Ionic crosslinking involves the association of polymer chains by noncovalent interactions. A crosslinked hydrogel network is formed when molecules containing opposite charges are blended, e.g., polyelectrolyte solution and multivalent ions (Gulrez et al., 2011).The ions of opposite charges electrostatically attract each other giving rise to a crosslinked polymeric network. The network can also be disrupted by using specific chelators to remove the multivalent ions from the polymeric network to reverse the gelation process.

In some hydrogels, hydrophobic interactions or hydrogen bonding interactions play a significant role in crosslinking. These interactions often tend to be temperature dependent and alter the rheology of the hydrogel network with changing temperature. Some hydrogels exhibit a random coil conformation at high temperatures; upon lowering the temperature, the polymer chains adapt a more ordered conformation forming junction points and consequently aggregating as physical gels (Jeong et al., 2002). Others undergo the reverse process. In reverse gelation, polymeric strands mostly contain both hydrophobic and hydrophilic regions (Pradines et al., 2015). At low temperatures, water (solvent) molecules dissolve the hydrophilic parts of the molecule leading to complete dissolution of monomer chains. There is a strong binding affinity between water molecules and monomeric chains, leading to highly ordered water molecules along the monomeric chains. The forces of interaction here are primarily hydrogen bonding forces, where the binding is an enthalpy-driven process; however, when the temperature increases, entropy overcomes enthalpy as water molecules become randomly oriented and hydrophobic interactions between the monomeric chains become predominant. Thus the chains dehydrate as the water molecules break free. With the increase in temperature, the polymeric interactions become more dominant as compared to the hydrogen bonding forces.

The association of disordered molecules into an ordered state by virtue of its nature or by electrostatic or covalent interactions is referred to as “self-assembly,” which is a primary mechanism of gelation observed in amphiphilic peptide molecules (Hauser et al., 2015). Synthetic peptides with a cationic polar head group and anionic groups along its backbone undergo rapid self-assembly from a random coil configuration to form ordered α helices and β sheets. With increased concentration, β sheets undergo further self-assembly to form nanofibrous networks that bear close resemblance to the native ECM. Some of the most commonly employed physically crosslinked hydrogels in bioprinting are agarose, alginate, chitosan, collagen, gelatin, Matrigel™, and Pluronic^®, which are further discussed in Section 3.2.1.1.3 in detail.

Schiff base is a compound formed by the nucleophilic addition of an amine to a carbonyl functional group. Amine containing amino acids, natural polysaccharides, or other synthetic polymers can be easily crosslinked under mild reaction conditions in the presence of crosslinkers containing aldehyde functionalities such as glutaraldehyde or other polyaldehydes (Hennink and van Nostrum, 2012). Other complementary functional groups that can also react with aldehydes to form crosslinked hydrogels include alcohol and hydrazides (Hennink and van Nostrum, 2012). The degree of crosslinking depends on the concentration of the crosslinker used. A high degree of crosslinking results in a hydrogel with strong mechanical properties; however, this reduces the degradation time of the hydrogel. Release of drugs, growth factors, or other biologics immobilized in a bioprinted hydrogel matrix might take longer to diffuse into the surrounding tissue region due to stronger encapsulation between the chemically crosslinked polymer chains. Some frequently used crosslinking agents (i.e., glutaraldehyde) are used to induce chemical crosslinking reactions between complementary functional groups. While glutaraldehyde might induce cytotoxicity (Takigawa and Endo, 2006), genipin is a benign naturally derived crosslinker, which has been used to crosslink gelatin (Bigi et al., 2002), collagen (Yan et al., 2010), chitosan (Chen et al., 2004; Moura et al., 2011), and fibrin (Gamboa-Martinez et al., 2015). The amine groups of the amino acids present in fibrin undergo a nucleophilic, ring closure type of reaction to form crosslinked hydrogels with genipin (Gamboa-Martinez et al., 2015).

