Designing Cell-Compatible Hydrogels for Biomedical Applications

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Science  01 Jun 2012:
Vol. 336, Issue 6085, pp. 1124-1128
DOI: 10.1126/science.1214804


Hydrogels are polymeric materials distinguished by high water content and diverse physical properties. They can be engineered to resemble the extracellular environment of the body’s tissues in ways that enable their use in medical implants, biosensors, and drug-delivery devices. Cell-compatible hydrogels are designed by using a strategy of coordinated control over physical properties and bioactivity to influence specific interactions with cellular systems, including spatial and temporal patterns of biochemical and biomechanical cues known to modulate cell behavior. Important new discoveries in stem cell research, cancer biology, and cellular morphogenesis have been realized with model hydrogel systems premised on these designs. Basic and clinical applications for hydrogels in cell therapy, tissue engineering, and biomedical research continue to drive design improvements using performance-based materials engineering paradigms.

Hydrogels, which are three-dimensional (3D) networks composed of cross-linked hydrophilic polymer chains, can be cast into practically any shape, size, or form and can absorb up to thousands of times their dry weight in water (Fig. 1). In medicine, dozens of hydrogel products based on synthetic or natural polymers have had an enormous effect on patient care in recent years. For example, soft contact lenses are typically made from poly(hydroxyethylmethacrylic)acid [poly(HEMA)] (1), and biological adhesives made from reconstituted fibrin or albumin are routinely used in surgical procedures (2). Wound dressings made from alginate polysaccharide have been used for over a quarter century (3), and fillers made from hyaluronic acid (HA) have been used for several clinical indications (4).

Fig. 1

Hydrogels can be cast into practically any shape, size, and form. (A) Polymer powder (far left) added to a drop of water (left) and cross-linked results in the formation of a hydrogel (middle). Even if the hydrogel network is dehydrated (right), it still retains its overall shape. A sesame seed (far right) is shown for scale. (Insert) The stages of the dehydration process are shown from left to right. (B) Water-borne microgels in suspension containing several different immobilized molecules (color-coded) are shown in a single drop of water so as to highlight the possibilities for injectable hydrogel drug delivery. (C) A transparent microgel containing smaller color-coded microgels is shown to highlight new “gel-in-gel” experimental tools for biomedical scientists. (D) A microgel encapsulating fluorescently labeled cells is shown to highlight potential uses in cell delivery for tissue regeneration. (E) Encapsulated fibroblast cells thrive in a semi-synthetic PEG-fibrinogen microgel, illustrating how highly compatible milieus used for studying cell behavior may one day replace conventional cell culture paradigms in cancer research, stem cell, and development biology. Scale bars, 500 μm.

New applications for hydrogels have emerged, most notably in stem cell and cancer research, cell therapy, tissue engineering, immunomodulation, and in vitro diagnostics. These applications require more than mechanical and chemical versatility from hydrogels; they require cell compatibility based on biomimicry of the extracellular matrix (ECM) (5). The emerging field of cell-compatible hydrogel materials is therefore defined by design strategies focused on tuning the biological and physical attributes of hydrogels in order to achieve specific interactions and responses from cellular systems. In stem cell research, for example—in which hydrogels are applied to investigate cell fate by regulating key variables of the biomimetic niche—complex patterns, and gradients of chemokines, cytokines, growth factors, and other biophysical properties are systematically designed into the bioartificial stem-cell microenvironments to mediate cell activity at the molecular level (68). For hydrogels that facilitate controlled drug release or optimized pharmacokinetics, the porous structure of the material and its affinity to its therapeutic payload can be chemically controlled without attenuating inherent physicochemical compatibility to natural tissues (9). In regenerative medicine, in which hydrogels are used to hierarchically organize cells into tissue-like structures, the architectural and/or molecular cues can be engineered with spatial and temporal presentations that mediate multicellular morphogenesis (10).

