Hyaluronic acid is part of the ECM [13], and a naturally occurring linear polysaccharide that is abundant in the vitreous and synovial fluid and plays important roles in wound healing [14], cell differentiation and cell motility [15]. Research on stem cell culture indicates that function of the HA is essential also to the control of self-renewal of human embryonic stem cells [16]. It is highly biocompatible, suitable for modification with drugs and other effector molecules [17]. HA and its derivatives have been used to compose drug delivery systems consisting of a wide variety of drugs and cell encapsulation [11,17].

HA is composed of beta (1–4) linked 2-acetamide-2-deoxy-D-glucose and beta (1–3) linked D-glucuronic acid, and is known to be non-antigenic, noninflammatory and generally non-tissue reactive [18]. HA can be degraded by HAse in vivo, which is ubiquitous in cells and in serum [19]. For the purpose of in vitro HA hydrogel study, different methods have been developed for measuring the hydrogel degradation, by such means as monitoring the release of uronic acid (a degradation component of HA) from a matrix, or monitoring the loss of weight [20,21,22].

The crosslinkage of HA hydrogel can be formed as a result of polyvalent hydrazide cross-linking, disulfide cross-linking, photo-crosslinking, or enzymatical crosslinking [11,17,22,23]. Pure HA hydrogel is a nonadhesive hydrogel for cells, and the promotion of cell adhesion can be achieved by mixing pure HA hydrogel with gelatin microparticle [24,25].

Another type of polysaccharide, chitosan has been widely used for biomedical applications, such as cell culture platforms [26,27], drug carriers [28] and non-viral gene delivery [29]. Chitosan is derived from chitin with a linear structure, consisting of β-(1–4)-linked 2-amino-2-deoxy-D-glucose and 2-acetamido-2-deoxy-D-glucose units [30]. Chitosan can be easily modified, usually via primary amine groups [31]. However, Chitosan can only be dissolved in an acetic environment, which limits its applications in injectable cell culture scaffolds. A common solution is to partially N-succinylate chitosan to make it dissolvable in a neutral solution [32]. Chitosan has excellent biodegradability as well [28]. Lysozyme is used to degrade chitosan for in vitro degradation evaluations [33]. Chitosan polymer chains can be crosslinked by glutaraldehyde or genipin to form chitosan hydrogel [34,35]. Other methods have also been developed to from chitosan hydrogels, such as maleic chitosan [36].

Alginate is block polymer composed of (1–4)-linked b-D-mannuronic acid (M units) and a-L-guluronic acid (G units) monomers, the ratio varying between M and G units [37]. Alginate polysaccharides covalently modified with RGD-containing cell adhesion legends are widely used for the settlement and attachment of cells. In addition, calcium alginate hydrogel surfaces coupled with GRGDY peptides can be fabricated to achieve cellular interaction [38]. They are well known for their uses in mineralized polymeric matrices, which justifies their application in bone tissue regeneration [39]. However, it is difficult to have alginate degraded in vivo and in a short time [40].

Collagen is naturally occurring proteins in the form of elongated fibrils and found in mammals [41,42]. It is the main component of connective tissues such as tendons, ligaments and skin, and is up to 35% of the body protein content, i.e., the most abundant protein in mammals [43]. Gelatin is collagen that has been hydrolyzed under basic and high temperature conditions [44]. Collagen or Gelatin can be degraded by enzyme collagenase in vivo [45]. The methods used to form the collagen or gelatin hydrogel are quite similar to those for HA, such as disulfide cross-linking, photo-crosslinking, and enzymatical crosslinking [46,47,48].

Synthetic biomaterials are favored as scaffold materials because their physical and biologic properties can be modified and they can be reproduced in similar and large quantities. The major classes of synthetic biomaterials are the glycolic acid derivatives, lactic acid derivatives, and other polyester derivatives [9].

Widely used in human medicine, PEG resists protein adsorption and is therefore endowed with unique nonfouling properties because of its nonadhesivity towards proteins and cells [8,49].

