Photocrosslinkable Biomaterials for 3D Bioprinting: Mechanisms, Recent Advances, and Future Prospects

Collagen is a class of fibrous proteins in mammalian tissues with a specific alpha-triple-helix structure that provides mechanical support and enables the adhesion and migration of cells [154]. Depending on its structure and composition, collagen is divided into three types: type I, II, and III collagen [155]. Among them, type I collagen accounts for more than 90% of the total collagen mass in the body [155,156]. Based on their excellent bioactivity, collagen hydrogels have been widely used in various applications, like skin, cartilage, and blood vessel regeneration [157,158,159,160,161].

In the presence of residual tyrosine, the photocrosslinking of unmodified collagen can be achieved by photosensitizer-mediated redox crosslinking [162,163]. Although no additional chemical modification is required, the curing process involves a long UV irradiation time [162]. To address this, collagen was modified with extra photocrosslinkable groups, like methacrylate, norbornene, and thiols, to enhance both the curing rate and mechanical strength (Figure 5) [164,165].

Although typical methacrylate collagen (ColMA) bioinks have been successfully developed in 3D bioprinting, some limitations are still worth noting. First, since collagen tends to be physically crosslinked at neutral pH and room temperature, cells are frequently dispersed within an acidic ColMA solution before pH adjustment to prepare cell-laden bioinks, which is not favorable for 3D bioprinting. Low-temperature 3D bioprinting is a potential method. In the work by Yang et al., the authors prepared low-concentration ColMA scaffolds with good biocompatibility and bioactivity by using a low-temperature DLP 3D printing technique [166]. In addition to this, ColMA also faces oxygen inhibition, ROS accumulation, and other problems [167,168,169]. To address the drawbacks of traditional ColMA, Guo et al. prepared norbornene-modified collagen (NorCol) by conjugating carbic anhydride (CA) to a collagen backbone [170]. With the extra carboxyl groups, NorCol not only exhibited improved solubility under neutral conditions but also showed excellent miscibility with alginate and gelatin. This strategy enables hybrid bioink to respond to multiple stimuli, resulting in continuous crosslinked NorCol networks in hybrid hydrogels (Figure 5) [170].

When collagen is heated to high temperatures, its triple-helix structure will irreversibly disintegrate, resulting in the formation of gelatin (Figure 6). Due to the disintegration of the supramolecular structure, gelatin has a lower molecular weight and better solubility than collagen. As the denatured product of collagen, gelatin preserves similar biological features to collagen, such as abundant RGD sequences for cell adhesion and growth [175], and retains matrix metalloproteinase (MMP)-cleavable sequences that facilitate cell migration, growth, and extracellular matrix remodeling [176,177,178]. Unlike collagen, neutral water-soluble gelatin requires no additional acid and pH adjustment operations during bioink preparation. The gelatin solution also has thermal-responsive properties but is temporarily crosslinked at low temperatures [179].

In 3D bioprinting, gelatin can serve as the sacrificial material to build constructs with vascular networks or porous structures. For permanent 3D scaffold fabrication, additional chemical modification is frequently required. Among photocrosslinkable gelatin derivatives, GelMA is the most commonly used for 3D bioprinting due to its excellent biocompatibility, good processability, and stable physical properties. In general, GelMA is synthesized through the reaction between methacrylic anhydride and the amino or hydroxyl groups of gelatin [180,181]. In the presence of photoinitiators and light irradiation, photo-reactive methacrylate groups can be activated by free radicals to form permanent covalent bonds between gelatin backbones. Due to their good bioactivity and stability, various types of cells (stem cells, progenitor cells, cancer cells, primary cells, etc.) were 3D-printed by GelMA hydrogels for tissue engineering and regenerative medicine. Despite the versatility of GelMA in cell culture and 3D bioprinting, ROS accumulation during gelation still affects cell viability [112]. In addition, cell viability is also related to the substitution degree (SD) of methacrylate groups of GelMA. He et al. prepared a series of GelMA bioinks with different methacrylate degrees and evaluated the viability of cells within GelMA scaffolds [52]. Notably, the cells embedded in GelMA scaffolds with a lower SD showed higher viability. The use of a GelMA hydrogel with low strength facilitated cell spreading in the model (Figure 7a).