Over the past 15 years, the challenges of designing complex hydrogel systems have been aided by major breakthroughs in synthetic polymer chemistry (11), 3D molecular patterning techniques (12), and biomimetic rational design approaches that are founded on the basic principles of cell and molecular biology (10). These advances have set the framework for overcoming some of the enduring challenges in applying biomedical hydrogels to additional areas of science and medicine. This review covers the design principles being applied to engineer cell-compatible biomedical hydrogels, focusing on biophysical and biochemical manipulations of 3D hydrophilic networks for controlling interfaces at the cellular and subcellular scales (Fig. 2). A number of recent publications review hydrogels in terms of their application and composition (6, 9, 1317). Here, we highlight the important role of performance-based engineering concepts in advancing the development of customized hydrogels that contain a set of desired properties based on specific molecular building blocks.

Fig. 2

Interface between cell and hydrogel. A false-colored high-resolution cryogenic scanning electron micrograph depicts the 3D interface between a fully hydrated hydrogel (blue) and an encapsulated cell (brown). Scale bar, 1 μm. The insert (bottom right) shows a magnified view of the cell-hydrogel interface, highlighting the cell membrane in the hydrated gel phase. Scale bar, 300 nm. (A to D) The four other inserts illustrate the cell membrane-hydrogel interface (blue, polymer network; white, void space) in different types of hydrogel structures. Cell receptor proteins (green) lie on the cell membrane (gray), whereas soluble bioactive molecules (red) and tethered bioactive factors (purple) lie within the hydrogel mesh. The length scale of the swollen polymer network can be used to control the transport of factors as well as receptor-ligand interactions and even cell motility. For example, hydrogel structure may facilitate amoeboid cell motility within micrometer-sized pores (D), but not through nanometer pores (A). Similarly, self-assembled and fibrous structures [(C) and (D)] of bioactive hydrogels enable focal adhesion contacts between cells and the network through receptor-ligand clustering, whereas in the dense or amorphous structures [(A) and (B)], these interactions are far less favored because of the spatial distribution and confinement of receptors and ligands (10).

Designing Hydrogel Molecular Building Blocks

A cell-compatible hydrogel is characterized by its ability to control specific molecular interactions at the cell-material interface. These include biological interactions such as receptor-ligand complexes that mediate cell adhesion and mesenchymal migration, bound or soluble molecule interactions facilitating proteolytic biodegradation or transcriptional events that govern cell phenotype, as well as focal adhesion interactions with compliant or rigid substrates to transmit mechanical stresses to cells. Designing the molecular structure of a hydrogel to control these spatial and temporal events enables their use for guiding cell response in vivo and in vitro. Many features that control cell response, which have been isolated from part of a cell’s in vivo microenvironment, can be built into a hydrogel using a top-down rational design.

Two seminal studies in the late 1990s demonstrated the top-down approach using recombinant protein domains that are responsible for the mechanism in which natural proteins undergo reversible temperature-induced gelation. Tirrell and co-workers designed hydrogels that undergo reversible physical cross-linking based on shifts in temperature or pH, using genetically engineered protein domains containing coiled-coil regions (18). Kopecek and co-workers developed similar stimuli-responsive hydrogels based on grafts of engineered coiled-coil or beta-sheet protein domains with a synthetic polymer backbone (19). They showed that the engineered hybrid system exhibited biomimetic self-assembly, as well as control over physical properties afforded by the synthetic constituent.

Many cell-compatible hydrogel materials have since been modeled on these seminal concepts, most notably the proteolytically responsive synthetic hydrogels pioneered by Hubbell and co-workers (20). They used poly(ethylene glycol) (PEG) macromeres that are cross-linked by oligopeptides that mimic collagenase substrates found in natural ECM proteins. The end-linked PEG-co-peptide hydrogels were formed by Click chemistry, in which a Michael-type addition reaction between the di-thiolated oligopeptides and vinyl sulfone groups on the PEG is carried out in the presence of cells and/or tissues. The chemical cross-linking reactions produced strong elastic gels in mild conditions and ensured high in situ cell survival. This methodology also represented a breakthrough in that it demonstrated the ability to transform an otherwise inert synthetic hydrophilic polymer network into a cell-responsive biomaterial that could undergo controlled protease-mediated dissolution. Other studies have combined a variety of synthetic polymers and protease substrates—showing that almost any synthetic hydrophilic polymer milieu can be designed to facilitate controlled cellular degradation and invasion. Other biomimetic features isolated from bioactive domains in natural proteins have been used in similar fashion, including cell-adhesive integrin-binding domains (21), controlled release affinity binding domains (22), and transglutaminase cross-linking domains (23).

