Light-activated decellularized extracellular matrix-based bioinks for enhanced mechanical integrity

Abstract

Decellularized extracellular matrix (dECM)-based bioinks have emerged as key materials in tissue engineering and 3D bioprinting technologies due to their ability to closely mimic the biochemical composition and structural organization of native extracellular matrices. These bioinks facilitate critical cellular behaviors, such as adhesion, proliferation, and lineage-specific differentiation, which makes them invaluable for constructing tissue analogs for applications in regenerative medicine, organ transplantation, and disease modeling. Despite their transformative promise, dECM bioinks face persistent challenges, including limited mechanical robustness, delayed gelation kinetics, and suboptimal printability, all of which constrain their translational utility. The advent of photocrosslinking technologies marks a paradigm shift, with light-activated functional groups such as methacrylates, thiol-enes, and phenols substantially improving the gelation efficiency, mechanical properties, and spatial fidelity of the printed constructs. The present review critically examines the state-of-the-art advancements in light-mediated dECM-based bioink crosslinking strategies, with a focus on innovations in bioink and photoinitiator design along with optimized crosslinking kinetics to address inherent limitations such as cytotoxicity and structural variability. Further, the review highlights the necessity of standardized dECM processing protocols and scalable biofabrication techniques to ensure reproducibility and clinical translation. By overcoming these challenges, dECM-based bioinks can enable the production of high-resolution, volumetric tissue constructs, thereby paving the way for transformative advances in regenerative medicine and translational biomedical applications.

Graphical abstract

1. Introduction

Tissue engineering and 3D bioprinting have revolutionized regenerative medicine by enabling the fabrication of tissue analogs that mimic the complexity of native tissues. The decellularized extracellular matrix (dECM), generated by removing cellular components from tissues, preserves a rich scaffold composed of structural proteins, growth factors, and bioactive molecules [1,2]. This distinct composition fosters a highly biocompatible milieu that facilitates cell adhesion, migration, and proliferation, which makes dECM an ideal candidate for bioink development [3,4]. Moreover, its capability to sustain the requisite microenvironment for tissue growth has driven its adoption in clinical applications, including tissue regeneration, organ transplantation, and disease modeling [5,6]. Despite these advantages, dECM-based bioinks suffer from inherent limitations, such as weak mechanical integrity and difficulty in maintaining structural stability after printing, largely because of their soft and fragile hydrogel nature [7,8]. These drawbacks compromise their ability to reproduce the structural intricacies and functional diversity of native tissues. In organs, spatial variations in stiffness and microarchitecture are critical for supporting specialized mechanical and biological functions [9,10]. Hence, the replication of these architectural and mechanical features is essential for functional fidelity; however, achieving such properties with dECM bioinks remains a formidable challenge.

To address these limitations, various crosslinking strategies—including thermal, chemical, and enzymatic methods—have been employed to reinforce dECM hydrogels [11]. Nevertheless, each approach presents inherent trade-offs between gelation time, mechanical reinforcement, and cytocompatibility. Photocrosslinking has recently emerged as a promising alternative, offering simultaneous improvements in mechanical strength, gelation kinetics, and structural stability, thereby facilitating the fabrication of geometrically and mechanically robust tissue constructs [[12], [13], [14]]. This technique employs photoreactive functional groups embedded within the bioink, which are activated by specific wavelengths of light to initiate polymerization and form robust covalent crosslinks [15]. This rapid crosslinking mechanism enables the creation of structurally complex constructs with enhanced shape fidelity and mechanical resilience, thereby facilitating the fabrication of high-resolution, centimeter-scale tissue constructs [16]. Photocrosslinking offers precise spatiotemporal control over bioink solidification, with key parameters such as light intensity, exposure duration, and photoinitiator concentration allowing fine-tuning of the crosslinking dynamics [14,[17], [18], [19]]. These adjustable parameters not only improve printability but also enable the fabrication of 3D constructs with spatially varied mechanical properties, thereby mimicking the heterogeneity of native tissues [20,21]. Photocrosslinking is effective for reinforcing the mechanical properties of dECM bioinks, surpassing those achieved by thermal crosslinking alone. It also enables rapid post-deposition stabilization, thereby improving the shape fidelity and structural integrity of printed constructs while preserving the native bioactivity of the extracellular matrix (ECM) (Fig. 1). However, despite these promising advancements, the development of light-activated dECM bioinks remains in its early stages, with current systems exhibiting limited versatility and functionality. Challenges include the functionalization of dECM with photoreactive groups without compromising its native bioactivity, as well as the selection of photoinitiators that strike a balance between efficient crosslinking and cellular compatibility. These challenges underscore the need for continued research and innovation to unlock the full potential of photocrosslinking in dECM-based bioinks.

Fig. 1.

Overview of dECM benefits, challenges, and the breakthrough of photocrosslinking.

This review provides a detailed exploration of light-activated photocrosslinking strategies in the development of dECM-based bioinks, with a focus on key mechanisms such as methacrylate-mediated crosslinking, thiol-ene chemistry, and phenol-mediated crosslinking. These strategies have emerged as essential advancements in addressing the critical limitations of dECM bioinks, including inadequate mechanical stability, slow gelation kinetics, and limited printability. By evaluating the unique advantages of these approaches, the review highlights their potential to enhance the gelation efficiency, mechanical robustness, and spatial fidelity, all of which are crucial factors in the fabrication of tissue constructs that closely replicate the structural and functional complexity of native tissues. Additionally, the review underscores the importance of optimizing key parameters such as photoinitiator selection, light intensity, exposure duration, and crosslinking density, which collectively determine the efficacy, reproducibility, and biocompatibility of the resulting constructs. While acknowledging these advancements, the review also addresses persistent challenges, including oxygen inhibition, cytotoxicity, and structural heterogeneity, which impede the scalability and clinical translation of dECM-based bioinks. By bridging the current limitations and future possibilities of dECM bioink, this review aims to advance the development of bioinks that are capable of reliably producing complex functional tissue constructs. Such progress has the potential to drive the scalable fabrication of clinically relevant engineered tissues, thereby opening up new frontiers in regenerative medicine and translational applications.