To realize fast and oxygen-tolerating crosslinking, gelatin-based thiol–ene bioinks have been introduced in 3D bioprinting, such as norbornene-modified PEG (PEGNB)/thiolate gelatin (GelSH), thiolate PEG (PEGSH)/norbornene-modified gelatin (GelNB), norbornene-modified hyaluronic acid (NorHA)/GelSH, GelNB/DL-Dithiothreitol (DTT), etc. Since pure PEG-based thiol-ene bioinks lack sufficient viscosity and bioactivity, GelNB or GelSH frequently plays multiple roles in extrusion-based bioink systems: (i) a thermal-responsive ingredient for temporary crosslinking, thus ensuring scaffold integrity; (ii) a curable agent for permanent photocrosslinking; (iii) a bioactive component to facilitate cell adhesion, migration, and extracellular matrix remodeling. The synthesis of GelSH is commonly performed through the ring opening of citiolone or γ-thiobutyrolactone or the conjugation of 3,3′-dithiobis(propionohydrazide) under an inert atmosphere to avoid thiol oxidation (Figure 6).

Currently, the synthesis routes for GelNB mainly include 5-norbornene-2-carboxylic acid and CA-based modification (Figure 7f). The reaction between 5-norbornene-2-carboxylic acid and gelatin is cumbersome and requires multiple steps to synthesize GelNB [180,182]. In contrast, the direct reaction between CA and gelatin was able to achieve GelNB with a substitution degree of about 44% [183]. However, excess water-insoluble CA was used during the reaction to ensure effective grafting. This may affect the purity of the final product. Alternatively, a dual-solvent system composed of DMF/water can achieve the synthesis of GelNB with a degree of substitution of about 90.1% [184].

Although GelNB/DTT is a versatile and powerful material for 3D bioprinting, bioinks based on GelNB and thiolate macromolecular crosslinkers have been developed to improve the curing speed and avoid the potential toxicity of DTT. Thiolate macromolecular crosslinkers also protect the cell from excess ROS damage. In the work by Zhao et al., who introduced thiolate heparin (HepSH) as a macromolecular crosslinker, lower ROS levels in the GelNB/HepSH hydrogel-loaded human umbilical vein endothelial cells (HUVECs) were found [112]. Compared with GelMA, HUVECs within a 3D-printed GelNB/HepSH hydrogel exhibit better viability and spreading ability (Figure 7b,c) [112]. As a cytotoxic molecule, the uncrosslinked traditional bifunctional thiol crosslinker DTT can lead to undesired cell damage during long-term storage or printing operations. This toxicity was greatly avoided by using a macromolecular crosslinker. In another work by Göckler et al., this cytotoxicity was greatly avoided by introducing GelSH as a macromolecular crosslinker [180]. In addition, GelNB/GelSH bioinks were able to undergo superfast gelation (1–2 s), even though the photoinitiator concentration was reduced dramatically (0.03%), and promoted continuous cell growth (Figure 7d,e) [180]. Hence, the thiol-ene-based gelatin system is a promising candidate to replace GelMA.

Hyaluronic acid (HA) is an extracellular matrix-derived glycosaminoglycan with repeat units consisting of D-glucuronic acid and N-acetylglucosamine [185,186,187]. The abundant carboxyl and hydroxyl groups endow HA with superior hydrophilicity and provide pendent sites for chemical modification. Due to its excellent biocompatibility and biodegradability, HA has been widely used for applications such as drug delivery, cell encapsulation, and wound dressings [105,188,189,190,191,192]. However, the pure HA solution is unable to form a stable hydrogel for cell encapsulation. In general, extra modification is required to endow HA with light-curing ability, such as alkenyl modification, thiolation, and tyrosine functionalization (Figure 8a).

Similar to gelatin, methacrylate HA (HAMA) is synthesized by grafting methacrylic anhydride onto the HA backbone via an esterification reaction [193,194,195]. To synthesize NorHA, HA is usually converted into intermediates like HA-TBA or HA-ADH, followed by amidation in the presence of a catalyst [196,197,198,199,200,201,202]. The synthetic route involves cumbersome and complex steps (Figure 8d). In addition, the NorHA hydrogel may become hydrophobic due to the consumption of hydrophilic carboxyl groups. More importantly, relevant studies have shown that the depletion of carboxyl groups also weakens the CD44 binding ability [203]. Alternatively, direct functionalization via esterification between CA and hydroxyl groups is a good choice to avoid the cumbersome synthesis step and the depletion of carboxyl groups (Figure 8a) [184,204]. In the work by Galarraga et al., the authors synthesized norbornene-modified HA (NorHA_CA) based on CA and found that the extra carboxylic group enables the accelerated degradation of thiol-ene hydrogels (Figure 8b,c) [204]. When combined with stable NorHA, the NorHA_CA/NorHA hydrogel exhibits a tunable degradation profile. NorHA_CA/NorHA showed great promise in fabricating hydrolytically degradable bioactive scaffolds through DLP.

