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Keywords:

  • hydrogels;
  • Tissue Eng;
  • cell biology;
  • cell-matrix interactions

Abstract

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

Cell culturing, whether for tissue engineering or cell biology studies, always involves placing cells in a non-natural environment and no material currently exist that can mimic the entire complexity of natural tissues and variety of cell-matrix interactions that is found in vivo. Here, we review the vast range of hydrogels, composed of natural or synthetic polymers that provide a route to tailored microenvironments.

1 Introduction

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

Every cell in the body is surrounded by an intricate network of fibers made of proteins and polysaccharides called the extracellular matrix (ECM). The ECM provides a physical support for tissues, and through variations in composition and organization of the macromolecular fibers an amazing variety of mechanical properties is obtained.[1] The ECM is not an inert medium, but strongly influences the survival, development, migration, proliferation, shape and function of cells that are embedded.[2] A range of physical cues from the ECM are integrated and converted into biochemical signals through a process known as mechanotransduction. It has been found that properties such as cell shape, matrix stiffness and porosity as well as surface topography, all affect cellular responses including those that direct stem cell fate.[3]

In contrast to embedding cells in an ECM in vivo, most research in cell biology has been dedicated to cell cultures in two dimensional (2D) monolayers, even though these conditions are not closely mimicking tissue physiology in vivo.[4] Superficially, the difference between two and three dimensions is that cells in 2D culture are growing on top of a surface, whereas cells in a three dimensional (3D) culture are fully surrounded by the matrix material. However, as shown in Figure 1, this change in dimensionality affects many aspects of the cellular environment.[5] In a 2D culture, cell adhesion and spreading are essentially unconstrained, but restricted to the horizontal plane, leading to forced apical-basal polarity in cells that are grown in monolayers. Although polarity is important for some cell types, such as epithelial cells,[6] for many other cells it is undesirable. As encapsulated cells can adhere in all three dimensions, their morphology remains more rounded, and no polarity is induced by cell-matrix interactions. Moreover, fully embedded cells are sterically hindered in their spreading and migration due to the confinement of the surrounding matrix. Cellular movement through an environment having pore sizes that are much smaller than the dimensions of the cells relies on proteolytic activity of the cell.[7] Migration through a confined environment in 3D largely relies on microtubule dynamics, in contrast to cells on 2D matrices, where migration depends on cycles of actin protrusion, integrin-mediated adhesion and myosin-mediated contraction.[8] Furthermore, limited diffusion of proteins and small molecules through a material results in gradients over the matrix and can cause oxygen tension within hydrogels.[9] Finally, the mechanical forces developing between adhering cells and their surroundings are very different under 2D and 3D circumstances. The polymer or glass substrates on which cells are traditionally grown to form a confluent layer is very stiff (Youngs modulus in the order of GPa) and coated with a continuous layer of matrix, compared to a much softer environment within hydrogels that sometimes consist of discrete matrix fibrils (Youngs modulus in the order of kPa).[5] As the stiffness of natural tissues is in the range of 0.1 kPa for brain tissue to 100 kPa within collagenous bone, and thus rather soft,[10] hydrogels provide a much more ‘natural’ mechanical environment for culturing cells.

image

Figure 1. Cell behavior including adhesion, spreading, polarity and migration is affected by the number of dimensions in which a cell is cultured. For example, cells respond in a different way on a 2D collagen surface (top) or being encapsulated within a 3D collagen matrix (bottom). Adapted with permission.[5] Copyright 2012, The Company of Biologists.

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In recent years, researchers have come to realize that new materials for cell culturing in both two and three dimensions are required. We need to understand how cells communicate with their 3D environment and which material characteristics are crucial to control these interactions. In this review we aim to provide an overview of the many different classes of polymer materials, both natural and synthetic that have been applied in 3D cell culturing. We especially hope to help researchers entering the field or those looking for new materials, to rapidly assess which materials have promising characteristics for their specific applications. Due to the large volume of literature on hydrogels for cell culture, we subdivided our review into hydrogels based on natural or synthetic macromolecules. We paid particular attention to key parameters as stiffness, porosity, ease of functionalization, biodegradability, and cell compatibility and we included a separate section on recent methods to control the microstructure of hydrogels. We conclude our review with a brief outlook how advanced materials can support the growth and differentiation of (stem) cells into desired lineages and how these designs can be exploited in the fields of tissue engineering and cell biology.

2 Considerations for Hydrogel Design

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

Hydrogel properties strongly depend on the macromolecular structure, methods of preparation, and degrees of crosslinking. Because cells can feel and respond to the physical properties of their environment, choosing the right material and corresponding properties is essential in cell biology research. This section provides an overview of several material properties that should be taken into consideration in the selection of a suitable hydrogel.

2.1 Stiffness

ECM stiffness plays a key role in regulating numerous cell functions including adhesion, migration and differentiation.[11] For example, lineage selection in naïve mesenchymal stem cells (MSCs) is sensitive to substrate stiffness.[10] Hydrogel stiffness can be varied by controlling the polymer concentration or crosslinking density. Soft polyacrylamide matrices (0.1–1 kPa) that mimic brain tissue are neurogenic, stiffer matrices (8–17 kPa) that mimic muscles are myogenic, and rigid matrices (25–40 kPa) that mimic collagenous bone tend to be osteogenic.[10] Lineage specificity by matrix elasticity has been shown in many studies (Figure 2), although recent work suggests that the mechanical feedback of the linkage between cell and substrate plays a role as well (Figure 2).[12] Fibroblasts and endothelial cells that adhere to a stiff surface (>2-3 kPa) show significant spreading and form more actin stress fibers than on a softer surface (<2-3 kPa).[13] Migration is affected by stiffness since anisotropic rigidity can induce directional epithelial growth and guide cell migration along the direction of strongest rigidity.[14, 15]

image

Figure 2. Differentiation of MSCs depending on hydrogel substrate stiffness. (A) Human tissue exhibits a wide range of elastic moduli depending on the type of tissue. The physical properties of in vitro hydrogels such as those made from poly(acrylamide) can be tuned with respect to the elastic modulus by controlling crosslinking density, to cell adhesion by covalent attachment of collagen I as well as to hydrogel thickness. Naive MSCs develop into various phenotypes depending on hydrogel stiffnesses similar to those found in natural tissue. (B) When comparing mimics of soft and stiff tissue microenvironments, cells anchor stronger to stiff than soft substrates, building focal adhesions and actin-myosin stress fibers, for example, in in vitro hydrogels which influence both cell physiology and cell fate. (C) MSC spreading and differentiation on PDMS substrates is not affected by the stiffness of the material, indicating that stem-cell fate not only depends on hydrogel stiffness, but also on the nature of the gel. (D) Substrate and stiffness-depending formation of focal adhesion points on PAAm and PDMS surfaces visualized by vinculin staining. Reproduced with permission.[10, 12, 51] Copyright 2006, Cell Press. Copyright 2009, AAAS. Copyright 2012, Nature Publishing Group.

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Hydrogel stiffness is described by its elastic modulus, which is a measure of the strain when stress is applied to a material and can be measured via rheology measurements or indentation experiments. In rheology, a controlled stress is applied to the sample of interest and the corresponding strain is measured, or a controlled strain is induced and the applied stress is measured. Since the force is applied to the entire sample, the resulting stiffness is an average value of the bulk stiffness.[16] In contrast, indentation studies on hydrogels measure elasticity at the mm to μm scale.[17] Samples are indented by small probes and the correlation between the applied force and resulting indentation is measured.[18] Despite the ability to probe hydrogel elasticity with high resolution, for example by using an atomic force microscope with a colloidal probe, inhomogeneities in the polymer network on the length-scale of the probe can result in significant deviations of elasticity data from those obtained by rheology. Moreover, hydrogel samples can adhere to the probe during retraction, affecting the outcome of the measurement if this is not corrected for in the chosen model.