2. Key characteristics of decellularized extracellular matrix for use in tissue engineering applications

Biomaterials for use in tissue engineering and 3D bioprinting applications should meet specific requirements in order to replicate native tissues and restore their functionalities effectively. For instance, they should exhibit excellent biocompatibility to avoid any adverse biological responses or degradation into toxic byproducts [22]. In addition, the compositional characteristics, structural organization, and biophysical properties of the biomaterials play a critical role in influencing cell-ECM interactions, which govern key cellular behaviors such as adhesion and spreading [23]. However, many commonly-used biomaterials limit their ability to mimic the natural ECM complexity due to being composed of single components. To address these challenges, dECM has emerged as a promising biomaterial by offering an optimized microenvironment that closely resembles that of native ECM, supporting cellular activities and tissue development. In the remainder of this section, the fundamental characteristics of dECM hydrogels are presented, with a focus on their intrinsic biochemical cues that support the replication of tissue-specific microenvironments and their adaptable physical properties that contribute to structural stability. Additionally, the section explores diverse crosslinking strategies and their pivotal roles in improving the mechanical properties and functional performance of dECM-based hydrogels for tissue engineering applications.

2.1. Intrinsic biochemical cues via tissue-specific microenvironment recapitulation

Decellularization of native tissues can closely mimic the microenvironments of native tissues by removing any immunogenic cellular components while preserving the functional components, such as structural proteins, and growth factors [24]. These retained components play a fundamental role in modulating cellular behavior. Notably, dECM provides tissue-specific structural and biochemical cues that further promote lineage-specific differentiation [[24], [25], [26], [27]]. Han et al. demonstrated that mesenchymal stem cells exhibit distinct gene expression patterns when cultured in various dECM materials [28]. Their study highlighted how compositional variations in the dECM influence stem cell behavior, thereby directing differentiation in a tissue-specific manner. Consistently, dECM serves as a superior cell culture matrix that significantly enhances the expression of genes associated with cellular maturation and specialized functions compared to traditional biological scaffolds such as Matrigel and collagen [[29], [30], [31]]. Moreover, dECM is increasingly recognized for its clinical relevance in regenerative therapies, where it provides bioactive signals that enhance cell-cell and cell-ECM interactions in order to support tissue regeneration [32,33]. Various formulations of dECM-based medical devices have been under clinical trials and commercialization, including patch-type grafts, such as Alloderm and GraftJacket, for skin or rotator cuff repair, and injectable hydrogel forms, such as Ventrigel, for cardiac function restoration [[34], [35], [36]]. These applications demonstrate the growing clinical adaptability of dECM as a biomaterial for regenerative therapies, further expanding its potential for diverse tissue engineering applications.

2.2. Versatile physical properties and reinforcement strategies for structural stability

The initial forms of dECM were developed as solid scaffolds [24]. The ability to solubilize dECM into hydrogel form has significantly broadened its utility in diverse tissue engineering contexts, allowing it to be used as a cell culture substrate and an injectable material for defect repair [37,38]. Moreover, recent studies have further explored its potential, particularly in the context of biofabrication, including its applications in 3D bioprinting. Depending on the bioprinting modality, dECM-based bioinks must satisfy different rheological and crosslinking requirements to ensure optimal performance. In extrusion-based bioprinting, bioinks must exhibit shear-thinning behavior, rapid viscosity recovery, and sufficient yield stress to retain the filament shape during layer-by-layer deposition [39]. These properties are essential for maintaining high print fidelity and construct stability. Pati et al. demonstrated the use of various dECM-based bioinks to fabricate 3D-printed structures through extrusion-based printing techniques [23]. The dECM hydrogels exhibit suitable rheological properties such as shear thinning behavior, wherein the viscosity decreases with increasing shear rate. This property is particularly advantageous for extrusion-based printing, as it allows the dECM bioink to flow smoothly through the nozzle under applied shear forces [40,41]. Additionally, the complex modulus of the dECM hydrogel varies depending on the dECM type, and can be further tuned by simply adjusting the hydrogel concentration (Fig. 2A) [23,42]. Furthermore, the dECM exhibits sensitivity to shear stress and demonstrates shear recovery capabilities with rapid sol-gel transitions, making it suitable for use as a supportive bath in embedded bioprinting techniques [[43], [44], [45]]. When the applied shear stress surpasses the yield point, dECM undergoes a reversible transition from a solid-like to a flowable state, thereby generating a localized void that facilitates the precise deposition of extruded bioink filaments [45]. Utilizing dECM as a supportive bath facilitates the direct fabrication of tunable 3D conduits with structural stability, while simultaneously providing a tissue-specific microenvironment niche for cells (Fig. 2B) [29,[46], [47], [48], [49]]. In contrast, vat polymerization-based methods, such as digital light processing (DLP) and volumetric bioprinting, demand prepolymer solutions with tailored viscosity, high optical clarity, and controlled light scattering or absorption characteristics [50,51]. For these systems, the selection of photoinitiators, the tuning of light penetration depth, and the incorporation of photoabsorbers must be optimized to achieve precise spatial resolution and homogeneous crosslinking. In particular, modulation of light penetration and scattering is critical, as these parameters directly influence the resolution and fidelity of printed constructs [52]. To address this, various photoabsorbers have been employed, including tartrazine (λ_max ≈ 405 nm), curcumin (λ_max ≈ 425 nm), anthocyanin (λ_max ≈ 510 nm), and Ponceau 4R (λ_max ≈ 510 nm) [[53], [54], [55]]. These dyes compete with photoinitiators for photon absorption, thereby attenuating light transmission and enabling fine control over the curing depth and thickness of each printed layer [51]. Thus, the choice of printing strategy strongly influences not only the design of crosslinkable systems, but also the overall formulation of dECM bioinks, highlighting the importance of application-specific customization.

Fig. 2.

Conventional crosslinking strategies of decellularized extracellular matrix bioinks.