PEG is non-ionic and soluble in water. PEG diacrylate is produced as a result of the modification of PEG with acrylate to get carbon double bone which will serve as a gelation functional group [36]. Application of PEG or its diacrylate derivative (PEGDA) to tissue engineering [49,50,51,52] is limited by their inability to support cell spreading due to their being non-adhesive to protein and the absence of cell adhesion ligand [49,53].

Pure PEG (molecular weight <20,000) is chemically stable under physiological conditions [40]. However, after modification some polymers based on PEG can be biodegradable, as instanced by oligo(poly(ethylene glycol) fumarate) (molecular weight ~10,000) [54].

The many crosslinking methods for PEG hydrogels, such as the thermal radical initiation system [54] and disulfide crosslinking system [23], have been extensively investigated so that nowadays both chemical hydrogels and physical hydrogels are available [55].

Similar to PEG, PVA is another type of hydrophilic synthetic material. Its excellent biocompatibility makes it one of the most popular materials in medical applications [56]. However, similar to PEG, PVA does not support cell spreading and adhesion. Modification therefore is required before its use in tissue engineering [57]. PVA hydrogel is not biodegradable, but its acrylate modified form is degradable [58]. Other methods, like incorporation of biodegradable crosslinkers into PVA hydrogel, is also developed [59].

P(HEMA), the homo-polymer or co-polymer from 2-hydroxyethyl methacrylate (HEMA), has been used as implant materials for a long time [60]. The contact lens is of the classic chemical cross-linking hydrogels developed by Wichterle and Lim as a result of the copolymerization of hydroxyethyl methacrylate (HEMA) [61].

The p(HEMA) hydrogel is non-biodegradable, a property that restricts its application in tissue engineering, but is advantageous for its micro porous structure (a preferable morphology for tissue engineering scaffold) as a result of the polymerization-induced phase separation polymerization [62].

Poly(amido-amine)s (PAAs) are a class of polymers characterized by the presence of amido and tertiary amino groups regularly arranged along the macromolecular chain. The backbone of poly (amido-amine)s are usually synthesized by poly addition of primary monoamines, or bis(secondary amines), to bis-acrylamides [63].

PAAs can be easily modified during synthesis by the introduction of functional co monomers [64]. For example, a new PAAs was synthesized to contain the disulfide linkage in the main chain by means of stepwise polyaddition of 2-methylpiperazine to N,N-bis(acryloyl)cystamine (BACy1) or N,N-bis(acryloyl)-(L)-cystine (BACy2). This kind of functionalized PAAs exhibit their good biodegradability when they are degraded by reduction of the disulfide group under physiological conditions. Another example is the biomimetic poly(amidoamine) hydrogels for cell culture that evolve from the incorporation of 4-aminobutylguanidine (agmatine) moieties to create RGD-mimicking repeating units for promoting cell adhesion [64].

PAAs are synthetic polymers endowed with biologically interesting properties, such as being highly biocompatible, free from toxicity, and biodegradable. When it is positively charged, however, because of the positive charge of tertiary amine group and the absent of negative charge group, this kind of PAAs will show cytotoxicity, which is partially dependent on its positive charge arrangement and ions density [65,66,67]. Most PAAs for biomedical application have carboxyl groups along the main chain to decrease cytotoxicity and form amphoteric PAAs [68].

PEO-PPO-PEO is a triblock copolymer consisting of poly (ethylene oxide) (PEO) and poly (propylene oxide) (PPO). It attracts interests in biomedical applications as it forms a thermo-reversible gel that shows gelation at around the body temperature (37 °C) via PPO segments aggregation [69,70] Furthermore, the block composition (PEO/PPO ratio) and the molecular weight can be used to control the final properties of the products so as to meet the specific application needs [71]. However, it has the same issues in biodegradation [72] and cell attachment [73] when it is applied in tissue engineering. The solutions to these issues are similar to those related to PEG or PVA.

Combination of natural and synthetic hydrogels has been utilized to overcome shortcomings of pure natural and synthetic materials. Biomaterials combined in different ways have been developed to generate hydrogels with biological properties (e.g., hydrophilicity, cell-adhesiveness, degradability), biophysical properties (e.g., porosity, branched vasculature), and mechanical ones (e.g., stiffness, viscoelasticity) [5].