For extrusion-based 3D bioprinting, which is the most common among 3D bioprinting methods, pure hyaluronic acid solutions cannot form self-supporting filaments [153]. Therefore, proper pre-solidification is frequently introduced for extrusion-based 3D bioprinting of HA. Meanwhile, combining HA with other natural or synthetic polymers, such as gelatin [181,189,195], collagen [205], methylcellulose [193,206], sodium alginate [207,208,209], polyethylene glycol [210,211], PF127 [212], or polycaprolactone [190,213], is another strategy to achieve scaffold building.

Alginate is a brown-algae-derived polysaccharide macromolecule consisting of β-D-mannuronic acid (M) and α-L-guluronic acid (G) repeats (Figure 9a). When multivalent cations Mⁿ⁺, such as Ca²⁺, Mg²⁺, and Ba²⁺, are inserted into the negatively charged polysaccharide chains of alginate, an egg structure is formed through ionic attractions, which results in a rapid sol–gel transition (Figure 9a) [214]. Since alginate can be cured mildly at normal temperature and pH, various cells, such as Schwann cells [47], BMSCs [215], fibroblasts [216], and HUVECs [217], have been 3D-printed with alginate bioinks.

Despite mammals lacking enzymes that can degrade sodium alginate, ionically crosslinked alginate scaffolds are still subject to degradation by ion exchange in vivo [220,221]. Apart from this, the impact of ionic crosslinkers on the viability of loaded cells is another question in alginate-based 3D bioprinting. As excess multivalent cations may cause undesired cell damage, the concentration of ionic crosslinkers should be carefully selected. Therefore, it is necessary to endow alginate with a light-curing ability to improve stability and biocompatibility. Photocurable AlgMA also can be prepared via methacrylic anhydride [214,222]. Compared to a purely ionically crosslinked alginate hydrogel, the structural stability of covalently crosslinked AlgMA scaffolds was significantly improved [214].

Generally, norbornene-modified alginate (AlgNB) can be synthesized through amidation between 5-norbornene-2-methyamine and carboxyl groups. Although the hydrophilicity of the thiol-ene hydrogel might be affected, this provides extra sites for bioactive modification. In particular, due to its non-fouling nature, the pure alginate hydrogel lacks a suitable environment to facilitate cell adhesion, migration, and growing [223]. Numerous studies have shown that the bioactivity of alginate can be improved by mixing it with hyaluronic acid [208], collagen [224], or gelatin [225] to form a composite bioink. In addition, the cell adhesion capacity of sodium alginate can be enhanced by RGD grafting of sodium alginate [218,226]. In a typical work, Ooi et al. created a thiol-ene crosslinking Alg-norb hydrogel that exhibited tunable mechanical and swelling properties, rapid gelation, and excellent bioactivity for extrusion bioprinting (Figure 9b) [212]. The thiol RGD sequence (CGGGRGDS) was grafted on Alg-norb through a photoinitiated thiol-ene reaction. Similarly, in another work by Liu et al., the RGD sequence (CGDS) was conjugated to maleimide-modified alginate (Alg-Mal) through a click reaction (Figure 9c) [219]. Although this requires a longer reaction time, these printed scaffolds exhibit improved biological performance.

Silk fibroin (SF) is a natural material derived from the Bombyx mori silkworm and is composed of an H-chain (350 kDa), L-chain (26 kDa), and amorphous glycoprotein P25 (30 kDa) with a ratio of 6:6:1 (Figure 10a) [227,228]. Considering its excellent biocompatibility, stable mechanical properties, biodegradability, and low cytotoxicity, it is no surprise that the FDA approved SF for biomedical applications. To date, SF has been selected for various aspects of regenerative medicine, including bone regeneration [229], cartilage repair [230], wound healing [231], and angiogenesis [231,232]. For example, it has been demonstrated that silk fibroin can promote the differentiation of mesenchymal stem cells toward an osteogenic phenotype, which makes SF an ideal choice for bone tissue engineering [233,234].