(A) Shear thinning behavior and a wide range of complex modulus of dECM hydrogel, supporting their suitability for extrusion-based bioprinting applications. Reproduced with permission from Ref. [23]. Copyright 2014 Springer Nature. (B) Shear recovery properties of dECM hydrogel, enabling their use as a supportive hydrogel bath for embedded bioprinting. Reproduced with permission from Ref. [47]. Copyright 2020 Wiley-VCH GmbH. (C) Thermal crosslinking of dECM hydrogels through self-assembly of collagen fibers (scale bar: 5 mm). Reproduced with permission from Ref. [23]. Copyright 2014 Springer Nature. (D) Silk-composite dECM bioink facilitating beta crystalline formation via self-assembly. Reproduced with permission from Ref. [73].Copyright 2024 Elsevier Ltd. (E) Chemical crosslinking of dECM bioink using tannic acid for enhanced mechanical stability. Reproduced with permission from Ref. [91]. Copyright 2020 Ivyspring International Publisher. (F) Ionic crosslinking of dECM bioink through the incorporation of alginate, improving structural fidelity. Reproduced with permission from Ref. [104]. Copyright 2018 WILEY‐VCH Verlag GmbH & Co. KGaA, Weinheim.

Nevertheless, the applications of dECM hydrogels are challenged by their intrinsic mechanical limitations, which are primarily due to the disruption of collagen fibers during the decellularization process [56]. This compromised structural integrity adversely affects the printability of the dECM hydrogel, thus resulting in constructs with low resolution and poor fidelity. To address these drawbacks, important factors include optimizing the gelation kinetics and enhancing the mechanical properties of the dECM hydrogel through rapid crosslinking during and immediately after printing [57,58]. These improvements led to enhanced printability and structural fidelity, making it possible to fabricate robust 3D-printed tissue constructs with better shape retention and resolution [59].

To assess these enhancements objectively, both printability and structural fidelity must be evaluated using standardized quantitative and qualitative methodologies. Printability is commonly characterized through filament fusion tests, collapse tests, and layer stacking fidelity measurements, which indicate the capacity of the bioink to retain its intended geometry during extrusion and deposition [60]. Moreover, rheological properties and crosslinking kinetics serve as critical indirect indicators of printability performance [61]. Structural fidelity is generally determined by comparing the printed construct with the original CAD model using such metrics as shape deviation, pore size consistency, and aspect ratio preservation [59,62]. Incorporating these evaluation techniques into bioink development provides a more robust framework for assessing and comparing the performances of dECM bioinks in 3D bioprinting applications.

2.2.1. Self-assembled crosslinking

The dECM hydrogel can undergo thermal crosslinking being a conventional approach when crosslinkers are not utilized. The viscous pre-gel dECM thermally crosslinks at physiological temperatures, where collagen within the dECM self-assembles into cross-striated fibrils that subsequently aggregate into elongated fibres. These fibers establish weak noncovalent interactions among collagen chains, including van der Waals forces, hydrogen bonding, and hydrophobic interactions [23,[63], [64], [65]]. This process improves the mechanical strength and thermal stability of the dECM hydrogel [66]. However, thermal crosslinking is hindered by a slow reaction rate due to restricted heat transfer within the hydrogel matrix, which reduces the structural stability, particularly for the fabrication of complex and volumetric tissue constructs [64]. Furthermore, as a physical crosslinking method, thermal crosslinking provides limited control over crosslinking density, unlike chemical approaches that allow precise modulation via stoichiometric control [65,67,68]. Recently, pre-fabricated polymer scaffolds with the desired geometry have been increasingly used to retain the construct shape while the dECM hydrogel is not yet fully thermally crosslinked [23,69]. Pati et al. infused polycaprolactone scaffolds with a dECM hydrogel, which was subsequently thermally crosslinked to conform to and preserve the intended shape, thereby facilitating the fabrication of implantable tissue constructs (Fig. 2C) [23].

Recently, other biomaterials with self-assembly properties have been explored to enhance the mechanical robustness of dECM hydrogels without crosslinkers. One representative strategy involves the incorporation of silk-based biomaterials. Silk fibroin is a natural biopolymer that is characterized by a combination of β-sheet and α-helix structures interspersed with random coil regions, which contribute to its crystalline nature [70]. In an aqueous solution, silk exhibits self-assembly properties, forming stable crystalline β-sheets through hydrogen bonding [70]. This nanostructured molecular organization provides silk with exceptional tensile strength and elasticity [71]. The integration of silk into the dECM hydrogel significantly enhances its mechanical robustness. Zhang et al. demonstrated that cartilage-based dECM hydrogels, when composited with silk, exhibit comparable mechanical properties to those of native tissues [72]. Furthermore, Major et al. demonstrated the spontaneous protein structural transformation of silk to achieve dynamic stiffening and increased crystallinity by integrating adipose-derived dECM with silk fibroin, thereby mimicking the ECM stiffening observed during breast cancer progression (Fig. 2D) [73].

2.2.2. Chemical crosslinking

In chemical crosslinking, covalent bonds are formed between polymer chains within the dECM by using either a synthetic or natural crosslinking agent. Among the various synthetic crosslinkers, glutaraldehyde (GA) and 1-ethyl-3-(3-dimethyl aminopropyl)-carbodiimide (EDC) have been widely incorporated into dECMs to enhance their mechanical strength. In particular, the GA molecule has an aldehyde group at each end of its chain, and these groups interact with ε-amine groups to form imine bonds within dECM [74]. Consequently, GA exhibits a high crosslinking efficiency, which significantly improves the mechanical characteristics and resistance to degradation of the crosslinked material. Conventionally, GA has been used for the fixation of clinical xenograft prostheses that require post-decellularization crosslinking [[74], [75], [76], [77], [78]]. However, GA-crosslinked decellularized tissue grafts are associated with calcification and cytotoxicity, which pose severe limitations for clinical applications [79,80]. To address these drawbacks, Xiaotong et al. introduced procyanidins, which are natural compounds with cytocompatibility and have anti-calcification properties, in combination with GA to crosslink vascular dECM matrices [77]. The resulting crosslinked vessel-derived dECM demonstrated enhanced tensile strength, along with improved chemical and biological stability. Moreover, it exhibited comparable anti-calcification properties and cytocompatibility by supporting the adhesion and proliferation of human umbilical vein endothelial cells.

