Biofunctional Hydrogels for Three-Dimensional Stem Cell Culture

C H A P T E R

22 Biofunctional Hydrogels for Three-Dimensional Stem Cell Culture Jenna L. Wilson1, Todd C. McDevitt2 1

Georgia Tech Bioengineering, Atlanta, GA, United States; 2University of California at San Francisco (UCSF), San Francisco, CA, United States

O U T L I N E 1. Introduction 1.1 Applications of 3D Stem Cell Culture 1.1.1 Studying Stem Cell Behavior and Niche Interactions 1.1.2 Creating Tissue-Engineered Constructs 1.1.3 Improving Cell Engraftment and Survival In Vivo 1.1.4 Stem Cell Expansion and Biomanufacturing 1.1.5 Drug Discovery and Toxicity Testing

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2. Hydrogel Materials 2.1 Naturally Derived Materials 2.1.1 Protein-Based Materials 2.1.2 Polysaccharide Materials 2.2 Synthetic Materials 2.3 Hybrid Materials

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3. Hydrogel Material Properties 3.1 Degradation 3.2 Topography 3.3 Matrix Mechanics 3.4 Mass Transport Properties 3.5 Quantitative Properties of Hydrogels

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1. INTRODUCTION There are significant differences between the native environment of stem cell populations and the artificial in vitro culture platforms used by the majority of scientists today. For example, pluripotent stem cells (PSCs) Biology and Engineering of Stem Cell Niches http://dx.doi.org/10.1016/B978-0-12-802734-9.00022-6

4. Engineering Hydrogel Bioactivity 4.1 Cell Adhesion Motifs 4.2 Growth Factor Presentation 4.3 Stimuli-Responsive Materials 5. Technologies for Hydrogel Production and Assessment 5.1 Hydrogel Production Methods 5.1.1 Formation of Microparticles 5.1.2 Microencapsulation Methods 5.1.3 Microfabrication Techniques for Bulk Hydrogels 5.2 Translating 2D Assays to 3D Systems

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6. Challenges and Future Directions 6.1 Spatiotemporal Control Over Biophysical and Biochemical Hydrogel Properties 6.2 Promoting Endogenous Tissue Morphogenesis

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7. Conclusion

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Abbreviations and Acronyms

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References

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exist transiently in the preimplantation blastocyst constituting the inner cell mass, a tightly packed cluster of cells, which is encased in a layer of epithelial trophoblast and surrounded by the fluid of the blastocoel cavity.1 In contrast, bone marrow mesenchymal stem/ stromal cells (MSCs), which are multipotent cell

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Copyright © 2017 Elsevier Inc. All rights reserved.

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FIGURE 22.1 In a two-dimensional monolayer (left), stem cells are polarized by attachment to a stiff plastic surface on a single side and have limited interactions with other cells and 3D extracellular matrix components. In contrast, a 3D hydrogel environment (right) provides cues in three dimensions and enables cellecell and cellematrix interactions in multiple dimensions.

populations present throughout adult life, are encompassed in a semisolid matrix consisting of various collagens, proteoglycans (PGs), and glycosaminoglycans (GAGs).2,3 Despite these distinct yet similarly complex microenvironments, in vitro studies of embryonic stem cells (ESCs) and MSCs are generally conducted in the same manner: as two-dimensional (2D) cell monolayers on plastic surfaces. In typical 2D culture, cells are specifically polarized by attachment to a plastic surface on one side and contact with liquid culture medium on the other (Fig. 22.1). Interactions between neighboring cells occur laterally within a single plane, if at all, and the threedimensional (3D) physical cues presented by interstitial extracellular matrices are absent entirely. Because celle matrix and cellecell interactions dictate much of a cell’s behavior, it is logical that the attenuation of these cues can lead to cell dysfunction, and this has been observed in cell types ranging from liver hepatocytes to ovarian and colorectal cancer cells.4e6 Therefore, a greater emphasis on 3D culture platforms for stem cell culture may provide more accurate insights into in vivo cell physiological function and enable more advanced applications, such as tissue-engineered constructs, biomanufacturing approaches, and platforms for drug discovery and toxicity testing. There are several standard approaches for 3D culture of mammalian cells, including the formation of scaffoldfree microtissues (often in the form of spheroids or aggregates), which offer limited exogenous control over extracellular matrix (ECM) properties, or the seeding of cells onto porous scaffolds or decellularized ECM from native tissues.7 However, perhaps the most pervasive technique has been culture of cells within hydrogel scaffolds. Hydrogels, which are simply hydrated polymer networks, share many key physical properties with native tissues, including a high water content, a similar range of elasticities, and mass transport characteristics that permit the diffusion of external stimuli and internal autocrine and paracrine signals. Additionally, depending on the source or modification of the

hydrogel polymer, the gels can serve as a platform to integrate specific biological elements in a wellcontrolled manner and consequently offer superior control when compared with cell or tissue-derived ECM.8 Thus, this chapter will focus on the use of hydrogels as platforms for 3D culture of stem cell populations, with a specific emphasis on systems that provide biofunctionality as opposed to simple biocompatibility (Fig. 22.2).

1.1 Applications of 3D Stem Cell Culture 1.1.1 Studying Stem Cell Behavior and Niche Interactions The function of stem cell populations in vivo is still being actively investigated to provide insights into cell regulators, which can be translated to in vitro differentiation protocols, and using typical 2D culture methods may lead to an unintentional bias in the interpretations of stem cell phenotype and function. Unfortunately, performing studies of complex niche interactions in vivo is technically challenging because of limitations in imaging and real-time monitoring. Thus, engineering hydrogel properties to mimic physical and chemical elements of the native stem cell niche may enable the discovery of novel insights into cell physiology and interactions with the surrounding microenvironment, both in terms of matrix properties and the soluble milleu.9,10 Depending on the degree to which a stem cell niche has been previously characterized, hydrogels can be designed to specifically simulate known niche components. For example, culture of MSCs in hydrogels constructed with collagen, a major component of the bone marrow niche, enhanced MSC secretion of angiogenic factors (compared with 2D cultures) and augmented chondrogenic differentiation of MSCs, as evidenced by increased GAG and collagen production, when compared with noncollagen-based hydrogels.11,12 In addition to mimicking niche material composition,

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FIGURE 22.2 Culture of stem cells in 3D hydrogels enables many downstream applications, including investigation of the stem cell niche, the creation of tissue-engineered constructs, advanced drug screening platforms, biomanufacturing systems, and improved cell survival and function posttransplantation.