Depending on its great bioactivity, SF is also a promising candidate to serve as a scaffold backbone for 3D bioprinting. Due to the existence of strong hydrophobic interactions and H-bonding, the SF hydrogel is more stable and stronger than other natural biomaterials [235]. However, the printability of pure SF bioinks is still limited by low viscosity. Hence, SF is frequently blended with other biomaterials to reach proper viscosity, thus improving the structural integrity and stability of the printed scaffolds [236]. For example, in the work by Moon and coworkers, SF was blended with iota-carrageenan (CG). The prepared SF/CG ink exhibited suitable viscosity and shear-thinning properties, which ensures high shape fidelity (Figure 10b) [237]. However, to build stable constructs, extra light irradiation is required to initiate redox-mediated dityrosine crosslinking (Figure 10b). Thanks to its natural tyrosine residues, which endow SF with an inherent redox-based photocrosslinking ability, no additional chemical conjugation of photo-reactive moieties is required [238].

Although it has been reported that SF bioinks can complete gelation within seconds under visible light irradiation, the influence of oxidated photoinitiators on the survival of embedded cells remains challenging [239,240]. To this end, photocrosslinkable SF derivatives such as norbornene-modified SF (SF-NB) [241] and methacrylate SF (SFMA) [242] were synthesized. Through photocrosslinking modification, it is possible to realize cell-friendly crosslinking and enhance physical performance. Bhar et al. prepared a composite bioink containing SFMA, GelMA, and photoactivated human platelet releasate (PPR) to build an immunocompetent human skin model, where SFMA was employed to maintain the strength, stability, and elasticity of the printed constructs (Figure 10c) [243]. With this design, the printed artificial model showed tunable physical properties to support the biomimetic skin epidermal and dermal structures.

A decellularized extracellular matrix (dECM) is a natural tissue- or organ-derived material obtained by chemical or physical decellularization techniques (Figure 11a). Apart from its great biocompatibility and biodegradability, dECM is superior in preserving the in vivo physical, mechanical, and biological microenvironment, thus facilitating tissue-specific cell proliferation and differentiation. Currently, dECM bioinks are extensively developed for cartilage [244,245,246], heart [247], liver [248,249], skin [250], and tendon regeneration [251] for their ability to offer tissue-specific cues to guide cell growth.

The viscous dECM solution is thermally responsive and experiences invertible sol–gel transition at body temperature, which makes dECM suitable for extrusion-based 3D bioprinting. However, due to unstable physical gelation, as-printed dECM scaffolds are often mechanically weak with unsatisfactory resolutions. Therefore, mechanical reinforcement strategies, such as incorporating support materials and extra crosslinking, are frequently introduced to improve printability and integrity. In addition to co-printing with synthetic or natural supporting materials (e.g., PCL, gelatin, alginate) and adding crosslinkers (e.g., genipin), photocrosslinking modification is another common approach to stabilizing dECM 3D scaffolds.

Moreover, a suitable mechanical clue is critical for cells to differentiate and function in an organ-specific way. Through proper photocrosslinking, it is possible to build a dECM hydrogel with enhanced and tunable mechanical strength similar to the original tissue or organ. Like silk fibroin, many dECM materials are readily available for redox photocrosslinking due to their abundant native tyrosine residues. One advantage of this approach is the effective protection of pristine bioactivity. In a typical work, Jang et al. realized the photocrosslinking of dECM when vitamin B2 was used as a UV-light-sensitive photoinitiator [256]. Although vitamin B2 is approved by the FDA for the treatment of corneal disorders, the use of UV light can lead to potential DNA damage. For the application of such dECM bioinks, visible light is certainly a more attractive option to initiate crosslinking. To address these issues, Ru/SPS was developed to realize safe and efficient visible-light-mediated crosslinking of dECM (Figure 11b) [253].

For 3D bioprinting, Kim and coworkers reported porcine heart (hdECM) and bovine eyeball (Co-dECM) dECM-derived bioinks, which can be crosslinked rapidly within 5 s through Ru/SPS-mediated visible light crosslinking. The printed scaffolds exhibit improved printability, physical performance, and excellent bioactivity at the centimeter scale (Figure 11c) [124]. However, the potential damage induced by excess oxidated photoinitiator is still the main obstacle to the further application of dECM bioinks.