Photopolymerization is one of the widely used crosslinking mechanisms, where low molecular weight monomers or oligomers undergo a process called curing to form a crosslinked polymeric network, when exposed to radiation²⁹. The crosslinking process is initiated in the presence of a photoinitiator that forms excited molecular species on irradiation with light and onsets the polymerization process. Often pure hydrogels do not crosslink by themselves unless an external initiator is added or the molecule is chemically modified with functional groups capable of undergoing free radical polymerization. The physical properties of the hydrogel can be efficiently controlled by proper modulation of the rate and degree of crosslinking by photopolymerization. Some commonly used photoinitiators in bioprinting include Irgacure (2-hydroxy-1-[4-(2-hydroxyethoxy) phenyl]-2-methyl-1-propanone) (Aubin et al., 2010), (Wang et al., 2015b), VA-086 (2, 2′-azobis [2-methyl-N-(2-hydroxyethyl) propionamide]) (Billiet et al., 2014), and Biokey (lithium phenyl-2,4,6-trimethylbenzoylphosphinate) (Fairbanks et al., 2009). While some photoinitiators produce free radicals via unimolecular bond cleavage, others undergo bimolecular reactions by producing excited species which undergo collision with another initiator molecule to generate free radicals. Irgacure is the most widely used photoinitiator in the bioprinting community, but it is only nontoxic below a certain threshold quantity. Even at concentration of 0.5% (w/v), this initiator has proven to be extremely detrimental to cells unless the excess photoinitiator is leached out of fabricated constructs (Arcaute et al., 2006). A few important requirements for photoinitiators are their water solubility, type of radiation, and exposure time required to generate free radicals. Although Irgacure is water soluble, it requires the use of ultraviolet (UV) light which is detrimental to cells with prolonged exposure. Long-term exposure to UV can damage the DNA, can induce undesired crosslinking, or affect cell functionality. In contrast, the use of photoinitiators such as Eosin Y and Biokey is advantageous as crosslinking can occur in visible light within a short span of time (Fairbanks et al., 2009). Using a low concentration (<0.1%) photoinitiator, as recommended by some studies (Fedorovich et al., 2009), can minimize toxicity but diminishes mechanical properties. Low mechanical properties often give rise to increased swelling and enhanced depth of curing, which deteriorates the bioprinted construct quality. Methacrylated gelatin (GelMA) and PEG are two of commonly used polymers capable of undergoing photopolymerization.

Alginate is a popular hydrogel used in bioprinting processes due its biocompatibility, various choices of crosslinking and bioprinting methods, low price, and ease of use in the creation of 3D structures (Ozbolat and Hospodiuk, 2016). It is a polysaccharide made of alternating β-D-mannuronate (M) and its C-5 epimer α-L-guluronate (G) units (Jungst et al., 2015). The G subunits are responsible for forming the gel phase of alginate. As the number of G subunits increase, the degree of gelation increases. Alginate undergoes ionic crosslinking in calcium chloride (CaCl₂) or calcium sulfate (CaSO₄) solutions. The divalent calcium ions form a bridge, due to the attraction of negatively charged carboxylic acid groups between two neighboring alginate chains (Lai et al., 2016). Alginate has been widely used for cell encapsulation, including neural stem cells (Purcell et al., 2009), skeletal myoblasts (Hill et al., 2006), and articular chondrocytes (Alsberg et al., 2002; Ab-Rahim et al., 2013). Moreover, alginate provides favorable conditions for adipogenic differentiation and does not alter cell morphology (Galateanu et al., 2012). In addition to cell encapsulation, alginate has been successfully used as a 3D matrix for in vitro culture of organoids and embryos, i.e., pancreatic islets (Tomei et al., 2014). However, despite the intrinsic properties of alginate that make it a favorable material for tissue engineering applications, chemical modifications are often required to promote desirable cellular functions, provide a greater range of mechanical properties, and facilitate controlled release of encapsulated factors. It was shown that, due to the highly hydrophilic nature of alginate, proteins are minimally adsorbed thus hampering cell attachment (Rowley et al., 1999). This can be overcome by modifying the alginate surface with peptides, such as RGD, to provide molecule binding sites for cell adhesion. EBB of alginate has been one of the most popular techniques in the bioprinting community (Ahn et al., 2012; Ozbolat and Koc, 2010; Yu et al., 2013). Alginate has several advantages for use in EBB, including the ease of crosslinking and a wide range of concentrations with superior mechanical properties. Alginate can be extruded in one of two forms, precursor and pre–crosslinked, depending on the application. Alginate in 2–4% (w/v) is extrudable, structurally stable and biologically acceptable and, after contacting the crosslinker, solidifies rapidly and maintains its 3D shape (see Fig. 3.2B) (Zhang et al., 2013a,b; Zhang et al. 2015; Ozbolat et al., 2014; Dolati et al., 2014). The concentration of alginate determines the viscosity of the solution, porosity and crosslinking time. It is possible to form pre-crosslinked alginate by mixing with low concentrations of the crosslinker, which increases its printability (Cohen et al., 2010b). Depending on the purpose, the bioprinted constructs can be strengthen by further addition of the crosslinker. DBB of alginate is feasible as long as its concentration allows for droplet formation [in general <2% in inkjet bioprinting (Gudapati et al., 2016)]. Exploiting the crosslinking mechanism of alginate with CaCl₂, heterogeneous tissue constructs comprised of amniotic fluid–derived stem cells (AFSCs), detrusor smooth muscle cell, and biliary epithelial cells were fabricated (Xu et al., 2013). Mechanical properties of bioprinted constructs can also be improved by combining alginate with other functional materials. For example, bifurcated vascular tissue constructs were fabricated within fluorocarbon employing a blend of 3% (w/v) low melting agarose and 3% (w/v) low viscosity alginate (Blaeser et al., 2013). The buoyant forces of the fluorocarbon solution supported soft tissue constructs. Similarly, zigzag cellular constructs were fabricated by printing 3T3 fibroblast-loaded alginate solution into a CaCl₂ crosslinker pool, where the crosslinker provided buoyancy support (Xu et al., 2012). Using an opposite configuration, another work also demonstrated bioprinting of alginate; heartlike tissue constructs were fabricated by bioprinting the crosslinker (CaCl₂) solution into an alginate pool using a modified HP inkjet printer (Xu et al., 2009). LBB with alginate of different concentrations has been widely described using MAPLE-DW, where an alginate suspension containing cells was placed on the ribbon then pulsed with a laser (Guillotin et al., 2010; Yan et al., 2013; Gudapati et al., 2014). Successful bioprinting depended upon alginate concentration and viscosity, surface tension, the bioink layer thickness coated onthe intermediate layer, gelation times, wettability, and laser fluence (Gudapati et al., 2014).