Some applications may require more extensive biomimetic features from the backbone constituents of the biomedical hydrogel. Certain stem cells, for example, require extensive biochemical stimuli inherent to their natural ECM niche for proper differentiation (6, 7, 16). These stimuli are difficult to replicate with synthetic biology, and a semi-synthetic approach may appropriately accommodate the discrepancy between the completely natural versus completely synthetic hydrogel microenvironments. The semi-synthetics use proteins or polysaccharides conjugated to low-molecular-weight hydrophilic polymers as the main building blocks of the hydrogel (13). The protein/polysaccharide provides biomimetic features, and the synthetic polymer provides important control features for regulating mechanical properties, molecular structure, and other physical attributes of the material.

One such family of semi-synthetic hydrogels is based on the covalent conjugation of linear or branched, nonionic hydrophilic polymers with reconstituted ECM proteins such as fibrinogen, collagen, or albumin (24). The liquid precursors to these hydrogels are synthesized by means of Michael-type addition under reducing conditions with acrylate-functionalized polymers such as PEG, whereas the formation of an elastic gel is accomplished by a light-activated chemical cross-linking. The biological consequence of conjugating the ECM proteins does not alter the inherent compatibility of these native components, and the physical properties of the conjugated polymers are retained or enhanced in the protein-polymer adducts. Shachaf et al. demonstrated this with reverse thermal gelation (RTG), a liquid-to-solid phase transition associated with increased temperature of certain amphiphilic polymers in solution. By conjugating fibrinogen to a triblock copolymer composed of a central hydrophobic chain of poly(propylene oxide) flanked by two hydrophilic chains of poly(ethylene oxide) [Pluronic F127 (BASF Corporation, Florham Park, NJ)], the RTG properties of the fibrinogen-Pluronic adducts were enhanced severalfold over the Pluronic alone while still retaining fibrin-like bioactive properties for 3D cell culture (25).

The semi-synthetic biomaterials are now routinely synthesized for biomedical research and clinical applications, including conjugates of HA, alginate, and chitosan (3, 15, 26). Conjugation in these materials has been achieved through stepwise copolymerization of the hydrophilic polymers with protein or polysaccharides, Schiff-bass formation reactions, disulfide bonding, free-radical–initiated copolymerization by using peroxides or a fenton reagent, photo-initiated free-radical copolymerization, or metal-free Click chemistry such as Michael addition (11). Although less defined than the synthetic ECM analogs, the semi-synthetic ECM hydrogels are relatively easy to manufacture, can be reproduced in large quantities and provide a more reliable material as compared with natural ECM hydrogels.

Although a top-down approach is beneficial, it has its limitations; for instance, it is difficult to precisely control spatial bioactivity given the structural complexity of natural protein domains. An alternative bottom-up approach provides the opportunity to control hydrogel molecular structure by arranging elementary chemical motifs together to give rise to a system possessing controlled yet complex patterns or gradients of bioactive factors, as well as other biophysical properties. The basis of this methodology is functional chemical features whose structure-function relationships are well characterized and which are easily implemented into the hydrogel’s cross-linking methodology. Some well-established chemical reaction schemes have therefore been recently adapted for the mild conditions often required with in situ formation of biomedical hydrogels (27).

Chemical reactions used for the immobilization of reactive macromolecules or localizing cross-linking reactions can be performed by using tightly regulated light-activated initiation either with traditional photolithographic techniques or more sophisticated approaches, such as multiphoton laser scanning lithography (12). Photolithography has been widely used for creating patterns in various hydrogels and, more recently, for patterning cell-laden PEG or alginate hydrogels with multiple functionalities by using free-radical photopolymerization (28). Burdick and co-workers developed a multiple-mode cross-linking methodology in which an ultraviolet (UV) light–induced radical cross-linking reaction morphologically confines 3D cell cultures in spatially controlled nondegradable stiff regions of an otherwise degradable HA-peptide hydrogel (29). The stiffer UV cross-linked regions in the hydrogel alter differentiation patterns of mesenchymal stem cells (MSCs) and thus may enable the development of niche microenvironments where heterogeneous cocultures are propagated through predefined stiffness and biodegradability.