The relationship between stress and strain in a hydrogel material can be described by the rubber elasticity model.[19-21] Both natural and synthetic hydrogels show a non-linear stress-strain relation, but many natural materials, including fibrin,[22] collagen[23] and actin[24] present a unique viscoelastic property, where the elastic modulus strongly increases the more the material is deformed.[25] This so-called strain-stiffening phenomenon could be an important factor in studying cell-matrix interactions as cells can locally and globally stiffen the gel.[22] Recently, a synthetic hydrogel of poly(isocyanopeptides) has been developed that possesses this fascinating physical property.[26]

On the contrary, fully swollen synthetic hydrogels usually consist of hydrophilic polymer chains that are fully extended in the aqueous media; additional stress can rupture the hydrogel which is why many synthetic hydrogel materials appear brittle.[27-29]

2.2 Biodegradability

Many natural matrices such as collagen or fibrin hydrogels are enzymatically degradable, enabling cells to degrade and remodel their surrounding environment.[30] As the pore size of many hydrogels based on natural materials is usually only slightly smaller than the size of a typical mammalian cell (several μm),[7] cells can spread, grow and migrate by remodeling the polymer matrix without degrading the material.[31] For hydrogels based on synthetic materials, which usually have pore sizes much smaller than the typical size of mammalian cells,[7] degradability is crucial since the cells require space to spread, grow and proliferate.[31] The importance of matrix degradability has been highlighted by Lutolf and co-workers in studies of cellular invasion of artificial matrices. It was found that fibroblasts could invade adhesive and degradable synthetic hydrogel and that invasion distances increased approximately linearly with culture time (Figure 3).[32] In a follow-up study, poly(ethylene glycol) hydrogels were described, either containing adhesion sites or protease sensitive crosslinkers, or both. Cell elongation was observed only in networks that contained both active adhesion ligands and degradable substrates. Digital time-lapse microscopy was used to quantify 3D cell migration (Figure 3). Within hydrogels that were functionalized with proteins for adhesion, cells extensively migrated, proliferated and formed interconnected cellular networks only when the scaffold material was sensitive for degradation (Figure 3).[33]

image

Figure 3. Cell migration and invasion in synthetic and naturally derived hydrogels. (A) Extensive cell migration inside PEG hydrogels leading to the formation of interconnected cell networks after long-time culture (upper image). Formation of blood vessels in subcutaneously implanted cell-laden hydrogels along tissue infiltration (lower image). (B, C) Synthetic hydrogel made by gelation of vinyl sulfone-functionalized PEG with RGD motifs as active cell binding sites and MMP-sensitive substrates that serve as elastically active crosslinking sites. (A) Time-lapse image of migratory behavior of a single cell within the PEG gel quantified by motile cell fraction. (D, E) Confocal microscopy images of GFP-AktPH-expressing fibroblasts migrating on the surface of a cell-derived matrix (CDM), inside a three-dimensional CDM (D) and inside 3D collagen matrix (E). Reproduced with permission.[32, 33, 35] Copyright 2003, Stanford University's HighWire Press. Copyright 2003, WILEY-VCH. Copyright 2012, The Rockfeller University Press.

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Synthetic materials are often used for migration studies as their degree of degradability can be controlled and degradable materials can easily be compared to non-degradable materials. Alternatively, the degradability of hydrogels based on pure natural polymers can be restricted synthetically. Khetan and Burdick have shown that an increased density of non-degradable crosslinks limited cell spreading in hydrogels that were patterned using sequential crosslinking. A primary crosslinking reaction formed a hyaluronic acid hydrogel with degradable peptide crosslinks; subsequently a UV light-induced radical reaction spatially introduced a non-degradable network. This property affected the morphology and eventually the fate of human MSCs cultured within these hydrogels.[34] Hydrogels based on natural materials are suitable when studying intracellular processes involved in 3D migration. For example, Yamada and co-workers investigated the role of PIP3 and Rac1 proteins in fibroblast migration on the surface as well as inside a cell-derived matrix, and inside a 3D collagen gel (Figure 3). They found that the mode of normal cell migration was dependent on the dimensionality of the matrix, the type of elasticity of the matrix and the activity of RhoA.[35]

2.3 Porosity

Regardless whether the matrix material can be degraded or remodeled by the encapsulated cells, the porosity of the material is an important environmental factor for cells and their viability, as large pores facilitate the efficient transport of nutrients, carbon dioxide and oxygen. In 3D, cell migration is fastest at pore diameters that match, or are slightly below, the diameter of polarized cells; migration speeds decrease in large pore size matrices due to the loss of cell-matrix interactions; but pore sizes much smaller than cell sizes trap cells in a physical ‘cage' and reduce cell migration.[7, 36]

Porosity refers to the maximum size of solutes that can diffuse in a hydrogel, which can be described by the mesh size (ξ). ξ quantifies the average linear distance between crosslinks and provides a measure of the space available between the macromolecular chains. The mesh size is an estimation of the average pore size assuming ideal crosslinking of the hydrogel molecules. However, a real hydrogel matrix usually contains larger and smaller pores due to non-crosslinked polymer chains contributing to an increase in effective pore size, while polymer chain entanglements decrease the pore size.[37]

The mesh size of hydrogels can be estimated by a number of experimental techniques, including mercury intrusion porosimetry (MIP), fluorescence microscopy employing dextran probes and scanning electron microscopy (SEM), each exhibiting distinct limitations.[38] While MIP is a rather standard method for porosity characterization, the hydrogel sample is kept under vacuum and thus the polymer chains are fully collapsed. Hydrogels samples for cryo-SEM measurements are also subjected to vacuum and are thus also at least partially collapsed.[39] Diffusion of fluorescent probes of precise molecular weight can be influenced by their interactions with the polymer host.

To obtain the pore size of a hydrogel in a theoretical way, one has to know the swollen polymer volume fraction. This para­meter is derived from the swelling ratio of a hydrogel from the dried to solvated state. The porosity of a hydrogel can then be predicted by the Flory-Rehner theory, which describes the thermodynamics of the equilibrium swelling of a crosslinked polymer network in a fluidic environment.[40] The theory describes the entropy of mixing of a solvent and a polymer network such as a hydrogel. The Flory-Rehner equation takes into account the average molecular weight between the crosslinks, which is closely related to the polymer volume fraction in the swollen state as well as the solvent-polymer interaction parameter. In complex situations when ionic interactions also play a role, a more sophisticated version of this equation can be used.[40] However, as the crosslinking of a polymer network is a random process, the polymer chain assembly inside the network is not uniform. By treating the polymer chains as Gaussian chains, their assembly can described by a Gaussian distribution function with sufficient accuracy.[41, 42] On the micro scale, porosity of a hydrogel can be controlled using solvent phase inversion whereas on the macro scale pore size can be varied by encapsulating inorganic particles that can be selectively dissolved after crosslinking the hydrogel matrix.[43] Other methods include stratifying hydrogel samples onto surfaces with controlled porosity as well as by growing salt crystals inside a hydrogel solution upon crosslinking.[44, 45]

The presence of pores greatly influences the water uptake capability of hydrogels. FTIR analysis has shown that the flexibility of polymer chains (and their solvation ability) in dense hydrogels is different from porous hydrogel materials.[44] The uptake of water can be calculated from the mass ratio of swollen and dry hydrogel.

3 Polymer hydrogels

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

Natural ECM-derived macromolecules have the advantage of inherent biocompatibility, but the formation of hydrogels from synthetic building blocks allows the design of artificial tissues with full control over the molecular structure (Figure 4). The latter has the potential to obtain materials with optimized properties and a more detailed understanding of the behavior of cells in complex 3D environments.

image

Figure 4. A variety of chemical and physical interactions underling hydrogel design. Bioactivity of hydrogels can be realized by grafting cell binding moieties, growth factors to hydrogels through carbodiimide chemistry (A). Degradation sites can be introduced by the incorporation of either hydrolytic (B) or proetolytic (F) functional groups into the backbone of the macromolecular network. Covalent bonds formed through the crosslinking will produce stronger and more stable network (G). Various interactions have to be taken into account when studying these cell-hydrogel systems, including hydrophobic interactions between the polymer chains (C), hydrogen bonding between polymer and cell-surface proteins (D) and ionic interactions between the polymer and cell membranes (E). Reproduced with permission.[212] Copyright 2012, AAAS.

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In general, hydrogel materials that are employed for 3D cell culture can be divided into promoting and permissive hydrogel materials.[46] Promoting means that they present natural binding sites cells can interact with, and that these interactions generate signaling cascades which in turn promotes cell migration, differentiation and remodeling of the gel matrix. Synthetic polymers form permissive hydrogels as they usually do not present natural binding sites that cells could make use of to probe their extracellular surroundings. Early examples of fully synthetic hydrogels only allowed for simplest cell adhesion and proliferation studies, as detailed below. The incorporation of integrin ligands in the matrix is crucial for cell adhesion and spreading.[47] In the absence of such ligands, some cells can excrete their own ECM to provide adhesion sites,[48] otherwise spreading is restricted, eventually leading to cell death.[49] Naturally occurring hydrogel precursors sometimes provide essential cues to allow for complex cell-matrix interactions that are also common for native cell tissue and thus do not require further functionalization. However, it is predicted that the future focus in the design of artificial ECM hydrogels will lie in highly customized cell matrices to perform specific functions or to investigate selective cellular behavior which fully naturally derived hydrogel materials cannot provide, as material stiffness, cell-matrix interactions and degradability are often coupled and cannot be individually tuned or studied.[50, 51] However, it is a challenging task to replicate the multiple chemical and physical functions arising from the large structural diversity the natural ECM provides.[52] Semi-synthetic hydrogels formed by a combination of synthetic building blocks and a second, natural polymer are promising to bridge this gap.[53] Although such hydrogel materials are less defined than purely synthetic analogs due to inherent variations in composition and quality of the natural source-based building blocks, 3D semi-synthetic systems will help to reassemble key characteristics of natural tissue by building biomimetic elements into a synthetic environment. In the sections below, we will summarize the main classes of hydrogels based on natural and synthetic building blocks. A more detailed summary of biopolymers and their respective applications, including materials that will not be discussed here such as cellulose derivatives, elastin and fibroin-based materials, can be found in other recent reviews.[54, 55]