Chitosan, a linear polysaccharide molecule obtained from deacetylation of chitin, has a wide range of applications in tissue engineering such as cartilage regeneration, devices for hemostatic and antibacterial activity, formation of sponge scaffolds, and fabrication of wound dressings (Croisier and Jérôme, 2013). The best concentration for optimal biological response is 1.5% (w/w) with a degree of acetylation between 30% and 40% (Montembault et al., 2006). Due to its unstable mechanical properties and limited bioprintability (see Fig. 3.2C), chitosan is an appropriate material for cell encapsulation, but not for the formation of large-scale scaffolds (Geng et al., 2005). Chitosan has been used in EBB for bioprinting of perfusable vessel-like microfluidic channels, where chitosan was crosslinked by an ionic crosslinker, sodium hydroxide (NaOH) (Zhang et al., 2013a,b). The polymer and the crosslinker solution were printed simultaneously using a coaxial nozzle unit with the crosslinker solutions in the inner core diffusing out toward the polymer resulting in hollow tubular constructs. Bioprinting of chitosan using DBB and EBB has not yet been attempted.

Collagen type I is a triple helical biocompatible protein obtained from natural sources which has been extensively used in bioprinting (Ferreira et al., 2012). It is one of the major components of connective tissues and occupies about 25% of the entire protein mass in most mammals. Collagen is a highly conserved protein cross-species causing minimal immunological reactions. Collagen matrix facilitates not only cell adhesion but also enhances cell attachment and growth due to abundant integrin-binding domains. Although collagen type I has been used in bioprinting, it has limitations as it remains in a liquid state at low temperatures and forms a fibrous structure with increased temperature or neutral pH. Complete gelation can take up to half an hour at 37°C. This slow gelation rate makes bioprinting of 3D constructs difficult, since deposited collagen remains liquid for more than 10 min. Also, cells deposited in collagen are not homogeneously distributed, as gravity pulls down the cells before gelation takes place. Low mechanical properties and instability along with the abovementioned issues necessitate the use of supportive hydrogels for collagen. EBB has utilized collagen alone as a bioink (Smith et al., 2004); however, mixing it with Pluronic^® has generated promising results (Homenick et al., 2011). The cell number increased significantly after 24 h postbioprinting demonstrating that collagen is suitable for cell growth. Fully crosslinked, bioprinted constructs can even be perfused (Lee et al., 2014b). DBB also takes advantage of collagen as a bioink material; however, collagen needs to be deposited before the onset of crosslinking (Deitch et al., 2008). In one study, Boland’s group used collagen as a bioink constituent to investigate cell adhesion and proliferation on collagen-coated cell repellant substrates (Roth et al., 2004). Additionally, the same group fabricated a bilayer skin graft that generated neoskins on mice that are near-identical to native skin with microvessels (Yanez et al., 2014). As collagen possesses a fibrous microarchitecture, its use in inkjet bioprinting is highly limited, therefore microvalve bioprinting has been preferred. For example, fibrin-collagen bioink incorporating one of two cell types, AFSCs or mesenchymal stem cells (MSCs), was bioprinted using a valve-based bioprinter into wound sites as a treatment for skin burns (Skardal et al., 2012). Similarly, LBB has been successfully used to coat thin layers of collagen on a laser-absorbing material; due to collagen’s sticky nature, it can be easily transferred using a laser source for fabrication of skin tissue constructs (see Fig. 3.2D) (Michael et al., 2013).