Patterning techniques with submicrometer-scale resolutions have also been adapted for cell-laden hydrogels based on the concepts reported by Kawata and co-workers using multiphotonic photopolymerization (30). This methodology is applied to a pre–cross-linked hydrogel containing functionalized photo-liable moieties designed to accommodate either immobilization or dissociation reactions initiated by femtosecond laser pulses in the near-infrared range (31). Anseth and coworkers used this approach to creating predefined microchannels within nondegradable or degradable synthetic PEG gels using photolabile cross-linker molecules made with nitrobenzyl ether cross-linking moieties (32, 33). The nitrobenzyl ether cross-links dissociate upon exposure to light from a multiphoton laser scanning microscope system, reducing the local cross-linking density of the hydrogel and giving way to physical tracks for cell invasion via contact guidance.

Hydrogel Structure and Properties

The interrelationship between the hydrogel processing, structure, properties and performance underlie the fundamental design rationale for most biomedical applications. With cell-compatible hydrogels, this interrelationship is complicated by a performance criteria characterized by a multitude of molecular interactions at the cell/material interface (Fig. 3). Therefore, the design should focus on those salient hydrogel features that give rise to the desired properties most suitable for the biomedical application, including transport properties (such as sustained release), tissue interactions (such as bioactivity), and chemical stability (such as degradability). In principle, most features can be engineered into a hydrogel; however, here we will focus on five important properties for biomedical applications: degradation, bioadhesion, bioactivity, transport (for example, controlled release of bioactive molecules), and mechanical properties.

Fig. 3

(A to G) A multitude of chemical interactions underlie the complexity of a cell-compatible hydrogel and its cellular interface. When a hydrogel’s polymer chains are held together by irreversible covalent bonds (chemical interactions), the network is relatively strong and stable. When the dominant connections between the polymer chains are reversible molecular entanglements or physical bonds (physical interactions), the network is generally less robust. Degradation can be designed into the hydrogel by introducing either proteolytic or hydrolytic moieties into the backbone of the polymer network. Bioactivity can be designed into the hydrogel by immobilizing growth factors and/or small molecules onto the backbone polymer, or by designing the polymer to facilitate physical interactions with cellular and extracellular biomolecules.

Biodegradation of hydrogels is essential for biomedical applications that require controlled resorption in vivo and/or local dissolution to facilitate cell morphogenesis and motility. Hydrogels have the ability to undergo local or bulk dissolution based on a number of mechanisms (such as hydrolysis, proteolysis, disentanglement, or environmental triggers); engineering the spatiotemporal aspects of this presents a challenge. Bulk hydrolytic resorption with specific temporal events in the body, such as bone regeneration, can be achieved by controlling the amount of hydrolytically liable cross-links in the polymer network, resulting in better tissue repair. Patterson et al. created such resorbable bioactive implants for bone repair using photopolymerized HA modified with different amounts of glycidyl methacrylate (GMA) cross-linkers. The relative concentration of GMA on the HA was proportional to the hydrogel resorption rate in vivo, and this in turn affected the organization of the new bone formation (34).

Cell-mediated hydrogel degradation can provide a more physiological control mechanism for both the removal of the provisional matrix and the liberation of matrix-bound bioactive factors (10). Strategies using cell-mediated control over degradation use a peptide-polymer hydrogel design with cross-linking oligomers that are known substrates for collagenases, gelatinases, and other matrix metalloproteinases (20). A large number of oligomer sequences that are known to be responsive to these cell-secreted proteases have now been characterized and can be synthesized and incorporated into the material design. In tissue repair of a critical-size rat calvarial bone defect, bioactive hydrogel implants designed with oligopeptides engineered to facilitate an in vivo resorption rate that coincides with the normal repair timeline outperformed all other material variants (35).