Chemically modified photo-reactive dECM (e.g., dECM-MA, dECM-NB) allows controllable rapid gelation and enhances mechanical stability while avoiding the use of oxidative photoinitiators [257]. Similarly, methacrylation is one of the main methods to modify dECM with photo-reactive ability, which mainly includes glycidyl methacrylate (GMA) or methacrylic anhydride (MA) functionalization. Although the structural integrity of dECM is not strongly affected by GMA or MA modification, more methacrylate groups can be grafted on dECM through a ring-opening reaction between GMA and amine or hydroxyl groups (Figure 11d). In addition, with the introduction of additional hydrophilic hydroxyl moieties, the GMA dECM hydrogel exhibited better bioactivity [254]. Similarly, in another work by Duong et al., CA was used to create norbornene-functionalized small intestine submucosa (dSIS-NB), which enabled the introduction of extra carboxyl groups (Figure 11e) [255]. The dSIS-NB bioinks showed versatility in building biomimetic scaffolds to support fast cancer cell spreading and angiogenesis through both DLP and extrusion bioprinting [255].

In contrast to natural biomaterials, synthetic polymers can be prepared in large quantities with controllable molecular weights and functional groups. Polyethylene glycol (PEG) is a linear polymer synthesized from ethylene glycol via polymerization. PEG has been approved by the FDA for pharmaceutical and medical applications due to its excellent biocompatibility and low immunogenicity [258]. It is no surprise that PEG and its derivatives have been widely explored in photocrosslinkable bioinks to serve as polymer backbones, crosslinkers, or sacrificial materials [259,260,261]. With terminal hydroxy groups, PEG can react with methacryloyl chloride or acryloyl chloride under basic conditions to synthesize PEGMA or PEGDA through esterification. The PEGMA or PEGDA precursor is readily crosslinked into hydrogels through chain-growth polymerization [55,262]. The physical properties of the PEG hydrogel can be efficiently adjusted by varying the molecular weight and monomer concentration. For thiol-ene-based photocrosslinking, PEG is frequently thiolated or norbornene-functionalized into various derivatives such as PEG-4SH and PEG-4NB, thus realizing versatile control over the hydrogel properties (Figure 12a) [263,264,265].

Providing a proper microenvironment to promote cell growth is essential for bioinks. Despite its good biocompatibility and hydrophilicity, it is hard for PEG to support continuous cell growth due to the lack of suitable cell adhesion and enzymatically degradable sites. Therefore, a variety of natural biomaterials have been blended with PEG to improve its bioactivity, such as gelatin, collagen, and hyaluronic acid [211,269,270]. Meanwhile, the bioactivity of PEG-based bioinks can also be effectively improved by incorporating biological ingredients like RGD sequences or heparin-binding sites [271,272,273,274]. For instance, Schwegler et al. incorporated a chemically synthesized RGD sequence into PEGDA, thereby giving the printed material cell adhesion sites without changing its mechanical properties (Figure 12b) [266]. The most common strategy to address the lack of MMP-degradable sites is introducing MMP-responsive crosslinking agents [275,276]. For instance, Chen et al. introduced a tunable-degradability hydrogel, which was synthesized by integrating MMP-sensitive peptides with norbornene-modified eight-arm polyethylene glycol macromers [277]. To endow the PEG hydrogel with dynamic mechanical performance similar to matrix remodeling, Wiley et al. developed a dual-enzyme (thrombin and MMP)-degradable peptide linker. The obtained PEG hydrogel allows a dynamic modulation process similar to matrix remodeling (Figure 12c) [267].

In another work, Zengin et al. developed a composite bioink containing MMP-sensitive peptide-modified mesoporous silica nanoparticles (MSN-MMPs), which resulted in an MMP-9-biodegradable hydrogel (Figure 12d) [268]. Furthermore, additional cysteine-modified RGD peptide incorporation enhanced cell-matrix interactions and supported the viability and proliferation of MG63 cells within 3D-bioprinted scaffolds (Figure 12d) [268].

Pluronic is a class of copolymers consisting of hydrophilic polyethylene oxide (PEO) and hydrophobic polypropylene oxide (PPO). Pluronic F127 (PF127) is a triblock polyether copolymer of PEO-PPO-PEO (Figure 13a). When the temperature is raised to the lowest critical solution temperature (LCST), the hydrophobic chains of PF127 will aggregate. In contrast, the hydrophilic chains will stretch into the aqueous phase to form microspheres. Finally, a hydrogel is formed through the aggregation of microspheres (Figure 13a) [278]. Based on its innate temperature-sensitive property, PF127 is widely used in 3D bioprinting to act as a sacrificial ingredient, polymer backbone, or supporting material [279,280,281]. Zheng et al. reported an electrohydrodynamic (EHD) inkjet printing system composed of GelMA and PF127, where PF127 served as a sacrificial matrix to create biomimetic microvascular structures (Figure 13b) [282]. Human dermal fibroblasts (HDFs) and HUVECs were successfully co-cultured to form a structure with tissue-specific morphology and high spatial resolution (30 μm). In another work by Lewis et al., PF127 was co-printed with liver dECM to fabricate biomimetic geometry to guide the directional formation of biliary trees.