Fibrin is a hydrogel formed by the enzymatic reaction between thrombin and fibrinogen, the key proteins involved in blood clotting. It supports extensive cell growth and proliferation (Cui and Boland, 2009), plays a significant role in wound healing, and has been used in fabrication of skin grafts (Skardal et al., 2012; Yanez et al., 2014). Rheology of gelled fibrin has been extensively studied because of its nonlinear elasticity (Janmey et al., 2009). The fibrin network is comprised of filaments forming as soft complex that allows a high degree of deformation without breakage (Janmey et al., 2009). Human umbilical vein endothelial cells (HUVECs) exhibit angiogenic behavior when cocultured with fibroblasts in fibrin. This is an essential factor in the vascularization of large tissue constructs and provides an effective in vitro model for analysis of the fundamentals of the angiogenic process (Nakatsu et al., 2003). The disadvantage of using fibrin for in vivo applications is the possibility of a severe immune reaction or transmission of infectious diseases when heterologous proteins are used. To increase the efficacy and safety of fibrin hydrogels, bacteria and viruses must be inactivated or fibrinogen and thrombin produced as recombinant proteins by mammalian cell lines. In addition, the degradation of fibrin is rapid, which is not conductive to long-term culture. Due to the non–shear-thinning nature of fibrinogen and thrombin, fibrin is rarely extruded (see Fig. 3.2E). However, despite these drawbacks, a few studies have utilized fibrin (Gruene et al., 2011). Weak mechanical properties do not allow manipulation on fibrin after gelation rendering EBB of precrosslinked fibrin a challenge. However, bioprinting of the two components of fibrin is an ideal option for DBB. Bioprinting of thrombin is feasible using inkjet (piezo- or thermal inkjet) bioprinting, but fibrinogen leads to clogging issues with unstable droplet formation due to its fibrous nature. In general, fibrinogen at low concentrations (<2 mg/mL) is usable, but such a concentration range limits the mechanical properties of bioprinted constructs. Therefore, microvalve bioprinting is more convenient for DBB of fibrinogen (Gudapati et al., 2016). For example, thermal-inkjet bioprinting was employed for bioprinting human microvascular endothelial cells (HMVECs), but cell were loaded in culture media and precisely bioprinted on crosslinked fibrin (Cui and Boland, 2009). Cells aligned in fibrin and formed into an extensive capillary network after 21 days of culture. Fibrinogen and thrombin can be bioprinted from separate cartridges and combined on a platform to form a hydrogel; alternating layers of neural cells and fibrin gel were bioprinted generating viable neural constructs with a potential in neural engineering applications (Xu et al., 2006). Because of its prolonged crosslinking time and dependence on thrombin concentration, fibrin is generally difficult to print into a designed shape. Fibrin is not a suitable bioink material for LBB due to its delicate structure; however, it can be mixed with other LBB compatible hydrogels such as hyaluronic acid (HA) (Gruene et al., 2011).

Gelatin, a denatured form of collagen protein, is used in food and pharmaceutical industries (Liang et al., 2004; Gómez-Guillén et al., 2011). It is extracted from bones, skin, and connective tissues of animals (Gómez-Guillén et al., 2002). At low temperatures, gelatin strands self-associate to form helical structures leading to a gel-like form, which reverts back to a random coil conformation as temperature increases (Chiou et al., 2008). Gelatin retains the Arg–Gly–Asp (RGD) sequence from its precursor, is less immunogenic, and promotes cell adhesion, differentiation, migration, and proliferation (Sakai et al., 2009).Gelatin in gel form can be obtained by thermally induced crosslinking, and it can easily liquefy at 37°C. The weight of unmodified gelatin gel decreases by 50% after around 10 h of incubation, and the gel completely dissolves within 24 h (Sakai et al., 2009). To avoid dissolution, a variety of chemical crosslinking methods have been examined; however, such methods are problematic in in situ and in vivo applications as crosslinking agents are toxic to cell-laden constructs (Liang et al., 2004). Popular enzymatic crosslinkers, such as transglutaminase, have been successfully used (Jin and Dijkstra, 2010). Other crosslinkers include horseradish peroxidase (HRP) and hydrogen peroxide (H₂O₂) (Sakai et al., 2009). Gelatin can be successfully gelled in vivo and remained intact for 1 week without inducing necrosis in surrounding tissues (Sakai et al., 2009). Gelatin-encapsulated cells demonstrate long-term viability, but limited cell elongation (Benton et al., 2009; Nichol et al., 2010). Gelatin is rarely bioprinted in its native form due to its poor mechanical properties. To employ gelatin for bioprinting, it has been chemically crosslinked by the addition of agents such as glutaraldehyde (Hellio and Djabourov, 2006). The free amine groups along the gelatin polymer backbone undergo a nucleophilic addition reaction followed by dehydration with the two aldehyde groups of glutaraldehyde that form an imine functional group commonly referred to as a Schiff base (Hennink and van Nostrum, 2012). Often the degree of gelation depends on the pH of the medium. Molecules which are capable of accepting or losing protons with an increase or decrease in pH are sensitive to a change in the local pH environment. Using EBB, hepatocyte-laden gelatin was bioprinted into spatially defined 3D structures. Hepatocytes remained viable and conducted their biological functions for more than 2 months (Wang et al., 2006). Gelatin is, however, not a popular hydrogel for DBB. Some studies have been performed by blending fibrin with gelatin for liver tissue biofabrication (Xu et al., 2007). Boland et al. (2007) presented DBB of gelatin/alginate blend; however, CaCl₂ solution was selectively printed onto the blended solution for fabrication of cardiac constructs as can be seen in Fig. 3.2F. Gelatin has been successfully used in LBB due to its viscoelastic properties, stability, and ability to hold cells in precise positions without damaging cells (see Fig. 3.2G) (Raof et al., 2011). Additionally, its thermosensitive properties allow for cell transfer and ease of removal postprinting to reveal an application-specific growth surface.