3.1 Natural Materials

3.1.1 Fibrin

Fibrin is a naturally occurring protein-based material that has shown great potential in recent years as a tissue culture model to study cellular behavior in 3D[56] and as an injectable scaffold in tissue engineering.[57] Fibrin is formed by thrombin-initiated aggregation of insoluble polypeptide chains of fibrinogen into a network of fibrils.[55] The preparation and structural characterization of 3D cell-seeded fibrin hydrogels formed at different thrombin concentrations has been extensively studied.[58]

A major advantage of using fibrin for 2D and 3D cell culture studies is that fibrin naturally contains cell binding sites, thus does not require additional chemistry to introduce these sites in the hydrogel fabrication.[59] Fibrin specifically binds many growth factors as well as clot components, such as fibronectin, hyaluronic acid and von Willebrand factor.[60] Fibrin has two pairs of RGD sites and one pair of AGDV sites through which it can interact with cell-surface integrins.[61] As a natural material, the fibrin network presents unique viscoelastic properties that differ in many aspects from those of synthetic hydrogels. Fibrin gels display non-linear elasticity: the elastic modulus of fibrin strongly increases the more the material is deformed.[25] The poor mechanical properties of fibrin gels (with elastic moduli ≤ 0.1 kPa) were once considered a disadvantage, resulting in efforts of making modifications in order to obtain a stiffer gel for 3D cell culture.[15, 61] Recently, Huang et al. showed that soft 3D fibrin gels can control differentiation and proliferation rates of tumor cell subsets, providing a mechanical method for selecting tumor cells.[56] It has also been pointed out that the softness and large compliance of fibrin gel may make it uniquely useful as a matrix model for cells such as neurons that normally reside in very soft tissues such as the brain.[61] Fibrin gels have been seeded with chondrocytes and cultured in vitro or in vivo to generate new cartilage tissues,[62] and served as a scaffold for angiogenesis, making it suitable for cardiovascular tissue engineering.[63, 64] Recently 3D fibrin scaffolds have also shown promises in cord blood (CB) transplantation applications.[65] A further expansion of the materials properties is possible via the formation of collagen-fibrin interpenetrating networks, which have also been studied extensively for 3D cell culture.[64, 66]

3.1.2 Collagen

Collagen is the most abundant fibrous protein in mammals and the major component of extracellular matrix.[67] Collagen consists of three left handed polypeptide helices wound together into a right-handed triple helix, and at the ends of each helix peptide bonds crosslink adjacent helices. The resulting long fibrils can form bundles of much larger diameters,[68] the thickness of which determines the tensile strength of the connective tissue. The orientation of collagen fibrils in the extracellular matrix is controlled locally by the cells that produce them, and influences cell migration and polarization.[69] For cell culture studies, the materials properties of collagen gels have been improved by crosslinking the fibers using glutaraldehyde,[70] formaldehyde,[71] carbodiimide,[72] or genipin,[73] enzymatic crosslinking by transglutaminase[74] or tyrosinase,[75] and photo-crosslinking methods like UV irradiation.[76] Riboflavin has been used in clinical treatment to increase the mechanical strength of corneal collagen and collagen type I under UVA light.[77]

3.1.3 Gelatin

Gelatin is the polydisperse peptide mixture produced through the irreversible process of partial hydrolysis of collagen.[67] Gels with mesh sizes in the range of tens of nanometers are formed when collagen triple helices are partially reformed by cooling down gelatin solutions below 30 °C.[78] Gelatin gels lack thermal stability and require chemical crosslinking for applications as tissue culture materials. Nearly all crosslinking methods introduced for collagen also apply for gelatine.[79] Covalent crosslinking not only strengthens gelatin gels, but also inhibits degradation, which can be tuned by controlling the degree of crosslinking. Alternatively, gelatin can be photocrosslinked after functionalization with methacrylate or free thiol groups.[80, 81] Both 2D and 3D culture of endothelial cells in photo-crosslinked gelatin methacrylate (GelMA) showed good adhesion, proliferation, elongation and migration of cells, but in 3D, cell elongation and interconnective network formation was impeded by stiffer gels.[80] Covalently crosslinked thiolated gelatin-poly(ethylene glycol) diacrylate hydrogels were shown to support improved cell spreading and intercellular network formation compared to networks of physically incorporated gelatin in PEG hydrogel.[81] In contrast to the specific enzymatic degradation of collagen by collagenase, gelatin can be degraded by many proteases.[82]

3.1.4 Matrigel

Matrigel is very widely used in cell culture studies and is the solubilised mixture of basement membrane proteins extracted from Engelbreth-Holm-Swarm (EHS) mouse sarcoma, a tumor rich in extracellular matrix proteins.[83] Matrigel has a heterogeneous composition and is rich in laminin, and collagen type IV. It also contains heparin sulphate proteoglycans, entactin/nidogen, and various growth factors that resemble the cell membranes. Matrigel gels rapidly and irreversibly between 24 °C and 37 °C, and the gelling speed is dependent on the concentration and incubation temperature. Although Matrigel provides cells with the pseudo-natural environment and enables stem cells to maintain self-renewal and pluripotency, its composition varies between different batches, which makes it unsuitable for studies into the effect of specific proteins or specific protein combinations on growing cells.[84] Most epithelial and endothelial cell types exhibit good differentiated morphology in 3D Matrigel matrices.[83, 85] Matrigel with desired properties such as higher levels of collagen type IV or lower ability of cell differentiation/proliferation can be extracted from mice fed on a lathrogen and β-aminoproprionitrile diet.[83] However, matrigel promotes tumorigenicity and the growth of tumor cells in vivo, due to the multiple growth factors present in the matrix that promotes cell adhesion and proliferation.[86]

3.1.5 Polysaccharides

Polysaccharides consist of monosaccharides linked together via O-glycosidic linkages and cover a wide range of macromolecules with varying chemical functionalities as well as physical properties.[87] Polysaccharides are an important group of materials for cell culture and tissue engineering applications due to their versatile ability to form hydrogels. The mechanism of hydrogel formation depends on the chemical structures of the polysaccharides involved. In the absence of chemical modifications, hydrogen bonding (e.g., agarose) or inter-molecular electrostatic (ionic) interactions (e.g., alginate) lead to gelation. Synthetic chemistry has greatly broadened the scope of hydrogel formation pathways, making it possible to obtain hydrogels with tailor-made properties. In this section, several important types of polysaccharide material that have been extensively studied as hydrogel platforms for cell culture purposes are reviewed.

Hyaluronic acid. Hyaluronic acid (HA), or hyaluronan, is an immunoneutral polysaccharide consisting of alternating disaccharide units of [β(1, 4)-D-glucuronic acid-β(1, 3)-N-acetyl-D-glucosamine) linkages (Table 1). HA is the only non-sulfated glycosaminoglycan that is widely distributed throughout the body and has been in clinical use for over thirty years.[88] HA is an essential component of the ECM and plays an important role in wound healing, cell motility, angiogenesis, cellular signaling and matrix organization.[88-90] Moreover, HA can be degraded in the body by hyaluronidase, rendering it an interesting candidate for generative medicine.[91] Chemical modifications of HA to introduce desired functional groups and material properties have been reviewed extensively.[88] Typically, the carboxylic acid, hydroxyl, or N-acetyl groups are targeted for modification.

Table 1. Selected natural hydrogel building blocks and their functionalization
Chemical structureFunctionalizationReference
ImageRNH2 + CDI-activated RCOOHHou, Biomaterials 2005
  Ishihara, J. Biomed. Mater. Res. 2000
  Burdick, J. Biomed. Mater. Res. 2005
Image(Meth-)acrylationHennink, Adv. Drug Delivery Rev. 2002
 Tyramine-functionalizationDordick, Biomaterials 2002
  Langer, Biomaterials 2007
 ImageFeijen, Tissue Eng. Part A 2010 Kohane, Biomaterials 2007
   
ImageMethacrylationChan-Park, Biomaterials 2010 Elisseeff, Matrix Biol. 2008
   
 EDC/NHS activation for further coupling reactionsElisseeff, Biomaterials 2010
  Prestwich, Biomaterials 2005
  Wang, Soft Matter 2011
ImageEDC/NHS activation for further coupling reactionsBellamkonda, Tissue. Eng. Part A 1999
  Shoichet, Nat. Mater. 2004
  Shiochet, Biomacromolecules 2004
 -CHX-funtionalization with sulfo-SANPAHConnelly, J. Cellular Phys. 2008
  Rotter, J. Bio. Mater. Res. 2011
 ImageFouque, Biomaterials 2010
ImageEDC/NHS activation for further coupling reactionsMooney, Biomaterials 1999, Nat. Mater. 2005, Nat. Mater. 2010, Proc. Natl. Acad. Sci. USA 2009
  Wee, Adv. Drug Delivery Rev. 1998
ImageEDC/NHS activation for further coupling reactionsStorch, Biomaterials 2009
  Werner, Biomaterials 2009
ImageEDC/NHS activation for further coupling reactionsPrestwich, Biomacromolecules2002
  Feijen, Acta Biomater. 2010
  Engler, Biomateials 2011
  Discher, Integrative Biol. 2012
 ImageMatyjaszewski, Biomaterials 2008
 ImageBurdick, Nat. Comm. 2012
  Matyjaszewski, Biomaterials 2008
 ImageLanger, Biomaterials 2004
  Drager, Biomaterials 2013