Modification of natural polymers for creating cross linkable functional groups is a basic step for hydrogel design and further application. Arylate is one of such functional groups for such modification. HA-acrylate can form a stable covalent crosslink hydrogel via a free-radical mechanism between carbon-carbon double bones (a covalent bond where two pairs of electrons are shared between the atoms rather than one pair) converting the double to a single bond and forming single bonds to join the other monomers. One of such efforts is to modify HA with methacrylic anhydride by adding ethyl eosin and triethanolamine as initiator, to result in the synthesis of a photocrosslinked polysaccharide hydrogel by irradiating using an argon ion laser [11].

Enzyme induced crosslink is also widely used to form hydrogels because of its low cytotoxicity. This crosslink system requires modification of natural polymer with an enzyme catalytic group. In one instance hydroxyphenylpropionic (HPA) acid was treated with NHS/EDC to link it to gelatin by reacting with tyrosine residues of the gelatin. Those conjugates can be catalyzed by hydrogen peroxide (H₂O₂) and horseradish peroxidase (HDP), and then crosslinked with each other to form gels because phenols can be crosslinked through either a more common C-C linkage between the ortho-carbons of the aromatic ring or a C-O linkage between the ortho-carbon and the phenolic oxygen. The stiffness of the hydrogels was readily tuned by varying the H₂O₂ or HDP concentration [48].

Other modification processes (e.g., disulfide linkage) can be affected by modifying a natural polymer with 3,3’-dithiobis(propionic hydrazide) (DTP) or cystine to form thiolated gelatin or HA macromer [20,74].

Cell anchorage is a strict requirement for the survival of most cell types, and it orchestrates critical roles in many cellular functions including migration, proliferation, differentiation, and apoptosis. Cell interaction with biomaterials is mediated through transmembrane receptors which recognize adhesion molecules at the materials surface [38].

PEG is a cyto-nonadhesive hydrogel and needs to be modified with cell adhesive peptides if it is to be used as a cell culture platform. Arg-Gly-Asp (RGD), a component of cell binding protein targeting αv-integrins, is widely used to promote cell adhesion [75,76]. RGD peptides was [49] incorporated into PEG diacrylate hydrogel through aminolysis of the N-hydroxy succinimide ester of acrylic acid by the a-amine terminus of the peptide sequence to produce an amide linkage between the peptide and the acrylic group.

PEG modified with polylysine [77] which serves as a functional cell attachment group for micro-vessel scaffolds is developed by activating PEG with N,N′-carbonyldiimidazole and then forming the gel by the reaction between free-amine groups in Poly-L-lysine hydrobromide [58].

Cross-links can be formed by means of chemical reaction initiated by heat, pressure, change in pH, or radiation [7]. As previously mentioned, the contact lens is a classic chemical cross-linking hydrogel developed by Wichterle and Lim, based on copolymerization of hydroxyethyl methacrylate (HEMA) with the crosslinker ethylene glycol dimethacrylate (EGDMA) [61]. This hydrogel is created by free radical chain polymerization which can be initiated by light, heat, or redox [78].

Reviewed here are the two major methods to initiate free radical chain polymerization for biomaterials: photo-initiated polymerization and radical-initiated polymerization for the preparation of hydrogels. The former is instanced by the modification of hyaluronic acid (HA) with methacrylate functional groups to result in the covalent cross-linking in the presence of a radical-initiated polymerization system. A concrete instance is the modification of HA with methacrylic anhydride to produce a synthetic, photocrosslinked hydrogel [11]. One more instance is the improvement of a polysaccharide hydrogel by using 2-Hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959) as the ultra violet light-sensitive photoinitiator to decrease cytotoxicity, and to encapsulate human embryonic stem cells and control cell proliferation and differentiation [16,36]. Irgacure 2959 was reported to have minimal toxicity towards six cell lines when it was used for cell encapsulation [79]. Upon absorption of UV light, Irgacure 2959, as a free-radical donor, is cleaved into two primary radicals which then react with the vinyl (C=C) groups of modified hyaluronic acid to initiate radical polymerization [80].