Hyaluronic acid, a linear nonsulfated glycosaminoglycan, is ubiquitous in almost all connective tissues and a major ECM component of cartilage (Zhang et al., 2009). It behaves similar to collagen type I and has properties that can be beneficial in the treatment of osteoarthritis (Migliore and Granata, 2008). It is comprised of repeating disaccharide units of D-glucuronic acid and N-acetyl-D-glucosamine moieties linked by alternating β-1,4 and β-1,3 glycosidic linkages (Zhong et al., 1994). The three major functional groups, which govern the chemical activity of HA, are the glucuronic carboxylic acid group, the secondary hydroxyl group, and the N-acetyl group. Hyaluronic acid is widely used in tissue engineering due to its excellent biocompatibility and ability to form flexible hydrogels (Luo et al., 2000). Chemical modification of HA is an illustrative example of a bioink molecule polymerized by photo-crosslinking. To use HA as a bioink, chemical modifications of these groups are often carried out to enhance its rheological properties (Burdick and Prestwich, 2011); however, HA has slow gelation rate and poor mechanical properties. Hyaluronic acid is favored due to its role in early embryonic development, cell-friendly nature, and controllable mechanics, architecture, and degradation. Human embryonic stem cells encapsulated in HA, as opposed to 2D culture cells, maintained their phenotype, normal karyotype, and full differentiation capability (Gerecht et al., 2007). Additionally, the high molecular weight of HA adds an essential structural and organizational element into the ECM of native tissues (Kreger and Voytik-Harbin, 2009; Toole, 2004). It is instrumental in cell migration (Baier et al., 2007; Turley et al., 2002), nerve regeneration (Seidlits et al., 2010; Faroni et al., 2015; Ikeda et al., 2003), neuronal (Mészár et al., 2008; Banerjee and Toole, 1991; Margolis et al., 1975; Rauch et al., 2005; Baier et al., 2007) and glial development (Back et al., 2005; Liu et al., 2004), and wound healing (Aya and Stern, 2014). Hyaluronic acid has only minor cross-species variation and excellent biocompatibility (Dana et al., 2004; Leach et al., 2003). Pescosolido et al. (2011) developed a semiinterpenetrating network (semi-IPN) bioink by combining HA with a photopolymerizable dextran derivative (see Fig. 3.2H). In that study, a hydroxyethyl-methacrylate (HEMA) derivative of dextran was prepared by activating the HEMA group with N,N′-carbonyldiimidazole and coupled it to dextran, forming dex-HEMA. Hyaluronic acid/dextran-HEMA solutions were exposed to UV radiations in the presence of a photoinitiator inducing polymerization of the dextran chains. Hyaluronic acid moieties became entangled in the chemically crosslinked dextran network. Hyaluronic acid has been used in EBB but is generally blended with other hydrogels to enhance its bioprintability and solidification ability. For example, Skardal et al., 2010c used hyaluronic hydrogels crosslinked with tetrahedral PEG derivatives for bioprinting vessel-like constructs. To form extrudable ECM hydrogels, the symmetrical tetra-PEG molecules were acrylated to form tetraacrylates (TetraPAcs) then cocrosslinked with thiolated gelatin and HA derivatives. In another study conducted by the same group (Skardal et al., 2010a,b,c), methacrylated derivatives of gelatin and hyaluronic acid were synthesized separately and photo-crosslinked by a two-step process, before and after extrusion, to form more compact firm structures. Dynamic crosslinking of thiol-modified HA by gold nanoparticles was also studied to develop a hydrogel capable of reforming itself during and after bioprinting (Skardal et al., 2010a,b,c). Use of HA in DBB has not been demonstrated so far due to its viscous nature and slow gelation rate; however, it has been employed in LBB when combined with other hydrogels such as fibrin (Gruene et al., 2011) to facilitate faster crosslinking. The bioink consisting of human adipose-derived stem cells or endothelial colony-forming cells combined with hyaluronic acid and fibrinogen was coated on the donor slide and bioprinted using LIFT. The collector slide was blade-coated with HA-fibrinogen as well and subsequently crosslinked by thrombin. Alternating layers of HA-fibrinogen combined with both cell types were bioprinted to study the interaction between the two cell types.