Michael addition. Thiol-modified HA (HA-SH) is obtained through a carbodiimide-mediated reaction using hydrazide reagents containing disulfide bonds.[92] The modified HA can spontaneously form a gel through the formation of disulfide bridges, or crosslinked by a Michael addition reaction with difunctional or multifunctional electrophiles. Engler reported HA hydrogels based on HA-SH and PEG-diacrylate (PEG-DA) that showed time-dependent increases in stiffness.[93] Hydrogels obtained from HA-SH and PEG-vinylsulfone (PEG-VS) have shown great potential in cartilage tissue engineering.[94] To provide additional cell-binding sites, PEG tetra-acrylate (PEG-TA) was used to crosslink HA-SH and thiolated gelatin (Gel-SH). The obtained extrudable hydrogel supported the growth and proliferation of NIH 3T3 cells for up to 4 weeks.[95]

HA can also be chemically modified to present acrylate groups, and gelled by adding thiolated matrix metalloproteinase-degradable peptides (MMP-PEP-SH). The migration and proliferation of mouse MSCs in this type of hydrogels was studied as a function of the concentration of cell-binding ligand.[96]

Schiff-base reaction. Aldehydes can be introduced to HA by reaction with sodium periodate, resulting in the oxidation of vicinal hydroxyl groups to dialdehydes to give HA-aldehyde (HA-ALD) and the opening of the sugar ring. The in situ crosslinking of HA-ALD with adipic dihydrazide modified HA (HA-ADH) with the formation of hydrazone bonds happens fast, with gelling times ranging from a few minutes down to less than 1 min.[97] These hydrogels were also shown to form in situ in the presence of viable myocytes, opening new possibilities for myocardial tissue engineering.[98]

Huisgen cycloaddition (click chemistry). HA bearing a terminal azide (HA-AA) was crosslinked with HA containing a terminal alkyne (HA-pA) in the presence of a Cu(I) catalyst.[99]S. cerevisiae yeast cells were encapsulated directly into this click-gel and preliminary results showed 80% cell viability after 24 hrs. In a similar approach, HA-AA was crosslinked with the acetylene groups of propiolic acid (PA) modified gelatin, resulting in hydrogels supporting the adhesion and proliferation of chondrocytes.[100]

One should bear in mind that the use of Cu catalysts is not really practical for cell–related applications, especially when cells need to be encapsulated directly during the hydrogel formation. It is thus more desirable to use copper-free click chemistry[101] in order to prepare a more biocompatible and clickable hydrogel material.

Enzymatically crosslinked HA hydrogel. In situ crosslinking reactions using enzymes is a relatively mild approach for making hydrogels. Tyramine conjugates of HA (HA-TA) rapidly form hydrogels (within minutes) through the addition of horseradish peroxidase (HRP) and H2O2.[102] Similarly, hydrogels were obtained when the tyramine was substituted for dextran-tyramine (DA-TA). The resulted HA-DA-TA hydrogels exhibited potential for cartilage tissue regeneration applications as they were shown to be biocompatible when encapsulating bovine chondrocytes, enhancing chondrocyte proliferation and matrix production.[103] The elastic moduli of HA-DA-TA hydrogels could be tuned in the range of a few hundred Pa to tens of kPa, by varying the degree of TA conjugation, the polymer concentration, and the peroxide concentration in the crosslinking reaction. Although these enzymatically crosslinked HA hydrogels could in principle be formed in the presence of cells to serve as 3D matrices for cell culture, the use of HRP and H2O2 could be problematic in terms of biocompatibility.

Photo-crosslinking. Controversy remains about the potential damage to cells by radicals involved in the gelation reaction, but radical polymerizations are currently applied in certain clinical settings[104] and allow spatial and temporal control over the reaction as well as control over material complexity and properties.[105] Free radicals can be generated by heat, light or redox reactions. For 2D and 3D cell culture purposes, hydrogels are ideally generated using mild initiation conditions. For this reason UV initiated crosslinking (photo-polymerization) is by far the most commonly applied form of radical polymerization for generating HA hydrogel for cell studies. By carefully adjusting the initiator concentration and light (UV) intensity, cell damage is minimized and direct cell encapsulation has been realized.[106]

In order to prepare photo-crosslinkable HA hydrogels, HA is first chemically modified with methacrylate anhydride or glycidyl methacylate to present UV-crosslinkable methacrylates (HA-MA). Both methods produce HA-MA with tunable degrees of functionalization of up to nearly 100%. The elasticity of HA-MA hydrogels can be adjusted between ∼3 and ∼100 kPa by varying the concentration of the macromer or by a sequential crosslinking process.[107] The latter presents a novel and important way for the preparation of a dynamic cellular environment. HA-MA hydrogel itself provides poor adhesion for cells. However, the properties of the photo-crosslinked HA hydrogel can be tailored through the incorporation of integrin binding proteins, peptides as well as sulfate derivatives.[108]

Agarose. Agarose is one of the two main components of agar, a polysaccharide present in red algae. Agarose is a linear polysaccharide consisting of (1[RIGHTWARDS ARROW]3)-β-d-galactopyranose-(1[RIGHTWARDS ARROW]4)-3,6-anhydro-α-l-galactopyranose as the basic unit and contains a few ionized sulfate groups (Table 1).[109] The gelling mechanism of agarose resides in the formation of intermolecular hydrogen-bonds upon cooling, resulting in the aggregation of double helices by the entanglement of anhydro bridges.[110] The viscoelastic properties of agarose strongly depend on the degree of de-sulfation. Depending on its molar mass and solution concentration, agarose forms physical gels with tunable elastic moduli between ∼1 kPa and few thousand kPa, well in the stiffness range of natural tissues.[111] A lot of research has been carried out using agarose hydrogels to seed chondrocytes in order to fabricate tissue constructs[112] or to study chondrogenic differentiation of adult stem cells.[113] However, native agarose is bio-inert, lacking bioactive signals just like other polysaccharides. Physically blending-in proteins and peptide sequences into agarose is a suitable method to enhance cell attachment. Nomizu et al. reported the blending of laminin-derived peptides into agarose matrices for 3D cell culture.[114] Strong cell attachment to agarose was only observed in the case of stiffer matrices, probably because peptides leaked out of less dense gels. Using a similar mixing-in method, agarose-collagen co-gel was made as a 3D tissue culture scaffold for the study of the invasive behavior of human glioma cells.[115] Scanning electron microscopy revealed that agarose forms an intercalating web-like network between the entangled collagen fibers. The mechanical properties of agarose-collagen co-gels were varied and mesenchymal-to-amoeboid transitions in glioma cell motility were observed.

In contrast to the blending-in method introduced above, chemical modifications ensure that the cell-binding moieties are covalently bound to the hydrogel matrices. A general way to chemically modify agarose is through the use of hetero-bisfunctional crosslinkers. Bellamkonda et al. have reported the covalent attachment of laminin to agarose via reaction with 1,1′-carbonyldiimidazole (CDI).[116] The lamimin modified agarose was shown to significantly enhance neurite extension from 3D cultured cells. A popular crosslinker is the UV-activatable crosslinker sulfo-SANPAH. Various synthetic peptide sequences, including the RGD integrin-binding sequence[117, 118] and active peptide sequences from fibronectin[119] have been successfully conjugated to agarose in this way. The effect of ligand density on the chondrogenesis of bone marrow stomal cells (BMSCs)[118, 119] and chondrocyte redifferentiation[117] have been studied using these peptide-agarose 3D culture matrices.