Compared to GA, EDC offers several advantages by facilitating amide bond formation between the carboxylic and amino groups of dECM without inducing cytotoxic reactions [81,82]. Moreover, EDC crosslinking enhances the mechanical strength without the risk of calcification, reduces inflammatory responses, and exhibits a slow degradation rate [81,83]. Cai et al. demonstrated that EDC crosslinking significantly improves the mechanical properties of vascular dECM scaffolds to achieve a comparable performance to that of native arteries [83]. Additionally, EDC treatment enables the immobilization of heparin on a dECM scaffold, which effectively inhibits thrombogenesis and presents an ideal scaffold for tissue-engineered vascular grafts.

Although synthetic chemical crosslinkers enable the robust crosslinking of dECM hydrogels, a general drawback is the potential cytotoxicity of unreacted residual crosslinking agents in vivo, which necessitates extensive washing to eliminate these remnants [84]. In this respect, natural crosslinking agents such as genipin, tannic acid (TA), and transglutaminase (TG) demonstrate superior biocompatibility and have been investigated for their efficacy in the crosslinking of dECM hydrogels. Genipin reacts exclusively with primary amine groups and interacts with a wide range of proteins in dECM hydrogels under neutral pH conditions [85]. Elder et al. achieved an approximate 90 % degree of crosslinking in cartilage dECM at a genipin concentration of 0.1 % [86]. Furthermore, by regulating the genipin concentration and crosslinking duration, the instantaneous compressive modulus was restored to levels comparable to those of native cartilage tissue. Notably, genipin is approximately 10,000 times less toxic than GA, which makes it a highly favorable option for tissue engineering applications [87,88]. Additionally, Vyborny et al. demonstrated that genipin crosslinking outperforms EDC in enhancing the rheological properties and biostability of a human umbilical cord-derived dECM hydrogel. This improvement highlights the potential of genipin as a promising scaffold for neural tissue applications [89].

Meanwhile, TA has been used to improve the structural stability of dECM hydrogels via hydrogen-bonding interactions and hydrophobic effects [90]. Kim et al. utilized TA to fabricate a finger-like villus structure for an intestinal model using small intestine-derived dECM [91]. They reported that the TA-crosslinked dECM hydrogels exhibited increased modulus and viscosity, along with improved printability and cell viability at optimal TA concentrations. However, excessively high TA concentrations led to unstable extrusion, which was attributed to non-homogeneous flow or elevated viscosity with excessive yield stress, which negatively affected the printing process (Fig. 2E).

The enzyme TG catalyzes selective crosslink formation between glutamine and lysine residues on collagen fibers [92,93]. Williams et al. demonstrated that, by varying the TG concentration, the Young's modulus of a dECM-based scaffold could be precisely adjusted between 13.7 and 32 kPa, which spans the range required to mimic both developing and mature heart tissues [94]. Moreover, You et al. successfully crosslinked tendon dECM hydrogels by using microbial TG (mTG), which is a non-calcium-ion-dependent enzyme with a lower molecular weight and fewer specific requirements for reaction substrates compared to mammalian TG [95]. Moreover, mTG demonstrates superior biocompatibility compared with other crosslinking agents, including GA and EDC, and exhibits a pro-healing effect on wounds. This effect was attributed to the promotion of collagen deposition by mTG, which facilitates the regeneration of scar tissue, thereby underscoring the potential of mTG as a highly efficient crosslinking agent for skin tissue engineering.

2.2.3. Ionic crosslinking

One promising strategy for enhancing the mechanical robustness of dECM hydrogels is to blend them with polymers that exhibit ionic properties, such as alginate. Alginate is a naturally derived linear polysaccharide with anionic properties, and is typically obtained from brown algae [96]. The gelation mechanism of alginate, known as the egg-box model, involves ionic bridging between two carboxyl groups on adjacent alginate chains with divalent ions (e.g., calcium, barium, strontium) [97,98]. The rapid gelation of alginate allows the dECM hydrogels to maintain their shape with high fidelity. Recently, Potere et al. developed a dECM-alginate composite bioink reinforced through a two-step strategy for fabricating aorta-like constructs [99]. The method involved blending dECM with alginate, where primary internal crosslinking was induced using calcium carbonate and glucono-δ-lactone. This approach enhanced the storage modulus and printability of the dECM bioink during the pre-printing phase, thereby facilitating precise shape morphing of the printed structure. After printing, the construct was immersed in a calcium chloride solution to trigger secondary ionic crosslinking of the alginate component. This dual-crosslinking process achieved a high degree of fidelity between the printed structure and the intended design, while maintaining a wall thickness comparable to that of the native aorta. Despite these advantages, a major disadvantage of using ions for gelation is the challenge of controlling the gelation kinetics, which often results in non-uniform structures [100]. Additionally, alginate lacks cell binding and adhesion capabilities, which arise from its polysaccharide-based composition [101].

2.2.4. Integrated sequential crosslinking

Recently, various crosslinking strategies have been integrated to further enhance the mechanical properties of the dECM hydrogels through sequential crosslinking. Hwang et al. developed a dual chemically and thermally crosslinked dECM hybrid bioink [102]. The primary chemical crosslinking was achieved via an aza-Michael addition reaction by combining the methacrylate groups of hyaluronic acid methacrylate (HAMA) with the amine groups of vascular tissue-derived dECM. In this process, dECM acted as a Michael donor, thus enabling chemical crosslinking without the need for an exogenous catalyst. The resulting improvement in printability enabled the successful printing of multi-layered patches with various degrees of crosslinking, thus allowing for the customization of the chemical crosslinking densities in order to modulate the release profiles of loaded growth factors. After printing, the patches were subjected to thermal crosslinking to further enhance their structural integrity. Meanwhile, Gao et al. engineered hollow vascular tissues by leveraging a dual crosslinking strategy involving ionic and thermal mechanisms to obtain a hybrid bioink composed of vascular tissue-derived dECM and alginate [47,103,104]. For example, a coaxial bioprinting technique was used to fabricate perfusable tubular structures with various pattern designs (Fig. 2F) [47]. In this approach, the hybrid bioink encapsulating endothelial cells was extruded into the outer layer, while the sacrificial ink containing calcium ions was extruded into the inner layer, simultaneously crosslinking the alginate hydrogel and maintaining the hollow tubular structures during extrusion. The printed tissue was both perfusable and functional, as demonstrated by the sprouting of neovessels in response to proangiogenic factors [104]. These step-wise multiple crosslinking approaches can improve the mechanical strength of dECM bioinks; however, they are often time-consuming and may result in heterogeneity within the hydrogel networks [105,106].