Matrigel™ is a gelatinous ECM protein mixture secreted by Engelbreth-Holm-Swarm mouse sarcoma cells (Kleinman and Martin, 2005). It promotes cellular outgrowth from tissue fragments, vascularization, and differentiation of various cell types (Gaetani et al., 2012; Kleinman and Martin, 2005). One of the major advantages of culturing cells on Matrigel™ is that it promotes complex cellular behavior which is difficult to observe on other surfaces. For example, endothelial cells create capillary sprouts in Matrigel™ (Kleinman and Martin, 2005). Molecules promoting endothelial cell network formation (Paul et al., 2013) are potential candidates for use in tissue regeneration therapies while the molecules that inhibit such formation can be used for anticancer therapies. Matrigel™ has been vital in the development of tumor xenograft models in rodents for discovery of novel cancer treatments (Benton et al., 2011). Matrigel™ is an expensive material; however, it has the ability to create mechanically strong, 3D bioprinted constructs with higher cell survival rates than other popular hydrogels such as agarose and alginate (Melchels et al., 2012). Thermal gelation properties of Matrigel™ and collagen type I are similar; however, gelation of Matrigel™ is thermally reversible. Crosslinking occurs between 24 and 37°C, and gelation takes about half an hour, beginning at temperatures above 4°C. In EBB, Matrigel™ needs to be bioprinted before becoming fully crosslinked; thus, a cooling chamber is essential. To accelerate the crosslinking process after deposition and maintain the printed shape, a heated plate should be employed. In EBB, Matrigel™ has been used in a radioprotection study on human hepatic carcinoma and human mammary epithelial cells (see Fig. 3.2I) (Snyder et al., 2011), and osteo- and endothelial-progenitor cell printing (Fedorovich et al., 2011). For DBB, Matrigel™ has not been used as a bioink material but employed as a substrate to place and pattern cells (Horváth et al., 2015). Its thermal crosslinking property and optimal viscosity make it an attractive bioink for LBB. Matrigel™ has been used for developing cellular constructs with precisely placed cells using MAPLE-DW cell transfer. A study conducted by Schiele et al. (2009) involved encapsulation of many cell types, such as dermal fibroblasts, neural stem, myoblasts, breast cancer cells, and bovine pulmonary artery endothelial in Matrigel™. The donor slide was coated with Matrigel™ containing cells while the receiver was coated only with Matrigel™. In a similar study, 3D cellular constructs were fabricated using multiple layers of human osteosarcoma cells bioprinted onto Matrigel™ substrates, using LIFT technique (Barron et al., 2004).