Alginate. Alginate, also known as alginic acid, is a linear anionic polysaccharide containing homopolymeric blocks of 1,4-linked β-d-mannuronate and α-l-guluronate (Table 1). Alginate is generally regarded as biocompatible, and the gelation can proceed through different mechanisms. At acidic pH values below 3, alginate self-assembles into acidic gels by the formation of inter-molecular hydrogen bonds.[120] Alternatively, alginate forms a physical gel by cooperative binding with divalent cations such as Ca2+ or Ba2+. However, cation-crosslinked alginate gels have the disadvantage of un-controlled degradation due to the diffusion of ions under physiological conditions. The anionic nature of alginate makes it possible to form polyplexes with cationic polymer, such as chitosan and polylysine. The formation of polyplexes enhances the mechanical properties of alginate hydrogels. For example, human umbilical cord MSCs encapsulated in alginate hydrogel microbeads were incorporated into calcium phosphate cement and chitosan fiber paste to obtain a hybrid gel with enhanced mechanical properties.[121] In another example, a hybrid gel based on alginate/lactose-modified chitosan and CaCO3 powder was shown to promote chondrocyte growth and proliferation.[122]

Just like HA, native alginate provides poor adhesion to cells. Covalent linkage of integrin-binding peptides to alginate can promote cell adhesion to the matrix (Figure 5).[123, 124] Since alginate contains carboxylate moieties on the polymer backbone, well-established carbodiimide/reactive ester chemistry can be used to attach amine containing peptides. For tissue engineering purposes, RGD-alginate hydrogels were shown to promote the adhesion and differentiation of MC 3T3 pre-osteoblasts that could be useful for bone regeneration.[125] For the study of 3D cell-matrix interactions, Mooney et al. have utilized a FRET technique to quantify the number of bonds formed between cellular receptors and the adhesive peptides in the oligopeptide-alginate hydrogels. It was shown that the average number of bonds each cell establishes is actually larger than the number of peptides available to it, indicating that either cells are actively probing the gels or they are expanding their surface area to probe locations further than the nearest ∼10 nm distances.[126] It was also reported that the spacer length between alginate and the peptide plays an important role in the adhesion and proliferation of fibroblasts in both 2D and 3D gels in vitro.[127]

image

Figure 5. Murine mMSCs cultured in 3D alginate matrices with varying elastic modulus (elasticity: 5 to 110 kPa) and constant RGD density showing that cell and nucleus morphology are not strongly correlated to mechanics of 3D matrices. This can be visualized from the differential-interference contrasts (DICs) (A), the F-actin staining by Alexa Fluor 568-palloidin (B) and nulear staining by ehidium homodimer (C). Reproduced with permission.[124] Copyright 2010, Nature Publishing Group.

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Alginate gels generally present pore sizes between 5 to 200 nm.[128] Very recently, a macroporous (pore sizes in the dry state in the range of 16-200 μm) and injectable alginate cryogel has been developed as preformed scaffolds. To this end, pendant methacryloyl groups were first introduced into the alginate main chain through the reaction of alginate with 2-aminoethyl methacrylate (AEMA). Cell-adhesive RGD peptides were introduced by the covalent coupling of the peptide to polymerizable acryloyl-PEG-N-hydroxylsuccinimide (ACRL-PEG-NHS). These so-called cryogels were shown to present shape-memory properties and could withstand reversible deformations, providing a promising and robust way for cell engraftment in vivo.[129]

Chitosan. Chitosan is a linear polysaccharide composed of randomly distributed β-(1,4)-linked D-glucosamine and N-acetyl-D-glucosamine units (Table 1), structurally similar to glycosaminoglycans. Chitosan displays excellent biocompatibility and immunostimulatory activities, and has been extensively explored in various biomedical applications. Due to the pH sensitive nature of the weak polybase of chitosan, a subtle balance between the interchain electrostatic interactions, hydrogen bonding and hydrophobic interactions can result in spontaneous gelation by careful adjusting the pH of chitosan solutions. It has been shown that this type of chitosan hydrogel could support mouse 3T3 fibroblast culture.[130] β-glycerol phosphate disodium salt (GP) induces a sol-gel transition in chitosan solutions at physiological pH and temperature.[131] A few chitosan/GP hydrogels have been developed for clinical trials. A MSC seeded chitosan/GP hydrogel is investigated for clinical use for the regeneration of degenerated Intervertebral disc (IVD).[132] An injectable starch/chitosan/GP hydrogel was shown to maintain chondrocyte phenotype.[133] Collagen/chitosan/GP hybrid gel effectively induces osteodifferentiation in rat bone marrow MSCs. The primary amines on chitosan can be easily modified via carbodiimide chemistry for protein or peptide (e.g., RGD) presentation to facilitate cell adhesion.[134] In a similar way, chitosan can be modified with the photo-reactive azidobenzoic acid[135] to produce photo-crosslinkable hydrogels used for myocyte cell culture.[136]

Heparin. Heparin (HEP) is a highly sulfated glycosaminoglycan (Table 1) that is abundant in the liver.[137] Due to its exceptionally high negative charge density, HEP can interact with many functional proteins such as growth factors and extracellular matrix components through electrostatic interactions.[138] For 3D cell culture, RGD sequences are incorporated into heparin following EDC/NHS activation of the carboxylate groups after which a 3D hydrogel network is formed by crosslinking amine-functionalized star-shaped PEG with EDC/NHS activated heparin. It has been shown that the BFGF loaded PEG-heparin gels effectively trigger the propagation and differentiation of neural stem cells, making it a promising candidate for cell replacement treatment of brain diseases.[139] The same type of gel model was used for the 3D culture of HUVECs, whose proliferation appeared to be dependent on the mechanics and adhesiveness of the 3D matrices.[140] Star PEG-heparin hydrogels with decoupled physical and biochemical characteristics can be designed using mean field theory and have been tested for directing desired cell behavior.[141] Similar to HA, HEP can be modified to contain thiol functionalities available for thiol-ene click reaction with e.g., PEG-diacrylate, resulting in hydrogel networks that are potentially useful for cartilage tissue regeneration.[142] A heparin/dextran co-gel prepared by HRP mediated crosslinking of tyramine-HEP and tyramine-dextran showed short gelation time (minutes) and promise in supporting chondrogenesis for cartilage tissue engineering.[143]

Chondroitin sulfate. Chondroitin sulfate (CS) is a type of glycosaminoglycan comprised of alternating units of β-1,4-linked glucuronic acid (Glca) and β-1,3-N-acetyl-d-glucosamine (GalNAc) sulfated on different positions of the two different residues and classified into four major types: CS-A, CS-C, CS-D and CS-E (Table 1). Chondroitin sulfates are abundant in the ECM and provide mechanical resistance for certain tissues (e.g., cartilage) to external compression through the electrostatic repulsion among the sulfate groups.[89] Since CS is a major component of the cartilage ECM that presents many useful related biological properties and activities, it has been studied for cartilage regeneration applications.[144] UV-crosslinked hydrogels of methacrylated chondroitin sulfate and PEG-DA has been shown to support the chondrogenic differentiation of MSCs in 3D.[145] In order to avoid any possible adverse effect from the UV, similar hybrid gels based on CS and PEG-NH2 through EDC/NHS chemistry have been reported.[146] Similar to the modification of HA, CS could be modified with disulfide containing hydrazide to display reactive thiol groups, which is available for thiol-ene crosslinking to prepare hydrogels. A comparison showed that hydrogels based on CS-SH better supports neovascularization in vivo than those based on HA-SH.[147] A thermo-responsive hydrogel based on poly(N-isopropylacrylamide) (PNIPAAm) grafted chondroitin sulfate was fabricated by first functionalizing CS with hydrazine groups and the potential of using this hybrid gel for both 2D and 3D cell culture of HEK 293 cells has been investigated.[148]

Dextran. Dextrans consist of branched polysaccharides of repeating α-linked d-glucopyranosyl units of varying lengths and degrees of branching (Table 1).[149] Crosslinked dextran is commercially available as a microcarrier, either surface coated with collagen or diethylethanolamine (DEAE) derivated. Cell culture on crosslinked microcarriers (DEAE Sephadex A-50) has been in use since the 1960's.[150] Various methods have been developed to chemically crosslink dextran to form a hydrogel. Methacrylate derivatives of dextran (Dex-MA) can be gelled in the presence of an initiator,[151, 152] and these gels are biocompatible as evidenced by the proliferation of human fibroblasts.[151] The vicinal hydroxyl groups of dextran can be oxidized in a step-wise fashion to obtain dextran-aldehydes, (Dex-CHO) which can be crosslinked in situ with hydrazide-modified carboxymethyldextran (CMDex-ADH) to give a hydrogel that showed efficacy in preventing peritoneal adhesions.[153] Feijen et al. have reported the HRP-mediated crosslinking of dextran-tyramine (Dex-TA) conjugates to obtain hydrogels as scaffolds for cartilage tissue engineering.[143] Langer et al. described the preparation of bioactive 3D scaffolds based on dextran-acrylate (Dex-Acr) for enhancing the vascular differentiation of human embryonic stem cells (hESCs) by introducing RGD peptides through UV-crosslinking with Acr-PEG-RGD with or without vascular endothelial growth factor (VEGF165) loaded microparticles.[154] For the 3D culture of smooth muscle cells (SMCs), methacrylate and lysine modified dextran (Dex-MA-Ly) was UV-crosslinked with gelatin-methacrylamide (Gel-MA) and extensive cellular network formation was observed.[155]

3.2 Synthetic Polymers

3.2.1 PAAm and PAA

Poly(acrylamide) (PAAm) hydrogels have been used extensively in the 2D culture of cells, and in mechanotransduction studies as their mechanical properties can be tuned over a wide range from less than one kPa to several MPa.[156] In a seminal work, Discher and coworkers cultured MSCs on PAAm gels coated with collagen I, which showed a correlation between lineage specification into myoblasts, neurons and osteoblasts and increases in hydrogel matrix elasticity.[10] A particularly interesting feature of PNIPAAm is its lower critical solution temperature (LCST), that is close to 37 °C, and this has been utilized in the design of grafted surfaces containing PNIPAAm and derivatives[157] that allow cells to grow and spread under typical culturing conditions at 37 °C, but spontaneously detach due to enhanced surface hydration below the LCST. Under these conditions, the fully extended polymer network can also shield integrin binding sites on the hydrogel surface, hence cell-matrix affinity can be thermally switched.[158] The same principles hold for bio-conjugates of thermo-responsive PNIPAAm with gelatin and collagen.[159] As cells are completely detachable by lowering the substrate temperature below the LCST of the polymer, no conventional detachment agents such as trypsin or EDTA are required. In addition, copolymers of PNIPAAm and cell-adhesive polymers can be utilized to form thermo-responsive 3D artificial ECMs.[160] By forming a copolymer of PNIPAAm and a more hydrophilic second block such as poly(acrylic acid) (PAA), the local aggregation of PNIPAAm in the hydrogel is pushed back (Table 2),[161] and the temperature-dependent volume change behavior of PNIPAAm can be precisely controlled. In case of copolymers of PNIPAAm and acrylic acid, hydrolytic degradability is introduced as another parameter in the artificial ECM. However, to induce interactions between a fully synthetic PNIPAAm copolymer and (mammalian) cells, peptides that interact with cell-surface receptors, RGD and FHRRIKA, for instance, need to be covalently attached to the hydrogel.[162] Such peptide-functionalized hydrogels can provide an artificial ECM in which cells are viable for several weeks.