3. Photocrosslinking system requirements for robust polymerization and superior spatiotemporal control of dECM-based bioinks

Photocrosslinking has emerged as a cutting-edge bioink crosslinking strategy that is capable of significantly enhancing the mechanical robustness of dECM bioinks within a short time, thereby enabling the precise bioprinting of intricate tissue architectures [107]. Recently, dECM-based photocrosslinkable hydrogel systems have been actively developed and optimized to mitigate their inherent mechanical limitations, thereby enabling the construction of volumetric and structurally complex tissue models with high shape fidelity, which offer distinct advantages over conventional dECM crosslinking strategies (Table 1).

Table 1.

Comparison of conventional and light-activated crosslinking strategies for dECM-based bioinks.

Crosslinking strategy	Advantages	Disadvantages	Ref
Thermal	•Biocompatible	•Slow gelation kinetics	[64,65]
	•Easily implementable process	•Low mechanical strength
	•No need for additional crosslinkers	•Limited control over final geometry
Chemical	•Excellent mechanical integrity	•Complex reaction conditions	[77,172]
	•Strong and stable covalent bond formation	•Potential cytotoxicity of residual crosslinkers
	•Crosslinking tunability via stoichiometric control
Enzymatic	•Site-specificity	•Slow gelation kinetics	[173,174]
	•High cytocompatibility	•Limited substrate versatility
	•Adjustable gelation time	•Potential immune response due to residual enzyme
Ionic	•Adjustable mechanical strength	•Weak and unstable hydrogel networks	[84,100]
	•Rapid gelation under mild conditions	•Difficult to precisely control crosslinking density
	•Reversible network formation with self-healing capabilities	•Sensitive to ion concentrations and environment
Photo	•Excellent mechanical integrity	•Potential phototoxicity depending on light wavelength and exposure duration	[14,19,142]
	•Spatiotemporal control over crosslinking
	•Rapid crosslinking with high resolution fidelity

A typical photocrosslinking system is comprised of the following three fundamental elements: (i) a photoreactive polymer, (ii) a photoinitiator, and (iii) a light source. The present section explores the diverse photocrosslinking strategies incorporated into dECM hydrogels, with particular attention to the integration of photocrosslinkable functional groups and the crosslinking mechanisms underlying the chemistry of recently-used dECM bioinks (Table 2). Additionally, it examines key considerations and pivotal parameters required for the successful fabrication of advanced tissue constructs through photocrosslinking strategies using 3D bioprinting technology.

Table 2.

Photocrosslinking system utilized in dECM bioinks.

Photoreactive group	Polymerization mechanism	Photoinitiator	Light source	Ref
Methacrylates	Free-radical chain polymerization	Irgacure 2959	UV light	[120,[175], [176], [177], [178], [179]]
		LAP	Visible light	[4,118,119,[180], [181], [182], [183], [184], [185], [186]]
		Riboflavin	Visible light	[157]
		Eosin Y	Green light	[160,187]
		ethyl(2,4,6-tri- methylbenzoyl) phenylphosphinate	Visible light	[188]
		propidium iodide (PI)	UV light	[189]
		2-hydroxy-1-[4-(2-hydroxyethoxy)phenyl]-2-methyl-1-acetone	Blue light	[190]
Thiol-ene	Step-growth reaction	Irgacure 2959	UV light	[126]
		LAP	UV light	[[127], [128], [129],191]
Phenol	Redox reaction	Ru/SPS	Visible light	[12,13,16,139,140,148,151,152,192,193]
		Riboflavin	UVA light	[149,150,156]
		Eosin Y	Green light	[159]

3.1. Photocrosslinkable functional groups and the versatility of dECM crosslinking

Photocrosslinkable hydrogels are a specialized class of hydrogels that incorporate photoreactive functional groups that render them capable of forming covalent bonds and undergoing crosslinking upon exposure to light. A variety of chemically driven photocrosslinking strategies have been extensively developed and, more recently, successfully adapted for application in dECM hydrogels. The most commonly used photoreactive functional groups in dECM photocrosslinking reactions are methacrylates, thiol-enes, and phenol groups (Fig. 3). These functional groups can be integrated into dECM-based hydrogels either by blending with polymers that inherently contain the desired functional groups or through the chemical modification of the dECM itself.

Fig. 3.

Photocrosslinkable functional groups commonly used dECM-based hydrogels.

(A) Mechanism of free-radical chain polymerization. Reproduced with permission from Ref. [19]. Copyright 2020 American Chemical Society. (B) Crosslinking between methacrylate groups under light exposure. (C) Mechanism of radical-mediated thiol-ene photocrosslinking. Reproduced with permission from Ref. [19]. Copyright 2020 American Chemical Society. (D) Crosslinking of thiol and alkene groups under light exposure. (E) Mechanism of photomediated redox polymerization. Reproduced with permission from Ref. [19]. Copyright 2020 American Chemical Society. (F) Dityrosine bond formation under light exposure.

3.1.1. Methacrylate-mediated photocrosslinking

Methacrylate, often used as a generic term encompassing both methacrylate and acrylate groups, represents a prevalent reactive functional group employed for the photocrosslinking of dECM bioinks [108]. Methacrylates readily polymerize due to the high reactivity of their vinyl groups (i.e., C=C double bonds) [108]. Photocrosslinking via methacrylate occurs through a photoinitiated free-radical chain polymerization mechanism comprised of three primary stages, namely: initiation, propagation, and termination [19]. Upon light irradiation, photoinitiators undergo photolysis (light-induced cleavage) to produce free radicals that react with the methacrylate groups by propagating through the unreacted double bonds. During the propagation phase, the double bonds are consumed, thereby elongating the polymer chain. In the final termination stage, the reactive radicals are neutralized to form stable covalent bonds that effectively crosslink the dECM hydrogel network. Although both methacrylate and acrylate polymerize through free-radical mechanisms, they exhibit distinct polymerization kinetics owing to the differences in their molecular structures. Acrylate is generally more reactive than methacrylate because the hydrogen atom on the α-carbon leads to significantly lower steric hindrance compared with the methyl group in methacrylates [109].