simulating the mechanical properties of the native niche can enhance stem cell function, as has been observed with skeletal muscle satellite cells cultured on substrates with mechanical properties similar to native muscle.13 In many cases though, specific niche components are not well defined and/or their functional effects are poorly understood, and thus design criteria for engineering hydrogel materials are lacking. Recent advances in material microarray technologies have attempted to address this dearth of information by allowing for highthroughput combinatorial analysis of cell response under many different conditions. In recent work seeking to investigate factors capable of promoting the self-renewal of murine ESCs, an array of more than 1000 variations of synthetic polyethylene glycol (PEG) hydrogels was created with varying mechanical properties (elastic modulus of 300e5400 Pa), sensitivity to matrix metalloproteinase (MMP) degradation, and inclusion of ECM (collagen, fibronectin, and laminin) and cellecell interacting (E-cadherin, Jagged, and epithelial cell adhesion molecule) proteins.14 The wide range of hydrogel properties in combination with differential cell densities and soluble factor [fibroblast growth factor 4 (FGF4), bone morphogenetic protein-4 (BMP-4), and leukemia inhibitory factor (LIF)] in addition provided a large data set in which systems biology approaches identified key regulators of ESC self-renewal. Similar combinatorial platforms of 3D microenvironments could be similarly applied to other

systems to determine the role of specific niche components in cell processes in a controlled manner. 1.1.2 Creating Tissue-Engineered Constructs Since the original principles of tissue engineering first emerged in the 1980s, the model of seeding cells onto/ within a scaffold has served as a paradigm for the entire field. However, the functional role of scaffold materials has shifted away from simple physical support structures and emerged as an essential contributor to the overall biophysical and biochemical properties of engineered tissues. The combination of stem cells or stem cellederived cell types with growth factors on a scaffold remains the iconic vision of a tissue-engineered product.15 However, tissue engineering and regenerative medicine applications have not yet attained the widespread clinical use that was predicted, primarily because of a continued lack of understanding about cellular processes in vitro and in vivo and the inability to recapitulate known processes in an ex vivo setting. Thus, it is critical for the field of regenerative medicine that 3D platforms for stem cell culture continue to be created and assessed as platforms for complex tissueengineered constructs. 1.1.3 Improving Cell Engraftment and Survival In Vivo While greater numbers of cell-based therapies demonstrate efficacy and thus seem to be a promising

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clinical approach, most are limited by the lack of long-term survival and functional engraftment of transplanted cell populations. For example, engraftment rates of cells delivered to the heart are typically below 10% and more often in the range of 1e2%, if not lower.8,16 Part of the lack of stable engraftment may be attributed to the dysfunctional environment the cells are transplanted into, as the injured tissue typically includes damaged matrix and tissue necrosis that can adversely affect stem cell fate and subsequent function.17 Hydrogel biomaterial carriers, which tend to be conveniently injectable, can be used to provide a controlled niche with instructive cues for transplanted stem cells (or stem cellederived cells) to improve their survival, engraftment, and long-term function. In addition to protecting transplanted cells, the hydrogel itself can promote tissue regeneration through the release of soluble cues or the recruitment of endogenous stem cell populations.17 Hydrogels composed of hyaluronic acid (HA), heparin, gelatin, and PEG are able to recruit bone marrow MSC populations when modified with hepatocyte growth factor as were gelatin hydrogels containing BMP-2 and stromal cellederived factor-1 (SDF-1).18,19 In this sense, incorporating cues within the material that target host cell populations could lead to synergistic regeneration with the transplanted stem cellederived populations and ultimately yield more effective cell therapy products. 1.1.4 Stem Cell Expansion and Biomanufacturing The clinical translation of autologous stem cell therapies is dependent on the development of biomanufacturing processes to expand stem cells to the appropriate numbers, often estimated at 107e1010 cells per patient depending on the application.20 Because of the amount of surface area required to culture such large quantities of cells in 2D, moving to 3D platforms is a necessity for economic reasons and for ease of monitoring because many online sensors for temperature, pH, dissolved oxygen, etc., are designed for use in suspension bioreactors. Encapsulation of stem cells within hydrogel microbeads or microcapsules can be used to provide a defined environment for adherent stem cell populations [i.e., ESCs, induced pluripotent stem cells (iPSCs), MSCs, neural stem cells (NSCs)] while enabling culture in large-scale bioreactor systems.21 Culture within hydrogels can also protect the enclosed cells from experiencing shear forces inherent to stirred bioreactor systems, which is particularly important in the case of stem cells because of their phenotypic sensitivity to hydrodynamic forces.22 1.1.5 Drug Discovery and Toxicity Testing The successful development of new pharmaceutical compounds and screening of existing chemical libraries

is highly dependent on the context in which cells are screened for toxicity and other cellular responses. Some of the reasons for the failure of late-stage drug candidates lies with inadequacies in the predictive ability of traditional 2D cell culture methods.23 The advent of iPSCs now allows for individual patient- and phenotype-specific drug assessments, and differentiating PSC populations into more mature tissue types of interest will enable a higher degree of functional testing than is currently possible.24 Specifically, drug testing may benefit from analyzing the response of functional “organoid” structures, such as intestinal organoids and cerebral brain-like organoids, which have been formed from PSC aggregates suspended in MatrigelÔ , a natural hydrogel product derived from mouse sarcoma cells and consisting of a mixture of laminin I, collagen IV, entactin, and various PGs and growth factors.25,26 To translate organoid cultures from the basic research environment to high-throughput industrial screening processes, it will be necessary to design well-defined hydrogel systems to ensure scalability and reproducible models.27

2. HYDROGEL MATERIALS As a class of materials, hydrogel polymers can vary greatly in source and structure, making them extremely versatile for applications in the 3D culture of stem cells. Although hydrogels can be classified in several ways, such as by their structural composition, charge, degradability, cross-linking chemistry, or responsiveness to different types of stimuli, distinguishing by them by material origin allows for distinction between naturally derived “promoting” polymers, which consist of some inherent bioactive components, and synthetic “permitting” polymers, which begin as more of a blank slate for directed feature design.7,28

2.1 Naturally Derived Materials Because hydrogels are often desired as ECM mimics, it is rational that researchers often turn to polymers consisting of natural ECM components or biomolecules with similar chemical structures. Naturally derived hydrogel materials are intrinsically biocompatible and bioactive, which often leads to high cell viability and proliferation rates without the need for significant material modification. However, the specific bioactive components have the potential to convolute results, as it can be challenging to isolate the effects of individual cues on cell fate.8 For example, increasing the mechanics of a collagen hydrogel will also lead to an increase in the concentration of adhesive sites. Additional drawbacks