Table 2. Selected synthetic hydrogel building blocks and their functionalization
Chemical structureFunctionalizationReference
ImageImageLutolf, Adv. Mater. 2003
  Patterson, Biomaterials 2010
 ImageLee, Biomaterials 2008
  Hoffmann, Soft Matter 2010
  Kloxin, Adv. Mater. 2010
 ImageMetters, Polymer 2000
  Martens, Biomacromolecules 2003
  Benoit, Tissue Eng. Part A 2006
 ImageRice, Tissue Eng. Part A 2007
  Atzet, Biomacromolecules 2008
 ImageRossow, J. Am. Chem Soc. 2012
 ImageDeForest, Nature Chem. 2011
 ImageDeForest, Nature Mater. 2009
  Deans, Proc. Natl. Acad.Sci. 2012
 ImageAlge, Biomacromolecules 2013
 ImageYang, J. Controlled Release 2012
ImageImageRice, Tissue Eng. Part A 2007
 ImageAtzet, Biomacromolecules 2008
  Atzet, Biomacromolecules 2008
ImageImageNakayama, Langmuir 1998
 ImageNuttelman, J. Biomed. Mater. Res. Part B 2001
 ImageMartens, Polymer 2002
  Martens, Biomacromolecules 2003
 ImageSchmedlen, Biomaterials 2002
 ImageMillon, J. Biomed. Mater. Res. Part B2011
ImageImageChong, Small 2013
ImageImageStile, Biomacromolecules 2001
ImageImageJo, Biomacromolecules2001
ImageImageDargaville, Macromol. Rapid Comm. 2012
ImageImageMann, Cell. Mol. Bioeng.2010
3.2.2 PEG

Despite its inherent biocompatibility, simple synthetic scaffolds of PEG do not promote cell adhesion, and MSCs that are encapsulated inside such permissive gels often undergo apoptosis due to a lack of cell-matrix binding sites. PEG-based hydrogels can be formed under both physiological pH and temperature by conjugation of vinylsulfone-functionalized multi-arm PEG with cysteine-containing peptide sequences.[33] As these gels reach the gel point within minutes, their solidification kinetics enable in-situ curing in surgical applications. By employing cell-adhesion promoting RGD derivatives or peptides that are sensitive to cellular proteases, such as members of the matrix metalloproteinases (MMPs) family, PEG-based gels can be tailored for remodeling as well as cell traction studies, where cells proteolytically enter the 3D hydrogel network. Hydrogels seeded with primary human fibroblasts are remodeled into bone tissue within a month.[32] Further functionalization of the macromer starting material with growth factors facilitates subcutaneous implantation of the material for in vivo studies. Aside from tuning cell-induced enzymatic degradation by the concentration of protease-sensitive peptides, the biochemical properties of the peptides themselves can be altered to enhance degradation kinetics and thus cell-matrix interaction kinetics.[163]

The formation of PEG-based hydrogels is not limited to the use of PEG homopolymers. PEG-poly(propylene oxide) (PPO)-PEG, so called poloxamers that are commercially known as Pluronics, have been used to form semisynthetic 3D gels with fibrinogen[164] via UV-activated free radical polymerization or by thermal gelation at typical cell culturing temperatures (Table 2). However, the application of poloxamers as tissue scaffold may be limited to in vitro studies in the future, as several studies have found evidence for accumulation of Pluronics in animals.[165]

3.2.3 Poly(oxazolines)

Poly(oxazolines) (POZ) constitute a promising alternative to PEG[166] and are also applicable in the design of artificial 3D tissues.[167] Triblock copolymers of PEtOz and PLA form thermo-reversible gels with gelling properties depending on the block ratio and molecular weight of POZ. Human skin fibroblasts are viable and proliferating in 3D tissues made from PEtOz-PLA-PEtOz. To further tune the mechanical properties of the hydrogel, secondary crosslinking of POZs is achieved by copolymerizing PEtOz or PMOz with a photopolymerizable POZ derivative, 2-(dec-9-enyl)-2-oxazoline, which is then crosslinked under mild, cyto-compatible photo-polymerization conditions applying very low UV intensities of approx. 2 mW cm−2 (Table 2).[168] In addition, POZs can be applied to form micro-patterned surfaces with confined cell receptor membrane proteins, thus providing spatial control over cell adhesion.[169]

3.2.4 PVA

Poly(vinyl alcohol) (PVA) is a biocompatible hydrolysis product of poly(vinyl acetate) that offers great flexibility in the design of hydrogel precursors (Table 2).[170] Although PVA shows only little adhesion to cell membrane proteins, its multiple hydroxyl groups can be easily functionalized with esters and photo-polymerizable vinyl groups, as well as RGD(S) or fibronectin to form 2D hydrogel surfaces for cell spreading studies (Figure 6).[171, 172]

image

Figure 6. Expression of beta tubulin in PEG-PLA hydrogels. Immunofluorescence analysis of cells in PEG hydrogel culture using antibodies against b-tubulin and nestin. (B) Radial artery cells encapsulated inside physically crosslinked PVA hydrogels. To improve cell adhesion inside the hydrophilic polymer network, fibronectin is blended into the hydrogel as well as covalently attached via two different approaches (from left to right). (C) RGD and PHSRN presentation as an active way to introduce cell-binding motifs into otherwise bio-inert supramolecular hydrogels. Reproduced with permission.[172, 180, 187] Copyright 2007 and 2011, Wiley. Copyright 2005, Nature Publishing Group.

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In general, hydrolytically labile ester bonds can be employed to trigger the degradation of polymer networks at neutral pH and alter diffusion rates inside the material as well as its physical properties over time. By functionalizing PVA with methacrylamido acetaldehyde dimethyl acetal or mono-2-(acryloyloxy)ethyl succinate, photo-polymerizable PVA-acrylate esters are obtained, with degradation time and swelling rate of the synthetic hydrogel easily adjusted by the molecular weight of the PVA backbone and the number of hydrolytically degradable crosslinking points.[173] The overall degradation time, which is typically measured by the percentage of mass loss, is tunable from less than a day to more than a month. Further control over degradation profiles is achieved by copolymerizing PVA with acrylate-functionalized PLA-PEG-PLA yielding hydrogels with bimodal degradation behavior, which can be employed for cartilage tissue engineering.[174] Upon photo-polymerized encapsulation of chondrocytes into the hydrogel, PVA is released via chondrocyte medium-induced gel degradation, and the compressive strength increased during several weeks due to formation of cartilage-like tissue.