Methacrylates are not inherently present in dECM hydrogels; therefore, various strategies have been employed to incorporate methacrylate groups into dECM-based bioinks while preserving the biochemical microenvironments of native tissues (Table 3). One common approach involves integrating synthetic polymers that contain methacrylate groups, such as polyethylene glycol diacrylate (PEGDA), into the dECM hydrogels, where photocrosslinking occurs in the synthetic polymer network rather than in the dECM itself [110,111]. Shin et al. combined PEGDA with cardiac dECM to enable photopolymerization of the PEGDA component, while the dECM served as a bioactive material. Laponite-XLG nanoclay was further incorporated to enhance extrusion stability [111]. This approach yielded a bioink that demonstrated high fidelity in extruded filaments and maintained structural integrity, thereby enhancing the mechanical durability of the construct after photocrosslinking (Fig. 4A).

Table 3.

Methacrylate-mediated crosslinking system of dECM bioinks.

Bioink composition	Additives	Photoinitiator	Wavelength (nm)	Printing technique	Ref
heart dECM + PEGDA	laponite nanoclay	LAP	405	Extrusion	[111]
annulus fibrosus dECM + PEGDA	–	LAP	405	–	[110]
liver dECM + GelMA	–	LAP	365	DLP	[115]
cartilage dECM + HAMA	–	Ru/SPS	450	–	[194]
cartilage dECM + HAMA	–	Eosin Y	515	Extrusion, DLP	[187]
cartilage dECM + HAMA	–	Propidium iodide	450	–	[189]
cartilage dECM + HAMA-SBMA	–	LAP	365	–	[195]
Vagina dECM + silk + GelMA	–	2-hydroxy-1-[4-(2-hydroxyethoxy)phenyl]-2-methyl-1-acetone	Blue light	Extrusion	[190]
meniscus dECM + GelMA + HAMA	–	LAP	365, 405	extrusion	[196]
kidney dECM-MA	–	Irgacure 2959	365	Extrusion	[116]
SIS dECM-MA	Graphene oxide, ascorbic acid	Riboflavin	405	Extrusion	[117]
Liver dECM-MA + PCL-MA	Orasol yellow dye	ethyl(2,4,6-tri- methylbenzoyl) phenylphosphinate	Visible light	DLP, SLA	[188]
methacrylated bone dECM + alginate	–	Irgacure 2959	UV	Extrusion	[197]
Cartilage dECM-MA + GelMA	–	Irgacure 2959	312	–	[198]
Methacrylated pulmonary dECM + HAMA	–	LAP	365	–	[199]
Cartilage dECM + HAMA	TEOA, NVP	Eosin Y	515	Extrusion, DLP	[187]
methacrylated bone and cartilage dECM + ColMA	–	LAP	365	–	[200]
cartilage dECM-MA + gelatin + HA		Irgacure 2959	UV	Extrusion	[171]

HAMA-SBMA: hyaluronic acid methacrylate-sulfobetaine methacrylate; PCL-MA: poly (ε-caprolactone) methacrylate; TEOA: triethanolamine; SLA: stereolithography.

Fig. 4.

Methacrylate-mediated photocrosslinking utilizing dECM hydrogel.

(A) Blending with synthetic polymers containing methacrylates for versatile bioprinting (scale bar: 1 cm). Reproduced with permission from Ref. [111]. Copyright 2020 Acta Materialia Inc. (B) Gelatin methacrylate composite dECM hydrogel mimicking tissue-specific architectures (scale bar; i: 250 μm, ii: (left) 500 μm, (right) 250 μm, iii: 1 mm, iv: 200 μm). Reproduced with permission from Ref. [115]. Copyright 2018 Elsevier Ltd. (C) Chemical modification of methacrylated dECM enabling photocrosslinking without blending with other hydrogels (scale bar; i: 5 mm, ii: 10 mm). Reproduced with permission from Ref. [171]. Copyright 2020 Acta Materialia Inc. (D) dECM-MA crosslinking electrospun fibers to improve structural stability and enable uniaxial cell alignment. Reproduced with permission from Ref. [120]. Copyright 2019 American Chemical Society.

In addition to synthetic polymers, methacrylated derivatives of natural polymers are often combined with dECM bioinks to produce composite systems in which photocrosslinking occurs in the methacrylated natural polymer components, while the dECM provides tissue-specific bioactivity. Since the first synthesis report of gelatin methacrylate (GelMA), numerous methacrylated natural biomaterials, including collagen methacrylate (ColMA), HAMA, and methacrylated silk fibroin, have been developed [18,[112], [113], [114]]. Yu et al. demonstrated that the incorporation of GelMA into a dECM hydrogel enables tuning of the Young's modulus under ultraviolet (UV) light, along with high-resolution printing, thus allowing the biomimetic patterning of key histological features such as precise parallel lines for aligned myocardium in heart tissue modeling and hexagonal lobular structures that mimic the architecture of liver lobules [115]. In addition, the researchers demonstrated that light-based bioprinting allows for the rapid fabrication of microscale tissue designs, thus promoting spontaneous cellular reorganization into predefined structures through biophysical cues to obtain physiologically functional tissues (Fig. 4B).