Radical-initiated polymerization can be achieved by thermal radical initiation reactions. One instance is the design of an injectable glucose sensitive microbead hydrogel by using a host macromer with bi-functional reactive group (vinyl groups), and adding sodium persulfate (SPS) as an initiator and N,N,N′,N′-tetramethylethylenediamine (TEMED) as a catalyst [81]. Researchers have now proved the feasibility of this initiation system for its negative cooperative effect to cytotoxicity in these injectable hydrogels [82]. They have also demonstrated the cytocompatiblity of the initiator concentration used in the APS/TEMED initiation system for MSCs in OPF hydrogels [83].

Diverse approaches have been utilized to produce in situ hydrogels with tunable properties using different triggers such as pH, temperature, light, and targeting biomolecules, as instanced by this facile approach to the creation of a “living” controlled in situ gelling system based on a thiol-disulfide exchange reaction by changes in the pH of the pre-gel solution [84]. Another example is the self-setting injectable hydrogels that can be triggered by a decreased pH. Silated macromolecules, such as silated hydroxyl propylmethylcellulosc have been used for such purposes [85]. Silated macromolecules can be self-crosslinked by silanol condensation to form hydrogels [86].

Physical hydrogels are not homogeneous, because clusters of molecular entanglements, or hydrophobically- or ionically-associated domains, can create inhomogeneities. Free chain ends or chain loops also represent transient network defects in physical gels [78].

Electrostatic between a polyelectrolyte and a multivalent ion of the opposite charge accounts for the formation of a physical hydrogel known as “ionotropic” hydrogel [78]. For example, alginates are naturally derived anionic polysaccharides and used as hydrogels by adding calcium ions as the opposite charge [87]. Divalent cations like Ca²⁺ cooperatively bind between the G blocks of adjacent alginate chains, creating ionic interchain bridges which cause gelling of aqueous alginate solutions [38].

With the development of nanotechnology, supramolecular self-assembly has attained keen interest in science [88]. Trials have been made adding multivalent ions of the opposite charge to the heated and cooled peptide amphiphile solution to cover and weaken the ionic force, thus breaking the balance between the hydrogen bond, ionic force or van der Waals forces, and to have monomers or macromers aggregated to form the gel [89]. This is instanced by the design of an injectable b-Hairpin peptide hydrogel, which can be triggered by shielding its ionic repulse force in a salt environment, and in which the peptides are capable of killing Methicillin-Resistant Staphylococcus aureus [90]. Besides the ionic trigger, enzymes also can serve as triggers to achieve biocatalytic self-assembly and gelation. A number of related precursors based on aromatic peptide amphiphiles were synthesized to self-assemble supramolecular hydrogel initiated by subtilisin, a hydrolytic enzyme from Bacillus licheniformis, which hydrolyses the methyl ester to form a peptide derivative that will cause a decrease in hydrophilic force. Recent research has shown that the structures can potentially be controlled by appropriate assembling conditions [91].

The polymer volume fraction in the swollen state being a measure of the amount of fluid imbibed and retained by the hydrogel [7], the swelling property to some degree characterizes water thus held in hydrogels, as indicated in a 2010 study. A hydrogel was cast, soaked in ethanol for 48 hours, and then allowed to air dry. When properly dried, the material was placed in a phosphate buffered saline (PBS) solution at 37 °C and allowed to swell for 16 hours. Readings were taken every 30 minutes for the first four hours, then every hour for the next four hours, then every two hours for the remaining eight hours. Gel fraction of each hydrogel sample was measured by determining the weight of the gel before and after soaking in the ethanol. After the initial 16 hours of swelling the samples were monitored for degradation. Weights were taken every 24 hours for the next 30 days. A fresh amount of PBS, previously equilibrated at 37 °C, was employed every 24 hours during the time of measurement to replace the used ones [93].

The mass swelling percentage of the hydrogel was calculated from the following relation [94]: %S = [(m_t − m₀)/m₀]× 100 (1) where m₀ is the mass of the dry gel and m_t mass of swollen gel at time t.

The swelling rate constant of the hydrogel was calculated from the following relation: Percentage mass swelling = k_st^0.5 (2) where k_s is the swelling rate.