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include the inherent batch-to-batch variability and potential economic cost that come with the isolation and purification of a naturally derived hydrogel material.29 Furthermore, it can be difficult to tune the system for the desired mechanical properties, degradability, and bioactivity because the molecules have their own intrinsic properties that are inherent to the polymer backbone and native cross-linking mechanisms. Naturally derived hydrogel polymers can be further subdivided into protein-based materials, such as collagen, fibrin, and MatrigelÔ , and polysaccharide-based materials, including agarose, alginate, chitosan, and GAGs, such as chondroitin sulfate, heparin/heparan sulfate, and HA. 2.1.1 Protein-Based Materials 2.1.1.1 Collagen Collagen is the most abundant protein found in mammals, and therefore, it has been extensively studied as a hydrogel for 3D cell culture applications. Specifically, collagen I, which contains several cell adhesion sequences including the integrin-binding sites GFOGER (glycinee phenylalanineehydroxyprolineeglycineeglutamatee arginine) and DGEA, is present in the highest abundance and is most commonly used in hydrogel applications. Collagen I is typically isolated from bovine skin or rat tail, solubilized in an acidic buffer (typically acetic acid), and formed into a hydrogel by neutralizing the acidic solution with sodium hydroxide and heating to 37
C to create a collagen matrix.29 The resultant gels are relatively soft, typically with an elastic modulus of less than 5 kPa and thus have limited use in applications that might require the use of stiffer hydrogel materials. As a result, collagen is often combined with other materials to form hybrid hydrogels, or alternatively, the collagen is modified to enable greater chemical crosslinking, such as with glutaraldehyde, formaldehyde, and succinic anhydride.30 Degradation of collagen is mediated by several classes of proteases, including MMPs and cathepsins, which are secreted by many cell types and can promote hydrogel degradation. MSCs are often encapsulated within collagen matrices because of the naturally occurring adhesive peptide sequences in the collagen chains, which lead to improved cell viability. Self-assembled microspheres of collagen and human MSCs have been formed by suspending the cells in liquid collagen solution and inducing gelation of the resulting droplets.31 The resulting constructs could be cultured in suspension, and the collagen matrix supported MSC viability, growth, and differentiation while improving the self-renewal capacity of the cells relative to monolayer controls. Encapsulation of embryoid bodies (EBs), cellular aggregates derived from ESCs, within collagen I hydrogels has

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also enabled the recapitulation of blood vessel formation and has been used as a model to understand the sprouting processes of angiogenesis.32 2.1.1.2 Fibrin The motivation for fibrin hydrogel formation can be traced to the natural process of coagulation in which blood is transformed from a liquid to a gel-like state in response to an injury that entraps cells. After the accumulation of platelets at the site of injury, thrombin cleaves soluble fibrinogen molecules into fibrin monomers, which in turn cross-link to form a dense fibrous matrix. The resulting fibrin gel plays a significant role in the healing and regeneration of the damaged tissue and is able to interact with the surrounding cells through binding domains for integrins, heparin, and fibrinogen, and many growth factors.33 To induce fibrin hydrogel formation ex vivo, a chilled fibrinogen solution (typically 2e4 mg/mL) is combined with thrombin and cross-linked by raising the temperature to 37
C. By varying the thrombin concentration from 0.001 to 1 units of thrombin/mg of fibrinogen, a range of mechanical properties can be achieved, with the elastic modulus generally ranging from 1 to 30 kPa.29,34 In the presence of cell-secreted enzymes such as plasmin, fibrin gels begin to exhibit decreased mechanical stability and eventually degrade completely, though the inclusion of enzymatic inhibitors such as ε-aminocaproic acid can be used to slow degradation.35 Thus, alternative cross-linking chemistries, such as with formaldehyde, gluteraldehyde, and succinic anhydride, have been similarly investigated for fibrin as they have for collagen.29 As with collagen, fibrin has also been commonly used as a hydrogel scaffold for MSCs. Early studies identified appropriate fibrinogen and thrombin concentrations, which promote MSC proliferation and permitted injection in vivo, after which MSC migration out of the fibrin gel was observed.36 Fibrin gels have also been used as scaffolds for ESC culture, particularly for differentiation toward neural lineages in which EBs are cultured in fibrin gels and exposed to neurotrophic factors.37 2.1.1.3 Matrigel Unlike the single-protein composition of collagen and fibrin hydrogels, Matrigel contains a complex mixture of basement membrane proteins derived from mouse sarcoma tumors. In addition to the ECM proteins laminin I, collagen IV, entactin, and heparan sulfate proteoglyclans, a number of growth factors are also commonly present in Matrigel, including transforming growth factor-bs (TGF-bs), FGFs, epidermal growth factors, platelet-derived growth factors, and insulin-like growth factors.38 Although this complex mixture of growth factors has been observed to readily support the culture of many cell types, there is significant

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batch-to-batch variability of Matrigel and the number of factors present can obfuscate specific biological mechanisms. A growth factorereduced Matrigel was developed by using a precipitation step to remove many of the growth factors in an attempt to address some of these issues, though the molecular complexity of the hydrogel material remains.39,40 Matrigel is a thermoreversible hydrogel that exists as a liquid below 10
C and is prone to degradation by cell-secreted proteases (primarily MMPs) over time. Matrigel has been used extensively as a surface coating for 2D cultures of human PSCs as a replacement for stromal feeder layers.41 More recently, Matrigel has been used as an essential 3D matrix to permit the selforganization of stem cells into organoid structures. The formation of intestinal crypts from Lgr5þ intestinal cells was performed in Matrigel because of the similar presence of laminin in the native crypt.25 In the formation of cerebral organoids, neuroectoderm-biased EBs formed from human ESCs were embedded in Matrigel droplets and cultured in a stirred bioreactor system to induce neural tissues with features unique to specific regions of the brain.26 2.1.2 Polysaccharide Materials 2.1.2.1 Marine Origin Agarose, alginate, and chitosan are biocompatible linear polysaccharides of marine origin, with agarose obtained from red algae, alginate from brown seaweed, and chitosan from the shells of crustaceans. Agarose consists of b-D-galactopyranosyl and 3,6-anhydro-a-Lgalactopyranosyl units and exhibits a temperaturedependent gelation below a threshold temperature (usually 15e30
C, based on the extent of methylation of the source agarose). Agarose is relatively stable because mammalian cells do not produce enzymes capable of agarose degradation, but degradation is observed over long time periods (25% over the course of 60 days) because of interruption of hydrogen bonding.42 Alginate comprises two anionic monomers, a-L-guluronic acid and b-D-mannuronic acid, and is ionically cross-linked by divalent cations, typically calcium or barium.43 As with agarose, alginate is not directly degraded by cells; however, the material can break down because of chelation of the divalent cations or displacement of divalent cations with monovalent cations present in high concentration.44 Chitosan consists of b(1e4) linked glucosamine and N-acetyl glucosamine and is ionically crosslinked with divalent cations as described with alginate. Unlike agarose and alginate, chitosan can be degraded by mammalian cells that express lysozyme, an enzyme with a role in the innate immune system.45 All three polysaccharide species lack specific biological moieties; thus, they are often used in combination with other