Methacrylated gelatin, the denatured form of collagen consisting of methacrylate (MA) groups conjugated to its amine side groups, has been used for tissue engineering due to its favorable biological and adjustable mechanical characteristics (Nichol et al., 2010; Benton et al., 2009). Despite low cell proliferation rates, GelMA has relatively high mechanical strength and low swelling ratio and is amenable to blending with other hydrogels to increase cell survival. In 5% micropatterned GelMA, cells elongated, migrated, and formed interconnected networks with surrounding cells (Nichol et al., 2010). At higher concentrations, 10% or 15%, cells do not migrate; however, individual cells showed limited elongation and some multicellular networks formed. Methacrylated gelatin is frequently used in EBB studies such as bioprinting of chondrocytes (Schuurman et al., 2013), liver cells (Visser et al., 2013; Schuurman et al., 2013; Bertassoni et al., 2014; Billiet et al., 2014), and MSCs (Du et al., 2015). It has low viscosity at room temperature and is easy to extrude, and its crosslinking rate can be manipulated by the length of exposure to UV light. Methacrylated gelatin forms a biomimetic (Nichol et al., 2010) as well as an enzymatically degradable (Hutson et al., 2011) hydrogel that is mechanically strong when photo-crosslinked in the presence of a photoinitiator. For example, Demirci’s group bioprinted bone morphogenetic protein-2 (BMP-2) and transforming growth factor (TGF-β1) using DBB on a GelMA substrate to mimic a native fibrocartilage microenvironment. This substrate differentiated bioprinted human MSCs toward osteogenic and chondrogenic lineages in a spatial manner (Gutkan, 2014). Methacrylated gelatin has also been combined with other synthetic polymers such as polyethylene glycol diacrylate (PEGDA) and photo-crosslinked by Eosin Y-based photoinitiator (Wang et al., 2015b). Crosslinking ability with visible light, long-term biocompatibility, bioprintability, and enhanced rheology are some of the advantageous features of this bioink formulation. Scaffolds with a porous architecture based on GelMA have also been bioprinted using LBB, where HUVECs spread over scaffold within 4 days (Gauvin et al., 2012) (see Fig. 3.2J). Recently, Ovsianikov et al. (2014) demonstrated the first bioprinting of cells using 2PP. MG63 cells were encapsulated in GelMA hydrogel loaded with water-soluble photoinitiators. Ying-Yang structures were then bioprinted with solid and hollow parts, where cells exposed to the laser were damaged; however, they were able to proliferate and filled all the available space in 3 weeks of in vitro culture.

Pluronic^® F-127 is a trade name for synthetic polymer poloxamer-based polymeric compound. It exhibits a polymeric architecture consisting of two hydrophilic blocks between a hydrophobic block making it an effective surfactant (Mesa et al., 2005). There are 11 types of Pluronic^® polymers which differ by molar mass, percentage of composites, functionality, and temperature of crosslinking which can vary from 10 to 40°C (Wang et al., 2015a). Pluronic^® undergoes reverse gelation as it starts to crosslink with increasing temperature. Pluronic^® copolymer structures erode quickly and cannot hold structural integrity for longer than a few hours; however, Pluronic^®, in combination with other hydrogels such as PEG, is useful for drug delivery and controlled release applications (Gong et al., 2009). Increase of mechanical strength was shown also by addition of methacrylated hyaluronic acid (MeHA) and then bioprinted into a fluorescent tube (see Fig. 3.2K) (Müller et al., 2015). The study performed by Vashi et al. (2008) showed that bone marrow–derived MSCs were able to interact with each other in a 3D Pluronic^® environment and adipogenic induction started after 4 days of culture. Pluronic^®, without any additives, can maintain the viability of cells for up to 5 days with a dramatic decrease thereafter. However, when supplemented with 60 nM hydrocortisone, cell viability was preserved even in higher concentrations of Pluronic^® (Khattak et al., 2005). Pluronic^® can be cured by UV light for bioink applications; however, the radiation type and duration can impact cell viability, and the concentration of the photoinitiator can affect the metabolic activity of cells. Chemical crosslinking can make the gel more resistant to thermal degradation³³. Pluronic^® can be crosslinked enzymatically whereby two ends of PEO side chains can be formed into self-assembled micelles (Lee et al., 2011). Pluronic^® does not lose its thermally reversible properties, but its mechanical strength increases with enzymatic crosslinking. In EBB, acceptable bioprintability of Pluronic^® is obtained above 20°C as it becomes more viscous and exhibits shear-thinning behavior. Therefore, Pluronic^®, depending on its concentration, requires a heating system around the needle to transition from liquid to gel state; a heated plate to maintain the temperature of the construct after deposition should also be considered. Nonetheless, it is more advantageous to maintain Pluronic^® in a semiliquid state to preserve cell viability in homogeneous suspension. In the case of prolonged bioprinting times, a cooling system might be needed to maintain a lower temperature in the bioink reservoir. The reversible properties of Pluronic^® can be useful in fabrication of complex constructs. Pluronic^® in solid form (at room temperature or higher) can be surrounded by a second type of hydrogel and then placed at 4°C to liquefy the Pluronic^®. This procedure creates perfusable channels within bulky cell-laden constructs (Wu et al., 2011). The thermosensitive nature of Pluronic^® and its high viscosity is problematic for DBB and so has not been used. LBB has not been attempted as Pluronic^® is not viscoelastic, maintains a solid coating on the quartz support, and cannot transfer thermal energy to kinetic energy, which is essential for jet formation.