Bioadhesion, an important property that allows cells and tissues to adhere to hydrogels has enabled their use as tissue adhesives in surgical repair or as inductive scaffolds for tissue regeneration. Although some hydrogels such as fibrin or collagen inherently exhibit bioadhesive properties, most other natural and synthetic hydrogels do not. Bioadhesive features can be engineered into a hydrogel network by using linker molecules that enable covalent or non-covalent molecular interactions between the implant and its surroundings, including cell-adhesive oligopeptides derived from fibronectin’s central cell-binding domains (FNIII 9-10) (21). Cell-adhesion modifications to hydrogel scaffolds have been used effectively to promote enhanced osteogenic differentiation of MSCs for bone repair (21), to provide an essential foothold for neurite outgrowth in axonal regeneration (36), and to understand the regulatory role of mechanotransduction in stem cell fate determination (8).

Tissue-adhesive modifications to hydrogels can further improve performance of cell-delivery scaffolds by stabilizing the in vivo location of the graft. For example, Messersmith and co-workers developed a catechol-based modification to a PEG hydrogel implant in order to improve extrahepatic islet transplantation for managing type I diabetes mellitus (37). The hydrogel modification, using a catechol moiety [3,4-dihydroxy-L-phenylalanine (DOPA)], is derived from the tethering chemistry that allows mussels to adhere to wet organic surfaces (38). Under oxidizing conditions, DOPA rapidly forms cross-links when bound, for example, to an end-functionalized star-PEG precursor. In addition to participating in the nontoxic cross-linking reaction of the PEG hydrogel, the catechol moiety forms strong covalent interactions with nucleophiles such as thiols and imidazoles found in organic substrates (ECM). The effective immobilization of the islet-containing catechol PEG hydrogels on the surface of the liver of diabetic rats enabled revascularization with minimal inflammation and permitted effective glucose management comparable with the gold-standard intrahepatic islet delivery.

Bioactivity in hydrogels is instrumental for materials that are called on to mediate specific biological events in the body based on endogenous cell recruitment, local morphogenesis, and controlled cell differentiation. Many of these events can be induced by using exogenous growth factors that are delivered with spatiotemporal control (7). However, hydrogels do not inherently sequester growth factors and thus fall short of precisely controlling the sustained or localized bioavailability of their payload. In vivo, growth factor bioavailability is tightly regulated by nonspecific associations between the factor and ECM proteoglycans [glycosaminoglycans (GAGs)] through affinity binding domains (10). Using strategies premised on such non-covalent interactions with polymeric building blocks, design modifications to hydrogels have recently been used to improve the localized growth factor availability (39). For example, Martino et al. modified fibrin hydrogels with an oligopeptide derived from a non-GAG growth factor binding domain that was isolated from the 12th to 14th type III repeats of fibronectin; the peptide-modified fibrin hydrogels required far less growth factor than unmodified hydrogels to achieve regeneration in either chronic wounds or critical-sized bone defects (22). These concepts have evolved to include short heparin-binding peptide domains that can sequester cell-secreted proteoglycans, such as endogenic heparin. Murphy et al. immobilized such peptides onto materials in order to take advantage of endogenic heparin binding and consequent growth factor immobilization (40). Another prevalent approach to sequester bioactive factors within a hydrogel network involves covalent growth factor immobilization. However, affinity-based linker peptides may be preferred because they circumvent loss of biological activity typically associated with chemical cross-linking of proteins.