biological materials or as blank slate materials in situations in which extracellular biological stimuli are not desired. Agarose has been used in several instances to encapsulate PSC aggregates to prevent agglomeration while promoting differentiation.46,47 Similarly, alginate has also been used to culture embryonic stem cells, typically in a microbead configuration conducive to long-term bioreactor culture.48 Though alginate is often thought of as relatively inert for a naturally derived material, differences in ESC and MSC differentiation have been observed as a result of varying the physical properties of alginate hydrogels.49,50 A hybrid chitosan-alginate porous scaffold has also been developed for feederfree expansion and maintenance of pluripotency of human ESCs for up to 21 days.51 2.1.2.2 Glycosaminoglycans Unlike the relatively inert quality of other polysaccharide species, GAGs play critical roles in many biological processes, including attachment to PGs within the ECM, recruitment of water into connective tissues to increase osmotic pressure and allow tissues to withstand compressive loading, lubrication of joint areas, and sequestration and immobilization of growth factors.52 There are several classes of GAG molecules, all of which are negatively charged linear polysaccharides with varying degrees of sulfation. HA is the only nonsulfated GAG species and has been the most extensively studied for biomaterial applications because of its ease of chemical modification because GAGs do not have an inherent cross-linking ability. Unlike most other ECM components, HA does not have integrin-binding sites for cell adhesion, though it can reportedly bind to specific cell receptors (e.g., CD44) and intercellular adhesion molecules (e.g., ICAM-1).29,53 Other GAG species, including heparin/heparan sulfate and chondroitin sulfate, have a higher net negative charge than HA and thus have an even higher affinity for positively charged growth factors.54 Hydrogels of GAG species are typically created as hybrid materials with synthetic polymers, though pure materials of chemically modified heparin and chondroitin species have also been developed.55e57 HA is found in the greatest abundance during early embryogenesis and thus has been investigated for ESC culture. Indeed, hydrogels of HA were found to promote the feeder-free self-renewal of human ESCs, and human ESCs were determined to possess active HA-binding sites and receptors.58 Additionally, HA hydrogels have been investigated for tissue-engineered skeletal muscle constructs using MSCs and enhance chondrogenic differentiation of MSC in comparison with synthetic hydrogels.59 Hydrogels modified with chondroitin sulfate (PEG-based) and heparin (collagen-based) have also

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been used to promote osteogenic and chondrogenic differentiation of MSCs.60,61

2.2 Synthetic Materials Many of the inherent limitations of hydrogels made of natural materials can be overcome with the use of synthetic polymers. With chemically defined materials, there is a priori control over cross-linker density and spacing to promote uniform quality with the desired mechanical properties. Although synthetic materials may not intrinsically offer biological information, such as adhesive cues or presentation of growth factors, to cells, bioactive and/or stimulus-responsive epitopes can be engineered into the polymers to confer biological signals.7 Many synthetic polymers have been investigated for cell encapsulation, including PEG, polyvinyl alcohol, and poly N-isopropylacrylamide (PNIPAAm). The most widely used polymer is PEG, which exhibits high hydrophilicity with low protein adsorption and can be readily modified at its hydroxyl ends.8 Cross-linking conditions that are noncytotoxic are possible with Michael-type addition reactions in which PEG modified with amine or thiol groups is mixed with methacrylicmodified PEG. Alternatively, photocross-linkable systems can be used in which acrylic or methacrylicmodified PEG is mixed with a biocompatible photoinitiator and exposed to a specific, noncytotoxic wavelength of ultraviolet (UV) light.29 The mechanical properties of PEG hydrogels are typically modulated by varying the concentration and molecular weight of the polymers; however, these changes also alter pore size and permeability of the gels. PEG is often modified with adhesion proteins (e.g., fibronectin, collagen, laminin) or more commonly with short peptide sequences (i.e., ArgeGlyeAsp, typically referred to as RGD).62 Although PEG is nondegradable, PEG hydrogels can be copolymerized with degradable polymers, such as PLGA, in varying ratios to tune the degradation kinetics.63 Alternatively, protease-cleavable crosslinkers can be incorporated to permit cell-mediated degradation.64

2.3 Hybrid Materials To combine the biological functionality of natural materials with the improved control of synthetic materials, the design of novel hybrid or semisynthetic materials has been an active area of research. In hybrid biomaterials, the synthetic polymer imparts the key physical qualities of the material, such as the mechanics and molecular structure, whereas the natural protein or polysaccharide endows the material with bioactive

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features. Typically, hybrid materials are produced through copolymerization reactions between the synthetic polymer precursor and the natural species of interest, often using free-radicaleinitiated copolymerization or click chemistry.65 Often, the protein or ECM species is covalently modified with linear or branched, nonionic hydrophilic polymers, such as PEG, and then cross-linked using standard conditions. For example, a fibrinogenePEG hybrid material was created by PEGylating denatured fibrinogen fragments with PEG diacrylate (PEGDA) groups, subsequently mixing the polymers with a photoinitiator, and crosslinking with UV light exposure.66 Although there are advantages in using hybrid materials, there are also a number of challenges, including the dependence of downstream protein function on the specific chemical configuration. For example, the quantity, length, and specific architecture (linear vs. branched vs. dendrimeric) of the polymer chain in addition to the site of attachment have all been observed to influence the ultimate bioactivity.67

3. HYDROGEL MATERIAL PROPERTIES 3.1 Degradation In the body, cells often remodel their surrounding ECM via matrix production and degradation, particularly postinjury, and these processes allow cells to migrate and spread to form cellecell contacts and produce functional tissue structures.68 Local degradation of native ECM is typically mediated by cell-secreted proteases such as MMPs, a process which can be recapitulated to a certain extent when using hydrogels derived of natural materials.69 However, many synthetic hydrogels rely on dissolution or hydrolytic degradation mechanisms, which are typically based on the density of cross-linking or the quantity of hydrolytically unstable chemical moieties and are thus not dependent on celldriven mechanisms.7 This lack of control over material degradation has motivated a number of novel approaches, including the creation of specific biohybrid materials, the incorporation of protease-cleavable peptide cross-links, and the design of photodegradable networks (Fig. 22.3). In one classic example, the creation of hydrogels of PEG modified with an RGD cell adhesion peptide and a highly sensitive MMP cleavage site enabled greater rates of fibroblast invasion and a more functional spindle-like fibroblast morphology when compared with PEG hydrogels with less sensitive proteasecleavable sites.64 In the context of 3D stem cell culture, MSCs cultured within alginate hydrogels containing the MMP degradable peptide PVGLIG in addition to

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FIGURE 22.3 Many of the key design features of hydrogels, including degradation kinetics, fiber orientation and topography, matrix mechanics, and permeability, are critical to promote cell survival and function and can be specifically tuned to mimic the native stem cell niche. Adding additional bioactive components, such as sites for cell adhesion, presentation of specific biologic signaling molecules, and stimuliresponsive elements, can be used to further regulate the environment of the encapsulated cells.