Recent advancements in the modification of dECM hydrogels with methacrylate groups (dECMMA) to confer photocrosslinking capabilities have made extensive use of methacrylic anhydride (MA), methacrylic acid, and glycidyl methacrylate (GMA) [[116], [117], [118]]. A comparative analysis of the various materials used to impart methacrylate groups to dECM hydrogels was performed by Almalla et al. [118]. Notably, dECM hydrogels synthesized using GMA exhibit higher solubility in aqueous environments, along with improved cytocompatibility compared to those synthesized with MA due to the hydrophilicity of the free hydroxyl groups. Meanwhile, Visscher et al. compared dECMMA to GelMA in terms of cytocompatibility and cell functionality, beyond the mechanical properties, as dECMMA was sufficient for bioprinting [119]. The dECMMA bioink demonstrated enhanced chondrocyte cytocompatibility over GelMA, as evidenced by the typical triangular and ovoid morphology of chondrocytes in dECMMA constructs, which was not observed in GelMA (Fig. 4C). In addition to extrusion printing, dECMMA hydrogels have been employed in other printing techniques, such as electrospinning, to promote anisotropic alignment in muscle tissues. Lee et al. chemically modified dECM with methacrylate to improve the structural stability, and fabricated a nanofibrous dECM-based structure by using electrospun dECMMA with UV exposure during the electrospinning process [120]. The resulting dECMMA nanofibers were deposited on fibrillated poly(lactide-co-glycolide) constructs with uniaxial alignment, thereby significantly enhancing muscle regeneration (Fig. 4D).

Methacrylate-mediated photocrosslinking has become a widely adopted method for crosslinking various biomaterials. Methacrylate-based photocrosslinkable hydrogels offer tunable mechanical properties and degradation rates by modulating factors such as the degree of methacrylation. However, the basic mechanism of radical polymerization, including stepwise reactions, initiator decomposition, and chain transfer to the monomer, is accompanied by complex reaction kinetics [19,121]. Moreover, a notable challenge in methacrylate-based systems is oxygen inhibition, where oxygen scavenges the radicals required for crosslinking, thus resulting in incomplete network formation [122,123]. This affects printing fidelity and hampers the maintenance of construct shapes. Additionally, controlling the crosslinking kinetics in free-radical polymerization systems is challenging, thus leading to structural heterogeneities in the polymer network along with residual unreacted radicals.

3.1.2. Thiol-ene chemistry

Thiol-ene chemistry has recently gained prominence in biofabrication due to its ability to offer precise control over the crosslinking process with enhanced reaction kinetics, particularly when compared to free-radical polymerization processes such as methacrylate-based systems [124]. This chemistry involves reactions between thiols and unactivated carbon-carbon multiple bonds (i.e., alkene, alkyne), and is categorized into the following two primary types: (i) thiol-ene click reactions and (ii) thiol-Michael addition [114]. Thiol-ene click reactions follow a step-growth free-radical mechanism that is initiated by radicals generated through photoinitiator activation. Upon light absorption, the photoinitiator releases radicals, which then transfer energy to thiol groups. These activated thiols function as nucleophiles, attacking the carbon-carbon double bonds to form new covalent bonds. Conversely, thiol-Michael addition is a nucleophilic reaction that occurs in the presence of a catalyst, thus resulting in the direct formation of a covalent bond between the thiol and carbon-carbon double bonds without the need for light activation or a photoinitiator. Nevertheless, thiol-Michael addition is often coupled with a photoinitiator in order to enhance the spatial and temporal control during bioprinting via secondary photocrosslinking [125].

The customization of a dECM hydrogel can involve blending with a photocrosslinkable hydrogel containing thiol or alkene groups (Table 4). Alkene groups are typically introduced by using alkene-conjugated tetra-polyethylene glycol (PEG) polymers such as 4-arm-PEG-acrylate, 4-arm-PEG-alkyne, or polyethylene glycol-alpha methacrylate (PEGαMA) [[126], [127], [128]]. Similarly, thiolated synthetic polymers such as 4-arm-PEG-thiol (PEG4SH) are incorporated to enable efficient crosslinking [129]. Skardal et al. integrated the commercially-available thiol-modified hydrogel toolkit, HyStem-HP, into dECM hydrogels [126]. This toolkit is comprised of thiolated hyaluronan, gelatin, and heparin, which were combined in a two-stage polymerization strategy. Initially, the thiolated natural polymers were crosslinked with PEGDA to form a suitably soft, low-viscosity hydrogel for extrusion. Subsequently, secondary crosslinking was achieved via photopolymerization using 4-arm and 8-arm alkyne-conjugated PEG, thus resulting in increased hydrogel stiffness and improved structural stability (Fig. 5A).

Table 4.

Thiol-ene crosslinking system of dECM bioinks.

Bioink composition		Additives	Photoinitiators	Wavelength (nm)	Printing technique	Ref
For thiol groups	For alkene or Alkyne groups
Liver, cardiac, skeletal muscle-derived dECM + Heprasiland Gelin-S	PEG 4-arm Acrylate, PEG 4-arm Alkyne, PEG 8-arm Alkyne	–	Irgacure 2959	365	Extrusion	[126]
Thiolated lung dECM	PEGαMA	DTT, CGRGDS	LAP	365	–	[127]
Thiolated human dECM	PEGαMA	DTT, CGRGDS	LAP	365	–	[128]
Thiolated lung dECM	PEGαMA	DTT, CGRGDS	LAP	365	–	[191]
PEG4SH	Norbornene-modified SIS dECM	TEA, DTT, tartrazine	LAP	365	Extrusion, DLP	[129]

DTT: 1,4-Dithiothreitol; CGRGDS: Cys-Gly-Arg-Gly-Asp-Ser peptide.

Fig. 5.

Thiol-ene chemistry mediated photocrosslinking utilizing dECM hydrogel.

(A) Incorporation of diverse functional groups within thiol-ene click chemistry for tunable stiffness in printed constructs. Reproduced with permission from Ref. [126]. Copyright 2015 Acta Materialia Inc. (B) Thiolated dECM undergoing thiol-Michael addition crosslinking, followed by a light-activated click reactions for hydrogel stiffness regulation. Reproduced with permission from Ref. [128]. Copyright 2023 American Chemical Society. (C) Norbornene-conjugated dECM applied in diverse light-activated printing techniques for fabricating various structures. Reproduced with permission from Ref. [129]. Copyright 2024 Wiley‐VCH GmbH.