Transport of hydrophobic/hydrophilic molecules is an important property of a hydrogel that can benefit therapeutic techniques requiring sustained drug release or triggered pharmacokinetics. For example, tumor chemotherapy may be far more effective if drug molecules were targeted to and sustained in the tumor site, or if a drug is released only after it has been internalized within the cancer cells. Depending on the properties of the therapeutic drug, a hydrogel’s porosity (mesh size) may be used to regulate the drug’s availability by controlling its diffusion through the polymer network (Fig. 2). The mesh size of a typical cell-compatible dense hydrogel network in the swollen state is no less than 5 nm and far greater than the characteristic size of most small drug molecules (41). For relatively large protein or peptide drug molecules, the hydrogel structure and mesh size (~5 to 20 nm) can be engineered to limit mobility and modulate release kinetics (Fig. 2A). In order to achieve sustained release of small drug molecules by using size exclusion, hydrogels must use immobilization schemes that entrap these molecules in dense physical structures within the network (mesh size < 1 nm), including in self-assembled nanostructures, layer-by-layer constructs, liquid crystalline nanostructures, and polyelectrolyte complexes (41). Premised on these concepts, Grinstaff and co-workers designed expansile hydrogel nanoparticles for targeting in vivo tumor suppression in a subcutaneous mouse cancer model with the small anticancer drug paclitaxel (42). The hydrophobic polymer-drug nanostructures, which are stable in water at physiological pH levels, are designed to release the drug only after a pH-triggered transition into a hydrophilic hydrogel network occurs inside the mildly acidic endosomes of the tumor cells. This triggered release of paclitaxel was effective in preventing tumor growth in vivo using far lower doses than are required with systemic administration of the drug.

The mechanical properties of a hydrogel can convey important physical cues to cells through mechanotransduction pathways that mediate tissue homeostasis, morphogenesis, cell growth, contractility, differentiation, and pathophysiology. As an example, Weaver and co-workers controlled the modulus of collagen hydrogels to confirm that ECM rigidity can affect the malignant phenotype of mammary epithelial cells through focal adhesion assembly mediated by either ERK activation or Rho activity (43). With new insights into the mechanical basis of tissue regulation emerging, a number of design strategies have evolved to improve control over a hydrogel’s mechanical properties. Young and Engler used a slow in situ chemical cross-linking reaction of HA-PEG hydrogels to accomplish time-dependent hydrogel stiffening that mimics temporal events associated with cardiomyogenesis. They engineered the timing of the Michael-addition reaction between thiolated high-molecular-weight HA and PEG-diacrylate cross-linker to recapitulate embryonic stiffening of cardiac muscle in order to show that the temporal pattern of stiffening enhances in vitro cardiac myocyte differentiation (44). Gong and co-workers developed a method for overcoming the inherent limitations of a hydrogel’s mechanical properties (such as low strength) using physical interpenetrating double-networks (DNs) (45). The DN hydrogels exhibited extremely high toughness when, for example, they combined a rigid polyelectrolyte gel, poly(2-acrylamido-2-methylpropanesulfonic acid), with a flexible neutral poly(acrylamide) gel. Applying this concept to a cell-compatible but inherently weak polymer such as reconstituted collagen, they produced DN collagen-poly(N,N′-dimethyl acrylamide) gels with an order of magnitude increase in fracture stiffness while still retaining more than 90% water content.

Performance-Based Application of Biomedical Hydrogels

Many of the early cell-compatible hydrogel systems lacked essential features, mainly those that can concurrently control material properties, biodegradation, and bioactivity. Material engineering design principles have overcome some of these limitations, and with these advances, new applications for cell-compatible hydrogels uncovered seminal scientific discoveries, as exemplified by the pioneering work of Discher and co-workers (46). Here, Engler et al. used polyacrylamide hydrogels modified with a thin collagen layer and mechanically tuned through radical polymerization cross-linking to document an important responsiveness of MSCs to their substrate elasticity (47). This research represented a major advance not only because it showed the profound influence of ECM mechanics on stem cell differentiation, but also because it ignited a pursuit toward identifying other material properties that can potentially control cell fate.

Discoveries based on hydrogel technologies have also affected translational research in regenerative medicine. Take, for example, the evidence presented by Blau and co-workers that documents how adult muscle stem cells on complaint hydrogel substrates can be propagated in vitro with higher efficiency, leading to better in vivo engraftment (48). These stem cells would normally lose their pluripotency and undergo massive cell death within the first weeks of culture on rigid plastic culture dishes (elastic modulus > 106 kPa); however, when cultured on chemically cross-linked bioactive PEG hydrogels tuned to 12 kPa elastic modulus (and containing immobilized laminin and cell-adhesive oligopeptides), the cells showed signs of self-renewal and were engrafted with substantially better integration in a muscle implant model.