The initial swelling data were fitted to the exponential heuristic equation. F = M_t/M_∝ = ktⁿ (3) where F is the fractional uptake; M_t/M_∝, when M_t is the amount of diffusant sorbed at time t, M_∝ is the maximum amount absorbed; k is a constant incorporating characteristics of macromolecular network system and the penetrant; n is the diffusional exponent, which is indicative of the transport mechanism. The equation is valid for the first 60% of the normalized solvent uptake. For Fickian kinetics in which the rate of penetrate diffusion is rate limiting, n = 0.5, whereas values of n between 0.5 and 1 indicate the contribution of non-Fickian processes such as polymer relaxation. Diffusion coefficients are important penetration parameters of some chemical species to polymeric systems. Using n and k, the diffusion coefficient (D) of the solvent in the matrix can be calculated using the following equations: K = 4[D/πr²]ⁿ(4) 4Dⁿ = k(πr²)ⁿ (5) Dⁿ = (k/4)( πr²) (6) where D is the diffusion coefficient and r the radius of gel disc.

The hydrogel is of a cross-linked network. A scanning electron microscope (SEM) is utilized by some researcher for the observation of gel morphology [36,62,77,95,96,97]. The sample for SEM should, first, be frozen quickly to best keep the original morphology upon the samples having reached their maximum swelling ratio in dd-water at room temperature after 24 hours and then freeze dried in a Freeze Drier until all water is sublimed. The freeze-dried hydrogel specimens are then cut and fixed on aluminum stubs and then coated with gold for interior morphology observation with a scanning electron microscope [36].

Because dehydration for SEM to some degree leads to artifact in the highly water-saturated gels, their morphology can be better viewed by cryosectioning the gels followed by staining with fluorescein isothiocyanate (FITC), and visulated under fluorescence or confocol microscopy [77].

SEM can also show cell morphology attached to the hydrogel when it is applied to tissue engineering. The method of treating samples can be quite different because cell morphology is subject to changes when going through the freeze drying process. Accordingly, in some studies hydrogels seeded with cells were washed with PBS, fixed with 4% glutaraldehyde or 10% buffered formalin, and dehydrated by the use of graded ethanol followed by the addition of hexamethyldisilizane. Cellular constructs were sputter coated with gold and observed under the SEM [77,98,99].

Backscattered electron microscopy (BSE-SEM) is commonly used for characterizing repaired bone sections by showing the mineralized tissues at the implant-bone interface [100,101]. Furthermore, BSE-SEM provides the mineralized composition and density of bone regenerated from nonautogenous grafts in bone tissue engineering [102,103].

The key structural parameter determining the material modulus and diffusional characteristics is the network crosslinking density. Many means are available for bulk or local gel mechanical measurement, such as the Dynamic Mechanical Analyzer for bulk gel mechanical measurement (DMA), atomic Force Microscopy (AFM) or Tracer Particle Microrheology (TPM) for the measurement of local gel and cell mechanical changes [8].

Researchers have been able to determine the compressive modulus of the various swollen hydrogels on a mechanical tester using a parallel plate apparatus and loading of 10% of the initial thickness per minute (~200 mm/min). Samples for mechanical testing (n = 5 per composition) are cylindrical (~2 mm height, ~7 mm diameter) and are compressed until failure or until 60% of the initial thickness was reached. The modulus is determined as the slope of the stress versus strain curve at low strains(<20%) [21].

Mechanical testing can also be performed on a DMA Q800 Dynamic Mechanical Analyzer in a “controlled force” mode, whereby the swollen hydrogel samples in the shape of a circular disc are submerged in distilled water and mounted between the movable compression and fluid cup; a compression force from 0.01 to 0.05 or 0.30 N (depending on the gel strength) at a rate of 0.02 or 0.05 N/min was applied at room temperature; the compression elastic modulus (E) of the swollen hydrogel can be extracted by plotting the compressional force versus strain [104].

In a self-assembly physical hydrogel, AFM is used for to show the gel nanostructure formation and nucleation and early-stage structure growth [91,105].