the RGD adhesion peptide exhibited a more spread morphology and formed cellular networks and extensive cellematrix interactions. In contrast, MSCs cultured within alginate containing only the RGD peptide remained round and had fewer cellecell interactions.70 Although the incorporation of protease-cleavable sequences within hydrogels allows for cell-mediated degradation, there is little external control over the system. Thus, the advent of specific photodegradable hydrogels that allow for spatiotemporal control over degradation provides for enhanced control of stem cellehydrogel environments. Hydrogels constructed of PEG modified with a nitrobenzyl etherederived moiety permit hydrogel cleavage upon exposure to UV light (365 nm), leading to local decreases in crosslink density in a nontoxic and spatially controlled manner.71 In addition to changing physical properties of the gel, photodegradation can also release or remove bioactive components. Photolabile RGD peptides incorporated into PEG gels containing MSCs were removed via irradiation after 10 days in culture, at a time when MSCs differentiating to chondrocytes downregulate

fibronectin. The specific temporal removal of the fibronectin-binding RGD peptide led to a fourfold increase of GAG production by the MSCs, suggesting enhanced chondrogenic differentiation and that the MSCs were responding to the change in extracellular signaling.72

3.2 Topography In native cell microenvironments, oriented fibers of ECM molecules provide topographical cues that influence cell polarity, migration, and ECM interactions. In fiber networks, it has been observed that smaller fiber diameters enhance cell adhesion to multiple fibers and lead cells to adopt a round morphology, as they sense traction forces in multiple dimensions. In contrast, cells on fibers with larger diameters tend to adhere to a single fiber and thus adopt an elongated morphology along the axis of the fiber because of the unidirectional traction force.73,74 To create nanofibers with polymers, electrospinning techniques are typically used, in which the polymer is

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expelled from a charged needle and the resulting fibers are collected on a grounded surface and aligned. The relative diameter and porosity of the scaffold can be adjusted by changing the concentration of the polymer, the flow rate, or the voltage.75 Although most efforts have focused on semicrystalline polymers, recent work has translated electrospinning techniques to hydrogel fibers.69,76 By electrospinning hydrogel precursors and cross-linking them in a dry rather than a hydrated state, soft electrospun hydrogels can be formed by placing the materials in an aqueous environment and allowing them to swell.77 Additionally, one of the issues with electrospinning is that small pore size may limit the infiltration and migration of cells. A novel approach to overcome this problem is to electrospin photocross-linkable polymers such that photopatterned pores can be created. MSCs cultured in electrospun HA/polyethylene oxide hydrogels with photopatterned pores exhibited increased infiltration into the aligned scaffolds.78 In addition to electrospinning, there have been a number of alternative approaches to recapitulate the fibrillar structure observed in native tissues, including the use of self-organizing peptides, which form entangled nanofibers because of sequences of alternating hydrophilic and hydrophobic side groups.73,79,80

3.3 Matrix Mechanics The mechanical microenvironment for cells within the body can vary by orders of magnitude, from very soft (e.g., w0.1 kPa) in the brain to very stiff (e.g., w80 kPa) in precalcified bone.81 Thus, to properly mimic the mechanical aspects of tissue niches, hydrogels of varying composition and biological cues should be capable of different degrees of stiffness. A wide body of work has examined the influence of mechanical stiffness on stem cell differentiation and phenotype, with the majority of these studies conducted by culturing cells in 2D on different hydrogel substrates. To examine whether the findings of 2D studies translate to 3D environments, MSCs were encapsulated in RGD-modified alginate gels produced with a wide range of mechanical properties (elastic moduli from 2.5 to 110 kPa). Similar to previous results with 2D systems, softer environments (2.5e5 kPa) promoted adipogenic differentiation of MSCs, whereas materials of intermediate stiffness (11e30 kPa) promoted osteogenic differentiation. However, the previous correlation with 2D morphology was not observed in the 3D setting (i.e., greater spreading was not observed in stiffer hydrogels), and investigation into integrin binding indicated that the underlying mechanism for the osteogenic lineage induction may be because of changes in the presentation of adhesive ligands.49

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Mechanotransductive pathways that cause the observed phenotypic changes are not yet completely understood, though recent work indicates that the transcriptional coactivators YAP (Yes-associated protein) and TAZ (transcriptional coactivator with PDZbinding motif), both part of the Hippo pathway, may be key regulators that can sense cytoskeletal tension.82 The translocation of YAP to the nucleus indicates activation and seems to be “switched on” based on the surrounding mechanical environment; for example, YAP is localized to the cytoplasm when cells are cultured on “soft” substrates (w1 kPa) but localizes to the nucleus when cells are cultured on a “stiff” substrate (w40 kPa).83 This switch-like behavior has led to studies examining the mechanical “memory” of MSC populations. By using photodegradable PEG hydrogels that can be “softened” (10e2 kPa) upon exposure to UV light, it was observed that while YAP/TAZ activation is reversible when switched to a soft substrate after 1 day of culture on a stiff hydrogel, MSCs cultured on the stiff hydrogels for 10 days retained active YAP/ TAZ activity for at least 10 days even after the switch to the soft hydrogel.84 The results of this study indicate that stem cells not only have immediate downstream responses to their surrounding mechanical environment, but they in fact retain a mechanical memory of their past surroundings. Studies that continue to investigate the influence of three dimensionality may provide different clues to the mechanisms underlying celle matrix mechanical interactions. One of the primary challenges in deciphering the specific role(s) of matrix mechanics on cell fate is the difficulty in separating changes in mechanical stiffness from the associated changes in hydrogel porosity and solute permeability, as the properties are inherently linked. Efforts to decouple these variables have led to a number of novel strategies, including the design of hydrogels, which incorporate methacrylic alginate with PEG dimethacrylate. In this system, the elastic modulus can be adjusted by changing the concentration of methacrylic alginate and the degree of substitution of the methacrylic groups. However, these changes do not alter the bulk hydrogel permeability because of the hydrophilicity of the network.85 Another approach has been to use different polymer configurations that exhibit similar permeabilities but different mechanical properties, such as a structured a-helix [formed of poly(propargyl-L-glutamate) derived from g-propargyl-Lglutamate] versus a flexible random coil [formed of poly(propargyl-L-glutamate) derived from equal parts of gpropargyl-L-glutamate and g-propargyl-D-glutamate].86 In addition to the challenge of decoupling hydrogel permeability from matrix mechanics, hydrogels from natural materials often have the additional complication of separating the mechanical and biochemical factors.