Although PF127 can form a hydrogel at body temperature, a pure PF127 hydrogel tends to disintegrate rapidly upon the penetration of ambient water, which is unacceptable for permanent 3D device construction. Therefore, the modification of PF127 is frequently performed, like esterification between acryloyl chloride and hydroxy groups. Similar to PEGDA, the stability and fidelity of 3D-printed PF127 diacrylate (PF127DA) scaffolds are greatly improved by covalent crosslinking [285,286,287]. Since PF127 lacks cell adhesion sites, it is often co-printed with natural polymeric materials to improve bioactivity [280]. In the work by Millik et al., they reported a coaxial-nozzle-mediated tubular coextrusion printing system, where the pure PF127 of the inner phase served as a sacrificial material to guide tube formation. At the same time, the F127–bisurethane methacrylate (F127-BUM) of the outer layer was photocrosslinked to create permanent constructs [288]. After functionalization with collagen I to promote cell adhesion, it is possible to build luminal scaffolds (∼150 μm) with monolayers of HUVECs.

Polyvinyl alcohol (PVA) is a nontoxic, biodegradable, and biocompatible hydrophilic linear polymer that has been approved by the FDA for biomedical applications. Similar to PF127, PVA has also served as a sacrificial material [289], supporting agent [290], and hydrogel backbone [291]. One outstanding feature of PVA is its abundant hydroxyl groups, which allow physical crosslinking through H-bonding. Although toxic additives are not required for physical crosslinking, the PVA hydrogel prepared through H-bonding requires long, repeated freeze–thaw cycles, which is impractical for cell encapsulation [292]. In contrast, PVA can be crosslinked through chemical crosslinkers (boric acid, aldehydes, or epichlorohydrin) under mild conditions, which makes it possible to realize stable cell encapsulation [293]. However, the toxicity of small chemical crosslinking molecules is still one of the main obstacles to further biological advances. Therefore, it is necessary to perform extra chemical modifications to realize safe cell encapsulation and 3D bioprinting of PVA. The abundant hydroxyl groups provide enough sites for chemical modification, such as amine [294], carboxyl [295], methacrylate [296], norbornene [294], and thiol [297] groups.

Generally, PVA is often modified by photo-reactive moieties like methacrylate or norbornene to create a photocrosslinked hydrogel with stable and strong networks. The improved gelation speed, enhanced stability, and preserved biocompatibility have made the photocrosslinked PVA hydrogel a promising candidate for tissue engineering and regenerative medicine. For example, Guo and coworkers developed a composite hydrogel composed of methacrylate-modified PVA (PVAMA), AlgMA, and GelMA to induce pancreatic differentiation of induced pluripotent cells (iPSCs) [298]. The incorporation of PVAMA and AlgMA provided a high-water-content environment to support cell migration [298].

Meanwhile, PVA-based photocrosslinkable bioink also plays an important role in 3D bioprinting. In a study by Lim and coworkers, the authors developed a PVAMA and GelMA composite bioink that enables high resolution for DLP-based bioprinting (25–50 μm) [299]. The printed constructs were able to support the growth and maintain the stemness of encapsulated stem cells [299]. Based on the merits of thiol-ene chemistry, norbornene-modified PVA (nPVA) was also synthesized through esterification between CA and hydroxyl groups. Zhu and coworkers prepared a PVA-NB hydrogel through thiol-ene photocrosslinking, which showed great potential to serve as a tissue scaffold for pelvic organ prolapse treatment [300]. In another work, Qiu and coworkers developed a dynamic bioink based on nPVA and gelatin, which allows the rapid fabrication of cell-laden constructs through tomographic VBP (Figure 13c) [284]. The fabricated constructs exhibit controlled physicochemical performance that facilitates continual cell growth, supports the osteogenic differentiation of stem cells, and realizes aligned multicellular aggregates [284]. Although pure photocrosslinkable PVA can be 3D-printed into scaffolds with desired mechanical properties and fidelities, PVA is frequently functionalized by RGD or blended with natural biomaterials to improve the bioactivity due to its antifouling nature and lack of cell adhesion sites [299,301]. For summary, the Representative biomaterials for light-based 3D bioprinting were presented at Table 2.