The encapsulation of living cells in a hydrogel milieu is still one of the principal challenges of adapting these materials for further advancing tissue regeneration or studying cell biology using 3D cell culture techniques (49). Cell encapsulation has not lived up to its full potential using existing hydrogel technologies partly because each type of cell requires its own specific encapsulating microenvironment with cell-specific material properties and spatially controlled bioactive features. In order to begin to address these limitations, Anseth and co-workers combined photo-patterning and cytocompatible Click chemistries so as to improve spatial control of microenvironments in synthetic polymer hydrogels that support protease-mediated 3D cell invasion (33). Here, two orthogonal Click chemistries were used: a Huisgen cycloaddition reaction to cross-link a multi-arm PEG-tetra-azide with di-functionalized oligopeptides (protease-sensitive oligomers), and a secondary photo-activated thiolene coupling chemistry for site-specific biomolecule conjugation (for example, cell-binding oligopeptides). For proof-of-concept, they were able to show that only those niches engineered with a combination of protease sensitivity and cell-adhesive functionality enabled 3T3 fibroblasts to achieve spindled 3D cell morphologies.

Elucidating the specific bioactive requirements of highly specialized cell types (such as cardiac, bone, or liver) and designing hydrogels with complementary spatial features is still an arduous process. Furthermore, many advanced hydrogel designs will require more than just one or two of the aforementioned properties to mediate complex biological events such as cellular morphogenesis, differentiation, or self-renewal. Branching epithelial morphogenesis, for example—which is in part regulated by complex in situ geometries and soluble factors (such as transforming growth factor beta-1)—should be investigated with bioactive and spatially patterned 3D hydrogel constructs. To do this, Bissell and co-workers used micropatterned collagen hydrogels engineered with cellularized cavities in the form of curved tubules, bifurcated tubules, and fractal trees to show that geometry has an instructive effect on morphogenetic development (50). The important insight from such studies underscores the vast potential of using cell-compatible, custom-designed hydrogel systems in basic and applied scientific research.

Hydrogel Outlook

Hydrogels are already having a dramatic impact in many biomedical fields. For instance, doctors can treat injuries with experimental hydrogels, and stem cell researchers can make discoveries using hydrogels in place of tissue culture plastic (Fig. 1D) (51). As hydrogels improve through better design, their influence is likely to expand to other areas of science and medicine. For example, complicated drug regimens may be replaced with simple injectable microgels (Fig. 1B), systemic chemotherapy may be replaced with hydrogel-based cancer drug treatments (42), and viral gene therapy may be replaced with gene-delivery hydrogels. Beyond biology and medicine, hydrogels may affect the fields of biotechnology, pharmacology, and biosensors by providing solutions for large-scale protein production, drug-screening techniques, and individualized chemosensitivity assays. As an example, drug developers can test antitumor drugs in vitro by encapsulating tumor cell spheroids in protease-sensitive, bioactive hydrogels that more closely mimic physiological tumor growth conditions. Another unfulfilled opportunity for hydrogels is in biotechnology, in which cell manufacturing and protein production in large scales can be aided by encapsulating microgels that enable suspension cultivation of most anchorage-dependent cell types. Whether the objective is protein production or cell expansion, the use of encapsulating microgels would enable the transition from less efficient roller bottle cultures to efficient industrial suspension bioreactors. Thus, as the field moves to use new hydrogel designs, it gains the ability to develop sophisticated hydrogels for these and practically any other application.

References and Notes

  1. Acknowledgments: Support was provided by the European Union FP7 Program, the Lorry I. Lokey Interdisciplinary Center for Life Sciences and Engineering, the Russell Berrie Nanotechnology Institute, and the Singapore National Research Foundation. I thank I. Mironi-Harpaz, D. Yelin, O. Kossover, Y. Shachaf, R. Meller, and A. Wexler for helpful comments and support with the text and figures. I. Mironi-Harpaz and Y. Talmon are acknowledged for the scanning electron microscopy image in Fig. 2.

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