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For example, if the concentration of the natural polymer is increased to augment the mechanical stiffness, the density of bioactive groups, such as adhesive peptides, is also increased. Varying the density of adhesive moieties can have downstream phenotypic effects, such as changing cell spreading, growth, and migration.73

3.4 Mass Transport Properties Stem cells cultured statically within hydrogels generally rely on diffusive mass transport to receive the necessary oxygen and nutrients required for cell survival, and thus, examining mass transport properties is a key design parameter, particularly for large scaffolds. The rate of diffusion of molecules through the hydrogel is dependent on a number of factors, including the pore size of the hydrogel network and the charge of the material and the diffusing species.7 Several approaches have been applied to increase hydrogel pore size, including electrospinning (see Section 3.2), gas foaming, lyophilization, solvent casting with particle leaching, and the creation of hydrophilice hydrophobic hybrid hydrogels.87 Increasing the pore size with these methods can also facilitate cellular processes, such as proliferation and migration, in addition to changing the system permeability; however, the isotropic porous environment created by these techniques is often distinct from the more fibrillar structure of ECM in native tissues. Instead of creating pores throughout the hydrogel system, an alternative approach based upon native vasculature has been to incorporate microchannels using microfabrication techniques such as molding with microneedles, soft lithography, bioprinting, photopatterning, and modular assembly.88,89

3.5 Quantitative Properties of Hydrogels To compare between different hydrogel systems, it is important to be able to assign quantitative parameters that describe material properties (Fig. 22.4). Many of these properties are interconnected, and several

theoretical and empirical correlations have been established to quantitatively relate them. For example, the cross-linking density (rx), which denotes the number of polymer chains present in a set volume of material, is directly associated with volumetric swelling ratio, the mesh size, and the shear modulus. Briefly, the volumetric swelling ratio (Q) is the proportion of hydrated gel volume to dry gel volume, the mesh size (x) describes the diffusivity of species through the gel, and the shear modulus (G) defines the mechanical properties of the gel in response to shear stress.8 The mechanical properties can also be defined by the Young’s or elastic modulus (E), which also describes the stiffness of the system but alternatively describes the material’s response to uniaxial stress and is typically defined as the ratio of tensile stress to extensional strain specifically in the linear elastic region of the stressestrain curve. Some of these parameters are easier to empirically determine than others; for example, the swelling ratio (Q) is easily measured and the shear modulus can be determined using standard rheometric techniques. Using these two values, the harder to measure crosslinking density (rx) can be derived using the following Eq. (22.1), based on elastic theory: G ¼ RTrx Q1=3

(22.1)

where R is the universal gas constant and T is the temperature.90 The mesh size (x) can be determined by first estimating the average molecular weight between crosslinking points (Mc), which is described by the following Eq. (22.2) relating the cross-linking density (rx) and the elastic modulus (E): Mc ¼ ð3rx RTÞ=E

(22.2)

From there, the mesh size is described by the following Eq. (22.3):  1 Mc 2 x ¼ 2a  2:21Q1=3 (22.3) Mr where a and Mr are dependent on the polymer species of interest.91 The diffusivity (D) of a species through a

FIGURE 22.4 Key quantitative properties of hydrogels, including the volumetric swelling ratio (Q), the mesh size (x), the shear modulus (G), the Young’s or elastic modulus (E), and the diffusivity (D), are interrelated and can be used to compare across hydrogel platforms.

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hydrogel network of a known mesh size can be determined if the diffusivity of the species through a pure solvent (Do) and the hydrodynamic radius of the species (rs) are known. This relationship is described by the Eq. (22.4) below:  D ¼ Do

 rs Y 1 e x

Q Q1


(22.4)

where Y describes the ratio of the volume required for translational movement of the entrapped particle to the free volume per molecule of solvent (typically approximated as 1 because of empirical findings).92

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The sequence and density of adhesive peptides have been found to specifically impact stem cell phenotype. Although it has been generally observed that the presence of the RGD peptide in PEG gels leads to higher viability of encapsulated MSCs, it has been determined that the flexibility and spacing of the RGD peptide play a role in MSC survival. PEG gels with RGD attached via a single link as a pendant group promoted improved MSC survival (w80% viability) when compared with PEG gels with constrained RGD peptides (w60% viability), and the addition of a glycine spacer arm to the tethered RGD further augmented MSC viability to w88%.96 The results of this study indicate that the spatial configuration in which the RGD peptide is presented to the MSCs affects the cell’s ability to interact with it and that increasing accessibility can be achieved by slightly extending the peptide away from the PEG backbone.

4.1 Cell Adhesion Motifs A fundamental aspect to enable bioactivity in biofunctional hydrogels are adhesive ligands for cell attachment; without them, many of the key material properties, such as topography and matrix mechanics, cannot be accurately sensed by entrapped cells, and the cells will be unable to spread or migrate through the system. This is particularly important for stem cell populations, most of which are typically anchorage dependent. Depending on the cell type and its native microenvironment, different adhesive cues for specific cell surface receptors are more appropriate than others. Typically, cells bind to the surrounding ECM through transmembrane receptors known as integrins, which in turn bind to adhesive proteins found in the ECM such as fibronectin and laminin. Rather than incorporate full-length proteins into synthetic polymers, short peptide sequences have been identified which facilitate cell adhesion and can be more easily incorporated because of their small size, which is less disruptive to the overall hydrogel structure. The three-peptide sequence arginine-glycine-aspartic acid (RGD), which is found in many adhesion proteins, is by far the most commonly used because of its interactions with a broad range of integrins. Other common peptide adhesive sequences include isoleucineelysineevalineealaninee valine, which is derived from laminin and thought to bind a3b1, a4b1, and a6b1 integrins, and the collagenmimicking hexapeptide GFOGER, recognized by a2b1 and a11b1 integrin.93e95 Although many naturally derived hydrogels contain native adhesive sequences, synthetic polymers must be specifically modified to include them, and typical techniques include covalent grafting posthydrogel formation or chemical incorporation/physical entrapment during polymerization.73

4.2 Growth Factor Presentation In the native stem cell microenvironment, the ECM can bind and present a diverse array of growth factors that promote specialized cell functions and tissue morphogenesis. Thus, it can be important to include growth factors in hydrogel design schemes, either through direct incorporation and covalent tethering or through the addition of growth factorebinding motifs. Direct encapsulation of growth factors within the hydrogel has been the conventional approach to present localized soluble cues to entrapped cells and leads to sustained presentation when the hydrodynamic radius of the growth factor is larger than the hydrogel mesh size, as it will remain trapped such that it can provide signals to the cells in the vicinity. However, simple entrapment of growth factors has a number of limitations, including its restriction to large protein molecules and the general lack of control over subsequent release kinetics. Covalent tethering of growth factors to synthetic polymers, such as PEG, without adversely affecting bioactivity can be accomplished via a number of conjugation reactions, including succinimide or carboxylic acid with amine groups, maleimide with thiol groups, Michael type addition, or click reactions of azide with alkyne groups.97 Bioconjugated growth factors have been used in a variety of stem cellehydrogel combinations, including NSCs within 3D chitosan hydrogels modified with interferon-gamma, which improved differentiation toward neuronal lineages.98 Many of the studies performed in 2D, such as the immobilization of stem cell factor and SDF-1a on PEG hydrogels to promote HSC adhesion and spreading, will benefit from replication in 3D systems to see if the 2D functional results can be recapitulated in a configuration more