3.2.5 PLA, PLGA and PCL

Poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA) and poly(ε-caprolacton) (PCL) are common polymers for designing hydrolytically biodegradable hydrogels as well as improving the degradability of hydrogels made from other macromers, including PEG (Table 2).[89, 175] Examples of PEG-polyester-based hydrogels include photo-crosslinkable PLA-PEG-PLA,[176] and PEG-PCL.[177] Although PLGA degrades faster than PLA due to the more hydrophilic glycolide block, losing 90% of its mass after 10 weeks compared to 20% in the case of PLA, it is suitable for long-term studies on differentiation and proliferation of cells.[178] Many studies on PLA and PLGA have focused on the design of artificial ECM in which osteoblastic cells are grown and manipulated for skeletal tissue replacement or chondrocytic cells for cartilage regeneration. However, these gels often degrade too fast, thus neo-cellular tissue cannot fill-out the volume previously occupied by the hydrogel in time and tends to form structural defects that diminishes its load-carrying capacity. In other cases, the hydrogels degrade too slowly and the newly grown cells cannot fully replace the damaged tissue in a timely manner. Photo-copolymerization of hydrolytically/enzymatically degradable PLA-PEG-PLA with non-degradable partners such as PEG dimethacrylate enables precise control over the hydrogel degradation kinetics by adjusting the mass ratio of both polymers.[179] As the degradability of the hydrogel also influences the diffusion of cell-secreted minerals and collagen in the artificial ECM, the homogeneity of mineralization as well as overall mineralization can be precisely tuned by increasing the amount of PLA-PEG-PLA. Like most synthetic macromers, PLA-PEG-PLA copolymers lack adhesion moieties and structural features of natural ECMs cells can interact with. To transform a fully synthetic hydrogel into an active matrix and stimulate the proliferation of cells into the PLA-PEG-PLA polymer network, samples can be doped with ECM molecules and growth factors post hydrogel formation (Figure 6).[180]

The degradation kinetics of semi-crystalline, more hydrophobic PCL is considerably slower than PLA and PLGA.[181] Therefore FDA-approved PCL is a promising candidate for the design of synthetic ECMs in the area of bone recovery/replacement and tissue replacement in general, as it provides the mechanical properties to resist mechanical stress during long-term recovery of bone tissues,[43] and degradation rates that are similar to the growth rates of natural tissues. The osteoconductivity of the PCL hydrogel matrix is further improved by conducting cell growth studies in simulated body fluid instead of conventional cell culture media or by adding hydroxylapatite, which supports remodeling of the artificial ECM. The precise control over mechanical properties and porosity of synthetic hydrogel matrices also facilitates development of tailored PCL-based scaffolds for cartilage regeneration.[182] Elastic hydrogels that completely recover from applied strain are formed by copolymerizing PCL with PLA. Mechanical stimulation is efficiently relayed through the porous artificial ECM to the implanted chondrocytes, which continue to differentiate even under severe mechanical stress.

While it should be noted that the non-enzymatic degradation of tissue matrices is a rather uncommon motif in vivo, hydrogels containing ester blocks are also susceptible to enzymatic degradation, for example in the presence of serum albumin,[183] peptidases that are able to degrade the glycolic acid-block in PLGA[184] or lipases that degrade PCL.[185] In the latter case, addition of lipase to the artificial ECM at defined time steps introduces precise temporal control over hydrogel degradation and cellular response.[177] Negative effects of the biodegradation of polyester hydrogels in vitro and in vivo are often not detailed. Yet, pH measurements of culture media in hydrogel scaffolds show that cell migration and viability is inversely proportional to the matrix degradation rate.[45] Supramolecular ureido pyrimidone (UPy)-functionalized oligocaprolactone membranes can be precisely engineered to accommodate human renal epithelial cells and form a long-term stable, bioactive kidney membrane.[186] By simply blending the hydrogel with supramolecular linker-functionalized bioactive molecules, the degree of functionalization as well as the number of different cell adhesion promoters can be easily controlled in a modular approach (Figure 6).[187]

3.2.6 OPF

Oligo(ethylene glycol) fumarates (OPFs) combine the well-known biocompatibility of PEG with a high ester-group density thus rendering the polymer degradable by non-specific hydrolysis and enzymes. Jo and coworkers[188] have successfully synthesized OPFs, using post-modification with short peptide sequences (GRGD) to provide a basis for modulating cell-polymer matrix interactions (Table 2). Subsequent photo-crosslinking of the fumarate double bond with poly(propylene fumarate) (PPF) yielded polymer gels with mechanical properties suitable for orthopedic implants.[189] However, the rather harsh crosslinking conditions involving the use of methylene chloride and DMF as well as extensive UV exposure of approx. 30 min. prevent pre-encapsulation of viable cells inside pre-polymer solutions.

4 Libraries of Polymers and Hydrogel Micro-Structuring

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

The behavior and function of cells is greatly influenced by the mechanical and chemical properties of the surrounding microenvironment. Despite remarkable progress towards efficient stem-cell therapies as well as tissue regeneration and replacement, investigating the cellular response of a large variety of cell types or materials as well as their characteristics in artificial ECMs using conventional culture plates is time consuming and limits progress towards tailored cellular environments. A broader insight into the complex orchestra of chemical and physical sensors that are controlling the cell's response to its external environment can be provided by high-throughput micro-engineered platforms enabling rapid cell-based screening of combinatorial libraries of hydrogels,[190] as well as by micro-engineered hydrogels, so called microgels.

Considering the wide variety of synthetic or natural polymers that can be used, alone or in combination, with other polymers, used to form tailored hydrogels for cell encapsulation and cell studies, a combinatorial approach enables efficient screening of large numbers of novel photo-crosslinkable and biodegradable hydrogel materials.[191] By forming microarrays of hydrogels of several thousand of polymer combinations using robot printing, this approach helps to reveal unexpected relations between hydrogels and soluble factors, as shown by screening the ability of acrylate hydrogels formed from combinations of 24 different monomers to support human embryonic cell attachment, spreading and differentiation.[192] In a similar approach, libraries of more than 3,000 biodegradable polyesters have been screened to support adhesion and spreading of hMSCs (Figure 7).[192, 193] Utilizing micro-arrays of combinations of growth factors and fibronectin, cells that are grown in monolayers on these spots can even be captured in single states of differentiation demonstrating the unrivalled control that can be achieved combining artificial ECM and micro-environment design.[194]

image

Figure 7. Formation of hydrogel microstructures by robot printing and soft lithography. (A) High-throughput screening of more than 3400 combinations of cell-hydrogel interactions using a library of PLA, PEG, PCL and PLGA. One million human MSCs are seeded onto the microarray of polymer spots. Actin is labeled green, the cell nucleus and DNA, respectively, is labeled blue. (B) Soft-lithographic production of artificial microenvironments (niches) made form PEG hydrogels for probing stem cell fate depending on cell concentration (right) quantified by cell shape, differentiation as well as proliferation. Reproduced with permission.[193, 196] Copyright 2005, Elsevier. Copyright 2011, Nature Publishing Group.

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In general, sample spotting onto flat surfaces is limited to studying adherent cells as it makes use of arrays of ECM molecules and growth factors that are printed on non-physiological surfaces with no control over physical properties of the substrate. Micro-patterned hydrogels can be applied to study the effect of cell-cell interactions as well as hydrogel stiffness on the fate of single MSCs. High-throughput screenings of 3D environments for human MSCs have been used to gain a deeper understanding on how the activity of cell binding sites is affected by the choice of hydrogel polymer. Thereby, cell density, stiffness, porosity of the hydrogel and the concentration of bioactive ECM components was combinatorially modulated in artificial cell niches to allow for detailed cell-material interactions (Figure 7).[195, 196]

Studies on fibronectin- and laminin-derived adhesion ligands indicate that cell adhesion is more pronounced in non-degradable PEG diacrylate (PEGDA) hydrogels than in hydrolytically cleavable DTT-crosslinked PEGDA. This indicates that the cell adhesion process is more complex and dependent on more parameters than previously predicted. Even small libraries of hydrogel candidates are able to efficiently contribute to the identification of promising hydrogel candidates such as for scaffolding bone tissue regeneration screening degradation behavior and physical properties as well as the incorporation of growth factors that support bone cell invasion.[197]

The formation of micro-patterns on surfaces usually only allows for a 2D view on cell-matrix interactions, whereas cells generally require a 3D and heterogeneous microenvironment to generate a response as found in complex natural tissues. To overcome the limitations of homogenous 3D hydrogels, photo-sensitive hydrogels can be heterogeneously patterned using two-photon absorption laser scanning lithography.[198] Predefined ‘tracks' for 3D migration of human dermal fibroblasts can be formed by light-induced micro-patterning of fibronectin-derived RGDS and CS-1 with high resolution at the μm to mm-scale. Cellular response can be precisely tuned by the concentration of binding sites which is dependent on the exposure time of the hydrogel to induce functionalization. In addition, biocompatible, yet non-degradable hydrogel materials such as PEG can be rendered degradable towards collagenase secreted by encapsulated cells by incorporating collagenase-sensitive peptide sequences such as the LGPA protein with the same method.[199] This approach facilitates the removal of the hydrogel matrix such as in surgical tissue engineering and increases the availability of cell-binding sites.