PEG has been widely used in medical and nonpharmaceutical products (Lin and Anseth, 2009; Pasut et al., 2008; Alexander et al., 2013; Wang et al., 2015a). It is a linear polyether hydrophilic compound which can be conjugated with proteins (Alconcel et al., 2011; Zhang et al., 1998), enzymes (Pasut et al., 2008), liposomes (Ishida et al., 2005), and other biomolecules. It has been employed as a drug delivery agent (Veronese and Pasut, 2005; Lin and Anseth, 2009; Alconcel et al., 2011) and used in biosensing or signaling (Sharma et al., 2004) and pharmaceutical applications (Alconcel et al., 2011). Because PEG is a hydrophilic polymer, it resists protein adsorption and cell adhesion. Hence, some cells, particularly chondrocytes and osteoblasts, encapsulated in PEG survive well even without addition of biological constituents (Benoit et al., 2006). Other cell types may require additional cell adhesion components, such as RGD peptides or fibronectin coating. Fibronectin precoating improves cell adhesion, spreading, and the organization of the actin cytoskeleton (Vladkova et al., 1999). PEGylation inhibits protein adsorption making PEG a very useful synthetic polymer for bioprinting of tissue constructs for regenerative medicine. PEG is water soluble, and its mechanical properties can be manipulated through variation of its chemistry (Gao et al., 2015). Altering the composition of PEG-based hydrogels in tandem with photo-crosslinkage in the presence of a photoinitiator allows alterations to the structural, functional, and mechanical properties of fabricated tissues. One of the major limitations of PEG is its poor mechanical strength; therefore, addition of diacrylate (DA) of MA is highly beneficial. However, additives such as DA and MA require photo-crosslinking by exposure to UV light for a specific length of time to obtain the desired mechanical properties. Excessive exposure to UV light can dramatically reduce cell viability. For example, rat aortic smooth muscle cells have been successfully photoencapsulated in PEG with 20 s of UV exposure (Mann et al., 2001). PEG was also used in creation of microgels that were assembled in seconds using acoustic waves (Xu et al., 2011). PEGDA and polyethylene glycol methacrylate (PEGMA) hydrogels are used in all types of bioprinting modalities: EBB (Hockaday et al., 2012; Wüst et al., 2014; Bertassoni et al., 2014; Skardal et al., 2010c), DBB (Cui et al., 2012a,b), and LBB (Hribar et al., 2014). Photopolymerization of PEG-based hydrogels with tunable mechanical properties has been widely demonstrated using EBB. Hockaday et al. (2012) performed 3D printing of anatomically correct aortic valve scaffolds using PEGDA hydrogels (see Fig. 3.2L). Scaffolds were then seeded with porcine aortic valve interstitial cells, where cells demonstrated adhesion with minimum spreading or proliferation in 21 days. Earlier work, however, demonstrated that the immobilization of cell adhesion sites and growth factors during EBB promoted cell spreading and proliferation (Zhang et al., 1998; Fedorovich et al., 2007). Due to greater mechanical stiffness compared to naturally derived polymers, such as alginate, fibrin, agarose, and collagen type I, PEG has been extensively used in DBB. Using thermal inkjet bioprinting, Cui et al. (2012a,b) bioprinted human articular chondrocytes in PEGDMA into osteochondral plugs in a layer-by-layer fashion with UV exposure following each layer, producing a 3D scaffold with uniform allocation of cells and neocartilage formation. In a similar study, acrylated-PEG was bioprinted with acrylated-RGD peptide, followed by photopolymerization of the bioprinted layer (Gao et al., 2015). Bone marrow–derived human MSCs were suspended in PEGDMA in conjunction with hydroxyapatite (HA) nanoparticles and bioactive glass and bioprinted using a thermal inkjet bioprinter (Gao et al., 2014), which conferred spatial control of the distribution of cells and bioactive ceramic materials in the fabricated bone tissue constructs. PEG-based hydrogels have been widely used in LBB, including SLA and DOPsL. Bioprinting using SLA was first demonstrated with a commercial SLA 3D printer (SLA-250, 3D Systems), where Chinese hamster ovary cells were encapsulated in PEGDMA with over 90% cell viability (Dhariwala et al., 2004). In a similar approach, the same 3D printer equipped with He–Cd laser was used for encapsulation of 3T3 fibroblasts in PEGDMA to fabricate well-defined cylindrical scaffolds with multiple channels (Arcaute et al., 2006). In addition to SLA applications, DOPsL enabled fabrication of vascularized tissue construct using PEGDA hydrogels to study the migration behavior of HeLa cells (Huang et al., 2014). It was demonstrated that HeLa cells migrated significantly when the vasculature diameter was decreased.

Table 3.1 provides sample bioprinting and bioink formulation parameters along with bioprinters used in EBB, DBB, and LBB of all the hydrogels presented in this chapter.