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similar to the in vivo environment.99 A general caveat of covalent tethering approaches is the potential attenuation or loss of protein bioactivity because of a lack of control over site-specific attachment and orientation of the ligand such that its associated receptor can properly bind. The entrapment of growth factoreladen microparticles or microspheres into larger bulk hydrogels or directly within cellular aggregates has been investigated to improve the control over growth factor release while preventing complications from covalent tethering. There is a plethora of materialeprotein combinations available because of the breadth of research in the drug delivery field regarding materials with controlled-release profiles. For example, modified PEG microspheres loaded with TGF-b3 and BMP-6 entrapped within chitosan hydrogels were found to enhance the chondrogenic differentiation of adipose-derived stem cells, with the local delivery from the microspheres leading to a higher ratio of collagen II to collagen I when compared with soluble growth factor delivery, which may be because of the manner in which the growth factors were present or the difference in local concentration.100 An advanced system has recently been developed for the simultaneous formation of coacervate microparticleeentrapped hydrogels using BMP-2eladen methacrylated alginate and MSC-containing methacrylated gelatin to promote osteogenic differentiation.101 Using photo-cross-linking, the constructs can be fashioned in situ, reducing the complexity of current systems in a noncytotoxic manner. An alternative approach to incorporating specific growth factors within hydrogels from the outset is to sequester growth factors, either from the media solution or from those produced by cells. Systems of this nature more closely mimic native tissue environments in which PGs and GAGs interact noncovalently with growth factors, thus many of the current sequestering approaches incorporate native GAG species or peptide fragments that bind GAGs.102 For example, the incorporation of a heparin-binding peptide sequence into fibrin gels facilitated binding and controlled release of soluble betanerve growth factor and other neurotrophic factors.103 The recent identification of a domain in placenta growth factor-2123e144, which binds with very high affinity to a number of ECM molecules suggests that protein engineering strategies in which this sequence is fused to a growth factor of interest, may be used to enhance growth factor sequestration.104 In addition to using ECM-based protein sequestration strategies, peptidebinding motifs for specific growth factors can be directly incorporated into engineered materials. For example, a peptide derived from vascular epidermal growth factor (VEGF) receptor-2 incorporated into PEG microspheres promoted specific binding of VEGF with high affinity, which has implications for angiogenic applications.105 A similar approach could be used for a variety of

situations in which increasing or decreasing growth factor activity is desired.

4.3 Stimuli-Responsive Materials Native biomacromolecules are dynamically regulated by their surrounding environment; thus, the design of hydrogel systems which exhibit a specific response to an external stimulus adds an extra dimension to hydrogel capabilities. Common external stimuli include altering the pH or temperature, treatment with light or electromagnetic fields, introduction of specific enzymes or peptides, or application of mechanical forces.106 A caveat is that the applied stimulus should not cause cytotoxic effects; thus, there is a small range of acceptable pH and temperatures and limitations in the wavelength and duration of UV exposure that are generally considered tolerable. The inclusion of photolabile groups such that the hydrogel structure can be dynamically modulated has already been discussed in the context of changing the degradation and mechanical properties (see Sections 3.1 and 3.3), and the recent incorporation of photo-activated caged RGD residues in PEGDA hydrogels enables spatiotemporal control of hydrogel adhesive properties in vivo through transdermal light exposure.107 Thus, photo-activated material systems continue to be a powerful tool for dynamic regulation of physical and biochemical material properties. Novel culture systems that take advantage of the native material’s sensitivity to temperature changes also offer possibilities for real-time control. Thermoreversible PNIPAAm-PEG hydrogels, which form gels at physiologic 37
C but remain liquid at 4
C, have recently been investigated as a 3D culture platform for human pluripotent stem cell (hPSC) culture.108 Cells were mixed with the polymer in its low-temperature liquid state, grown in the hydrogel at 37
C, and easily retrieved for further analysis or expansion by simply reducing the temperature. Using this fully defined system, significant expansion of multiple hPSC lines was observed (approximately 20-fold expansion per 5 day passage) while the cells retained their pluripotent qualities, indicating that thermoresponsive hydrogel systems can serve an important purpose in developing scalable bioprocesses for hPSC expansion.

5. TECHNOLOGIES FOR HYDROGEL PRODUCTION AND ASSESSMENT 5.1 Hydrogel Production Methods Hydrogels are useful tools in 3D stem cell culture across a range of scales, spanning from microparticles

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at the micrometer scale, to microcapsules at the submillimeter scale, and finally to bulk gels at the centimeter scale (Fig. 22.5). All three configurations can be used to encapsulate cells within a matrix, though microparticles are more typically used in scaffold-free cell aggregate configurations or in combination with larger hydrogels. The production of bulk hydrogels typically uses molds that are filled with the polymer solution before cross-linking, and more sophisticated shapes can be achieved by using microfabrication technologies such as rapid prototyping and soft lithography. In contrast to forming whole hydrogel systems, a modular bottom-up approach can also be used in which small units are fabricated for future assembly into larger constructs.109 This is similar to the way in which many native tissues are composed; for example, muscle is made up of many myofiber bundles. The construction of larger structures from smaller units can be performed manually or through the use of self-assembled materials. With hydrogels, one can take advantage of hydrophobicehydrophilic interactions at liquideliquid and aireliquid interfaces to drive self-assembly of microgels.110 5.1.1 Formation of Microparticles Hydrogel microparticles have been used for controlled-release drug delivery purposes for some time by engineering material properties to control release profiles.111 For the most part, emulsion polymerization techniques are used to produce droplets of the polymer phase within a surfactant-stabilized carrier solution before cross-linking of the hydrogel.112 Microparticles loaded with small molecules or growth factors can be incorporated into bulk hydrogels to facilitate controlled release of soluble cues to encapsulated stem cells.113 Alternatively, microparticles can be incorporated into

