-linking components has gradually increased.

2.3.1

Self-Assembling Proteins/Peptides

Coiled-coil motif is one of the super secondary structures of proteins, which is ubiquitous in the native structure of those proteins accounting for 10% of the total protein in eukaryotes [127] (Fig. 4a). The canonical coiled coil contain a characteristic seven-residue sequence repeat, labeled from a to g, in which hydrophobic amino acids (usually Leu, Ile, and Val) at positions a and d are conserved [128–130]. The assembly of coiled coil is driven by hydrophobic interactions, stabilized by hydrogen bonding and electrostatic interaction, and reversibly responds to changes in temperature, pH, and ionic strength [131]. In the past two decades, coiled coils have been used as building blocks of hydrogels to develop multi-block protein hydrogels [132– 135] and hybrid hydrogels [136–139]. These engineered hydrogels are comprised of more than one coiled-coil domain within a polymeric chain, preserving each domain’s function and leading to an overall enhanced functionality of the multiblock system [140]. Their physicochemical and/or biological properties can be finetuned easily by modifying the assembly of the coiled coils (Fig. 4a). Typical examples of coiled-coil-assembled hydrogels can be found in the studies of Tirrell’s group. They are the first who tried to apply coiled-coil domains in the design of protein hydrogels and developed a triblock polypeptide comprising two

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Fig. 4 Self-assembling peptides that are used in design of tunable protein hydrogels. (a) Multiblock peptides composed of coiled coil can self-assemble into a hydrogel network with mechanical properties tunable through block number, linker length, aggregation number, and aggregate stability. (b) Amphiphilic peptides composed of alternating hydrophobic and hydrophilic amino acids can form a hydrogel in a suitable pH environment, and β-sheet is the most dominant conformation. (c) Caf1 has a prominent N-terminal sequence, which can specifically bind to another Caf1 molecule (stable at