When cells are seeded to gels in 2D, cytotoxicity is largely related to the chemical structure of the hydrogel; under three dimensional situations, that is, when the cells are encapsulated in the hydrogel, cell viability can also be affected by poor nutrient and gas exchange because of the inappropriate cross-linking density [21]. Assays generally used to measure hydrogel cytotoxicity and cell viability are described as follows.

MTT assay is a popular approach to evaluate cell viability on hydrogels [20,24,36]. MTT (3-(4,5-Dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide, a yellow tetrazole) is reduced to purple formazan in living cells [106]. A solubilization solution (usually either dimethyl sulfoxide, an acidified ethanol solution, or a solution of the detergent sodium dodecyl sulfate in diluted hydrochloric acid) is added to dissolve the insoluble purple formazan product into a colored solution. The absorbance of this colored solution can be quantified by measuring at a certain wavelength (usually between 500 and 600 nm, an absorbance maximum at 490–500 nm in phosphate-buffered saline) by a spectrophotometer. The absorption maximum is dependent on the solvent employed [107].

Researchers have managed to assess the viability of photoencapsulated fibroblasts immediately after encapsulation and after 1 week of in vitro culture using a commercially available MTT viability assay (ATCC, 30–1010 K). For this assay, 100 µL of the provided MTT reagent (tetrazolium salt solution) is added directly to the wells containing the constructs (n = 3 per macromer solution) and placed in an incubator at 37 °C for 4 hours. The purple formazen produced by active mitochondria is solubilized by construct homogenization in 1 ml of the provided detergent solution and orbital shaking for 2 h. The absorbance of these solutions is then read at 570 nm (SpectraMax 384) (Molecular Devices, Sunnyvale, CA, USA) [21].

Live/Dead assays are cytotoxicity and viability assays too. Live/Dead fluorescent staining (Invitrogen) can be used to test osteoblasts for viability and proliferation at Day 1, 3 and 7. The live (green) and dead (red) cells are observed after 30 min incubation and sufficient PBS wash. For quantitative evaluation, WST-1 (Basel, Switzerland) colorimetric assay can be used. The incubation lasted for 2 h and the absorbance at 450 nm is measured with reference to 620 nm in a microplate reader [108].

Gels as tissue engineering scaffolds are mostly capable of degradation and ultimately absorbed in the body [9]. Gels can be degraded by certain chemicals or enzymes. For example, reducing agents such as DTT or GSH can degrade hydrogels containing disulfide linkages [20,23,109]. Some hydrogels made from nature polymer can be generally degraded by particular enzymes, such as hyaluronidases for hyaluronic acid [19] lysozyme for chitosan [33] and collagenase for collagen or gelatin [45]. Cells can also have effects on hydrogel degradation since cells can produce enzymes. For example, alginate hydrogel can be partially degraded by fibroblasts encapsulated in it and release calcium ions from its construct as a result [110]. A lot of methods have been developed to conduct degradation analysis. After degradation, the weight of hydrogel main body will be reduced, so the degradation can be directly characterized by the changes of hydrogel weight [93]. Since some degradation components will also be released from the gel main body into the testing solution, another qualitative approach to characterize the gel degradation is showing the release profile of the degradation component by collecting the degradation component and then calculating the cumulative amount of this component [111]. For detecting local degradation in gel systems, some researchers incorporate a FRET(fluorescence resonance energy transfer)-based linker within the gel-forming macromers [8]. Once the gel is treated by degradation solution, the FRET-based linker will generate strong fluorescence where the cells will subsequently extend the process and move within the degraded gel, leaving tracks of increased fluorescence. Upon local gel degradation by encapsulated fibroblasts, the increased fluorescence is observed, allowing visualization of the degradation tracks generated by the migrating cells degradable linker within the PEG-based gel [45].

Characterization of chemical properties includes clarification of macromer molecular weight that can be measured by Gel permeation chromatography (GPC), analysis of chemical structure by NMR [36], and so on. Among these, gelation efficiency is the important parameter that describes the capability of a hydrogel and can be expressed by the following equation: G_f= W_d/Wp*100% (7) where W_d is the weight of the dry hydrogel and Wp is the total feed weight of the two macromer precursors and the photo-initiator [104].