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scaffold-free 3D multicellular constructs and act as local reservoirs of soluble factors that may otherwise face diffusional barriers if delivered externally.114,115 Spatial control over microparticle position within cell constructs can also be used to promote morphogenic events stimulated by soluble factor gradients.116,117 5.1.2 Microencapsulation Methods The formation of spherical microcapsules or microbeads can be achieved by a number of available methods including the use of micromolds or microfluidic droplet generators.7 Specialized microencapsulation devices use a variety of methods to produce small, uniform hydrogel spheres, with most relying on electrostatic potential, coaxial flow, vibration, or water jet cutting to form single beads from a stream of liquid. In all cases, microencapsulation techniques rely on rapid gelation of the hydrogel material after bead formation. A distinction exists between microcapsules, which have a solid shell with a liquid, and microbeads, which are a solid hydrogel, and the two configurations may be suited for different applications.21 For example, liquid core capsules facilitate aggregation of enclosed cells into a single aggregate, whereas cells in solid microbeads are more constrained and often proliferate to form multiple smaller and distinct aggregates. 5.1.3 Microfabrication Techniques for Bulk Hydrogels Techniques for producing hydrogels with a high degree of spatial control often rely on microfabrication techniques that have a fine control over positioning. Many complex hydrogel structures are enabled through microfabrication technology, including the formation of custom, multilayer tissue-engineering constructs.118 Unfortunately, many of the standard fabrication methods

FIGURE 22.5

Hydrogels are useful across a range of scales, spanning from microparticles at the micrometer scale, to microcapsules at the submillimeter scale, and finally to bulk gels at the centimeter scale. Each of these types of hydrogels can play a key role in controlling stem cell function.

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used to produce highly defined 2D patterns are not easily translated to 3D systems, prompting the development of novel fabrication techniques. Soft lithography techniques often have the highest resolution (w100 nm) and include a range of methods that facilitate the design of micro-features from elastomeric materials, typically PDMS, which can then be used as master molds for bulk hydrogel formation. Stereolithography approaches offer less spatial resolution (w1 mm) than soft lithography, but can be used directly with photopolymerizable hydrogels to create structures in a layer-bylayer approach. The advent of 3D printing technologies has also been applied to create hydrogel systems, though the spatial resolution (w10 mm) is currently less than that obtained with other microfabrication techniques.73

5.2 Translating 2D Assays to 3D Systems Despite the advantages of culturing stem cells in hydrogels, many of the routine assays to monitor outcomes, including cell number, viability, metabolic activity, and phenotype, are not easily translated to 3D systems. Common assays that make use of cellular dyes can be limited by diffusion through the hydrogel construct, though often increasing the dye concentration and/or incubation time or introducing convective flow (i.e., mixing) into the system can overcome mass transport constraints. Even if mass transport is not an issue, which it typically is not for common small-molecule stains (<1000 Da, including DAPI/Hoechst, calcein-AM, and propidium iodide), obtaining images of cells within hydrogels can be challenging.7 If the hydrogel is not fully transparent, live staining and imaging may not be possible, thus restricting visualization to after processing and sectioning. Even if the hydrogel is transparent, of densely packed regions of cells can limit imaging penetration, though systems such as two-photon confocal microscopy and light-sheet fluorescence microscopy may be able to optically section 3D cell constructs.7 In addition to real-time monitoring, retrieving cells embedded within hydrogels for phenotypic analysis or future applications can be difficult if the hydrogel materials are not degradable. Themoresponsive hydrogels, such as agarose, methylcellulose, and PNIPAAm, can be easily reversed from the gel phase, but many other hydrogels require proteases such as trypsin and collagenase to be degraded. To avoid the use of proteases, which often have secondary effects on cell viability and function that confound results, a novel hydrogel made of nanofibrillar cellulose was recently developed for hPSC culture.119 The simple addition of cellulase, which has no negative impact on mammalian cells, can be used to readily harvest the encapsulated cell population.

6. CHALLENGES AND FUTURE DIRECTIONS 6.1 Spatiotemporal Control Over Biophysical and Biochemical Hydrogel Properties Because of the dynamic nature of native tissues, techniques that facilitate spatiotemporal control over hydrogel constructs will lead to more advanced mimics of in vivo stem cell microenvironments. Because the physical and biochemical features of hydrogels both play important roles in directing cell fate, approaches which target multiple properties of the hydrogel are essential. Advances in photoreversible hydrogel systems are promising candidates; for example, click-based PEG hydrogels have been designed with independently activated sites for photocleavage and photoconjugation, enabling dynamic control over hydrogel cross-linking density and the activation of conjugated biomolecules.120 Innovations in material design will be essential to move the field from relying primarily on 2D systems to 3D practices and eventually to “4D” systems, which have controlled and temporal sensitive components.

6.2 Promoting Endogenous Tissue Morphogenesis In light of the striking results from recent organoidformation studies, it is essential that hydrogel systems that facilitate and even promote emergent cell behaviors are considered. Although it is now possible to design complex hydrogel systems with multiple functional attributes, it could be that only simple, supportive environments are necessary to facilitate self-organization and self-instructive stem cell processes. In organoidformation studies thus far, the cell-directed morphogenesis has occurred either in scaffold-free suspension culture or in the presence of Matrigel, a complex and difficult to reproduce material. Therefore, it may be that the necessary signals for stem cells to produce complex tissues may be provided entirely by the cells themselves and that challenges in future material design will be to determine strategies that enhance endogenous processes.121 Additionally, materials that allow for replicative manufacturing and downstream transplantation will be essential for translating cell therapies emerging from the formation of in vitro tissue constructs.

7. CONCLUSION In summary, hydrogels enable advances in the 3D culture of stem cell populations, which will be essential in future explorations of stem celleniche interactions and the production of cell-based therapies. Different

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REFERENCES

hydrogel sources are available, with natural ECM-based materials offering inherent bioactivity but low external control over specific material properties. In contrast, synthetic polymer systems provide users with greater control to engineer materials with very specific chemical and physical properties. Many of the key design features of hydrogels, including degradation kinetics, fiber orientation and topography, matrix mechanics, and permeability, are critical to promote cell survival and function and can be specifically tuned to mimic the native stem cell niche. Adding additional bioactive components, such as sites for cell adhesion, presentation of specific biologic signaling molecules, and stimuliresponsive elements, can be used to further regulate the environment of the encapsulated cells. Although standard techniques for hydrogel formation and monitoring exist, further advances in microfabrication methods for production of hydrogels will enable more advanced geometries, while improved imaging and analytics will be critical for supervising stem cell behavior within the hydrogels. Research in innovative materials that can provide precise spatiotemporal control and/or promote the endogenous self-organization of stem cells into complex tissue structures will be critical enabling technologies for future opportunities in the stem cell field.

ABBREVIATIONS AND ACRONYMS EBs Embryoid bodies ECM Extracellular matrix ESCs Embryonic stem cells GAGs Glycosaminoglycans HA Hyaluronic acid iPSCs Induced pluripotent stem cells MMPs Matrix metalloproteinases MSCs Mesenchymal stem/stromal cells NSCs Neural stem cells PEG Polyethylene glycol PGs Proteoglycans PNIPAAm Poly N-isopropylacrylamide PSCs Pluripotent stem cells RGD Arginine-glycine-aspartic acid peptide UV Ultraviolet

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