In general, 4-arm PEGs can be crosslinked under mild, cell-compatible conditions employing the Huisgen cycloaddition between an azide and an alkyne.[200] This approach can be utilized to form hydrogels from PEG-hydrazides and di-functionalized peptides that contain cyclooctynes and photo-reactive allyloxycarbonyl groups that allow for secondary crosslinking and functionalization of the hydrogel matrix via UV-induced thiol-ene click chemistry (Figure 8).[201] As allyl esters are not reactive under Michael-addition conditions, no non-specific crosslinking occurs. The motif of orthogonal, copper-free, non-cytotoxic alkyne-azide reaction can be extended towards the introduction of photo-conjugatable and photo-cleavable groups allowing polymer network formation and degradation using visible (thiol-ene reaction) and UV light (benzyl ether photo-degradation). The presentation of cell binding sites within the hydrogel matrix depends on the photoinitiator concentration as well as exposure time and is spatially controlled in 3D with micrometer-scale resolution (Figure 8).[101]

image

Figure 8. 3D micro-patterning of hydrogels. (A–C) Secondary patterning of polymerized hydrogels. (A) Two-photon laser scanning photolithography precisely locating RGDS in collagenase-sensitive PEGDA hydrogels. Fibroblasts migrate only within the highly defined geometry. (B) Photo-patterning of three different fluorescently labeled peptide sequences in space inside an PEG tetraazide-peptide cyclooctine-clicked hydrogel with free photo-crosslinkable allyl esters. (left). Formation of complex 3D patterns by scanning the focal point of a near-infrared laser (right). (C) Independent from a temperature-induced hydrogel formation via click coupling, additional thiol-containing adhesion peptides are incorporated via photo-coupling employing visible light. The presentation of fluorescent RGD within the hydrogel matrix depends on photo-initiator concentration as well as exposure time and is spatially controlled in 3D with micrometer-scale resolution which allows to control the outgrowth of fibroblast cells in a 3D PEG-peptide hydrogel by forming RGD-functionalized hollow microchannels. (D) Adhesion and growth of MSCs on the surface of hydrogel beads microfluidically fabricated from branched PEG-VS and PEG-SH. DAPI and phalloidin are used to stain the nuclei and the cytoskeleton, respectively. Reproduced with permission.[101, 199, 201, 205] Copyright 2008, Elsevier. Copyright 2009 & 2011, Macmillan Publishers Limited. Copyright 2011, RSC Publishing. Copyright 2013, ACS Publications.

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Taking these simple conjugation processes one-step further, Shoichet et al. developed a versatile approach to spatially control ligand binding to agarose in a mild way using two-photon photochemistry. S-(2-Nitrobenzyl)cysteine (S-NBC) can be grafted to the hydroxyl groups of agarose following CDI activation.[202] As shown in Figure 9, upon UV photo-irradiation, free functional sulfhydryl groups are generated, which react with sulfhydryl-reactive biomolecules through maleimide moieties, for instance.[203] The degree of immobilization can be controlled by the energy dose and distinct biochemical channels can be created in the 3D agarose matrix to guide cell migration.[204]

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Figure 9. Simultaneous immobilization of stem cell differentiation factors SHH and CNTF in agarose hydrogels. Maleimide-barnase (black circle) is immobilized using two-photon photochemistry and a femtosecond laser. The hydrogel is then washed in buffer to remove unbound maleimide-barnase. Next maleimide-streptavidin (orange square) is immobilized using two-photon irradiation followed washing. The fusion proteins barstar-SHH (green circle) and biotin-CNTF (red square) are soaked into the gel and bind to barnase and streptavidin, respectively. After washing out excess protein, both SHH and CNTF are simultaneously and independently immobilized in three dimensions. Confocal images show the simultaneous 3D patterning of biotin-CNTF-633 and barstar-SHH-488. Reproduced with permission.[202] Copyright 2011, Nature Publishing Group.

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While the examples described above make use of bulk hydrogels to study cell behavior, the combination of bio-orthogonal hydrogel formation and functionalization with materials engineering using droplet microfluidics enables the fabrication of 2D and 3D tissue in a bottom-up approach with defined architecture and composition.[110] Michael addition-chemistry can be employed to form cell-compatible, protein-functionalized PEG hydrogels with tailored stiffness to influence the growth of stem cells on the microbead surface (Figure 8).[205] Another advantage of microfluidic hydrogel fabrication over conventional bulk polymerization is that 3D micro-environments of varying elasticity can be formed by simply tuning the ratio of two hydrogel precursors online which facilitates high-throughput formation of hydrogel beads with up to 35-fold variation of elasticity in case of agarose and is thus suitable to efficiently evaluate stem cell fate versus hydrogel elasticity.[206]

As discussed above, cell-compatible polymerization of hydrogel precursors is an important prerequisite to attain high cell viability rates during long-term studies. This can be achieved by bio-orthogonal thiol-ene-click hydrogel formation from hyper-branched polyglycerols and PEG which is catalyst- and radical-free.[207]

While bulk gels are usually employed to investigate one cell type at a time, microfluidically prepared cell-laden hydrogel beads can be assembled into complex patterns not only to study cell-matrix interactions within single gel beads, but also to investigate cell-cell communications (Figure 8).[208] In addition to using single-emulsions droplets to encapsulate cells in well-defined hydrogel beads, monodisperse double emulsions can be as well employed to form cell-containing hydrogels.[209] Using alginate, gelation is induced upon removal of a mineral-oil layer that initially separates the double-emulsion's alginate core from an outer calcium solution.

Aside from microfluidically prepared droplet templates that are employed to form cell-laden hydrogel beads, microfluidic devices themself can serve as a platform to study cells in hydrogel matrices, whereas the conventional device building material poly(dimethylsiloxane) (PDMS) can be replaced by agarose, for instance.[210] By employing soft lithographic fabrication techniques, cell-laden microfluidic devices are fabricated where nutrients and oxygen are supplied via a microchannel network with improved resistance under compressive loading.[211]

5 Outlook

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

Virtually every cell in the body is exposed to an extracellular matrix (ECM) to which they adhere via a number of different cell surface receptors. This interaction allows the cells to sense mechanical cues from the ECM as well as degradability, porosity and distribution of binding sites and respond in an appropriate manner (for example, changes in cell shape and size and responses such as differentiation and proliferation). Thus, as a key component of the stem cell niche, the ECM is not just an inert scaffold, but instead can profoundly influence cell fate choices. In order to maximize the effectiveness of cell-based therapies – whether stimulating expansion of endogenous cells or transplanting cells into patients – it is essential to understand the environmental (niche) signals from the ECM that regulate cell behavior. Polymer hydrogels, composed of synthetic or natural polymers, have provided a powerful platform to precisely tune mechanical properties, degradability, biocompatibility, and porosity in accordance with demands of specific cell types. Many of the natural polymers have the advantage that they are biocompatible and contain many sites for further chemical modification. However, synthetic polymers provide much more precisely defined materials and are bio-inert. Although the latter requires further modification with cell adhesion sites, it prevents unwanted cell stimulation. However, it is also clear that thus far no single material has emerged (and will emerge) that suits every application, whether it be tissue engineering or fundamental cell biology research. In this review, we have attempted to provide an extensive overview of the polymeric building blocks, crosslinking strategies and chemical modifications applied to produce hydrogels with defined properties. Where possible, we have linked those properties with the cellular responses that were reported in the literature, but it is clear that there are significant gaps in our knowledge how to engineer the 3D environment. Much of our knowledge on cell adhesion and cell spreading derived from studies on 2D cell cultures evidently cannot be transferred directly to a 3D setting. The materials chemistry for the 3D environment is well-established; the first priority now will be to learn the fundamental rules for guiding cell behavior in 3D matrices, using the broad range of materials available.

Acknowledgements

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies

This article is part of an ongoing series celebrating the 25 th anniversary of Advanced Materials. Authors J.T. and Y.M. contributed equally to this work. J. T. is a Feodor-Lynen fellow of the Alexander von Humboldt Foundation. S. M. was funded by a CSC Cambridge International Scholarship. Work in the Huck group is supported by a European Research Council (ERC) Advanced Grant (246812 Intercom), a VICI grant of the Netherlands Organization for Scientific Research (NWO), and by funding from the Ministry of Education, Culture and Science (Gravity program 024.001.035).

Biographies

  1. Top of page
  2. Abstract
  3. 1 Introduction
  4. 2 Considerations for Hydrogel Design
  5. 3 Polymer hydrogels
  6. 4 Libraries of Polymers and Hydrogel Micro-Structuring
  7. 5 Outlook
  8. Acknowledgements
  9. Biographies
  • Image of creator

    Dr. Julian Thiele is a postdoctoral researcher at Radboud University. He received his B.Sc. from Lund University (Sweden, 2008) and his diploma in Physical and Polymer Chemistry from Hamburg University (Germany, 2008). Funded by grants of the German Academic Exchange Service (DAAD) and the Fund of the Chemical Industry (FCI), he obtained his Ph.D. from Bayreuth University (Germany, 2011) for which he was awarded the Culture Award Bavaria. Being a Feodor-Lynen fellow of the Humboldt foundation, his current research focuses on microfluidically prepared artificial cells and functional hydrogels.

  • Image of creator

    Dr. Yujie Ma received her B.Sc. in Polymer Science and Engineering (2001) from Shanghai Jiao Tong University (China), M.Sc. in Chemical Engineering (2003) and Ph.D. in Macromolecular Nanotechnology (2008) from the University of Twente (Netherlands). In 2011 she became a postdoctoral fellow in the group of Physical Organic Chemistry in the Radboud University of Nijmegen. Her current research is focused on the study of cell-matrix interactions in 3D and micro-tisue engineering based on hydrogels.

  • Image of creator

    Wilhelm T. S. Huck is a Professor of Physical Organic Chemistry. Prior to taking up his position at the Radboud University Nijmegen, he was Professor of Macromolecular Chemistry and Director of the Melville Laboratory for Polymer Synthesis at the University of Cambridge. His research is focused on the physical biology of the cell and aims to elucidate, using model systems and living cells, the influence of the special nature of the cellular environment on complex reaction networks in cells.