Recently, dECM hydrogels have been directly functionalized by introducing thiol and alkene groups in order to expand their crosslinking versatility. Thiol groups have been incorporated by converting free amine groups into thiols using reagents such as Traut's reagent (2-iminothiolane hydrochloride) [127,128], while alkene groups have been incorporated through norbornene conjugation [129]. Hewawasam et al. demonstrated this strategy by modifying dECM with thiol groups and crosslinking the thiolated dECM hydrogels using PEGαMA-blended bioinks [128]. This process involved an initial thiol-Michael addition reaction between the thiolated dECM and PEGαMA, followed by light-induced secondary crosslinking, thus resulting in stiffer hydrogels and enhanced fibroblast activation (Fig. 5B). In a similar approach, Duong et al. functionalized dECM with norbornene groups (dECM-NB) by reacting the amine groups of dECM with carbic anhydride under acidic conditions to create alkene-functionalized hydrogels [129]. These were subsequently blended with PEG4SH and a photoinitiator to enable the fabrication of tubular structures with high aspect ratios. Additionally, the bioactivity of the dECM-NB was shown to enhance cell-ECM interactions, thereby demonstrating its potential for advanced tissue engineering applications (Fig. 5C).

Thiol-ene chemistry is characterized by rapid kinetics and a step-growth mechanism, which enables efficient and selective crosslinking within seconds under mild conditions while achieving high functional-group conversions even at low concentrations of radical initiators [130]. A key advantage of thiol-ene coupling is its oxygen-insensitivity, which sets it apart from methacrylate-mediated photocrosslinking [131]. Furthermore, precise control over hydrogel network formation within the crosslinked hydrogels can be achieved based on the specific functional groups participating in the thiol-ene reaction, such as thiol-acrylate or thiol-norbornene [132]. Among the various enes (vinyl groups) used in thiol-ene chemistry, NB has gained increasing attention owing to its greater reactivity and superior precursor stability compared with methacrylates [133,134]. This is primarily because of the lower electron density in NB, which makes it highly reactive owing to the release of ring strain, whereas methacrylates are relatively unreactive owing to intermediate radical stabilization [135]. This provides better control over the crosslinking process, leading to enhanced printing resolution of the resulting constructs [133].

Beyond the type of functional group used, the stoichiometric ratio of thiol to alkene plays a critical role in determining the mechanical strength of the hydrogel. An imbalance, whether an excess of thiol or alkene, can result in insufficient crosslinking or poor mechanical properties [136]. However, residual thiol or alkene groups do not necessarily indicate a drawback; instead, they can serve as additional sites for subsequent biological modifications, thereby further enhancing the functionality of the bioinks [137].

3.1.3. Phenol-mediated photocrosslinking

While less extensively studied, phenol-mediated photocrosslinking systems hold significant promise for 3D bioprinting applications. Phenol groups are naturally present in the reactive amino acid tyrosine, which is present in collagen, the predominant protein in dECM. This inherent composition allows dECM to be readily photocrosslinked into hydrogels through phenolic oxidation without any chemical modification [12,15]. In the presence of a photoredox catalyst (i.e., a photoinitiator), the phenol groups undergo redox reactions with reactive species. Upon light exposure, the photoinitiator absorbs photons and enters an excited state, enabling the abstraction of a hydrogen atom from the substrate, which donates electrons and generates hydrogen-donor radicals. These radicals initiate the crosslinking process by forming covalent bonds between the phenolic groups through oxidation [19,138].

Recent studies have demonstrated the effective application of phenol-mediated photocrosslinking systems to enhance dECM bioinks in 3D bioprinting (Table 5). Kim et al. developed light-activated dECM bioinks using a tris(2,2-bipyridyl) dichlororuthenium(II) hexahydrate/sodium persulfate (Ru/SPS) photoinitiator (designated as dERS) [12]. A liquid chromatography-mass spectrometry analysis demonstrated that the tyrosine residues were reduced upon photocrosslinking of the dERS, thus indicating the formation of dityrosine bonds. The crosslinked dERS exhibited superior shape fidelity, forming cylindrical structures with high aspect ratios compared to the non-crosslinked dECM. This facilitated the fabrication of intricate constructs, including curved corneal and heart chamber models, while maintaining high cell viability and functionality (Fig. 6A). Additionally, Han et al. employed a dERS bioink in coaxial printing to construct a tubular structure with hollow lumens [139], while Lian et al. used a tomographic printing technique to direct dynamically evolving light patterns onto specific areas of photocrosslinkable bioinks within a rotating transparent vial [16]. The latter technique enabled the use of lower concentrations of dECM bioink, with the entire printing process being completed in under 90 s, which is significantly faster than conventional digital light processing (DLP) techniques. The final constructs, including hollow cubes, screws, and tissue-specific geometries such as hearts, menisci, and bones, were produced with exceptional resolution (Fig. 6B).

Table 5.

Phenol-mediated crosslinking system of dECM bioinks.

Bioink composition	Photoinitiator	Wavelength (nm)	Printing technique	Ref
Heart or cornea dECM	Ru/SPS	400–450	Extrusion, DLP	[12]
Colon dECM	Ru/SPS	400–450	Extrusion	[147]
SIS dECM	Ru/SPS	400–450	Extrusion	[13]
Heart or meniscus dECM	Ru/SPS	525	Volumetric printing	[16]
Heart dECM	Riboflavin	370	Extrusion	[150]
SIS dECM	Riboflavin	420	Extrusion	[149]
Adipose dECM	Ru/SPS	402	Extrusion	[148]
Liver or heart dECM	Riboflavin	350–370	–	[156]
Skin dECM	Ru/SPS	400	–	[193]
Cartilage dECM + HA-Tyr	Ru/SPS	405	DLP, FLight printing	[192]
Pancreas dECM + tyrosine modified HA	Ru/SPS	406	–	[152]

HA-Tyr: HA tyramine; Flight printing: Filamented Light printing.

Fig. 6.

Phenol-mediated photocrosslinking within dECM-based bioinks.

(A) Ruthenium-based photoinitiator mediating the crosslinking of intrinsic phenol groups within dECM for high shape fidelity. Reproduced with permission from Ref. [12]. Copyright 2021 Wiley‐VCH GmbH. (B) Versatile fabrication of complex structures using dECM bioinks supplemented with Ru/SPS through volumetric printing techniques. Reproduced with permission from Ref. [16]. Copyright 2024 Wiley‐VCH GmbH. (C) Triple crosslinking strategies utilizing dECM and its gelatinized derivatives for printed structures with high shape fidelity, toughness, and resilience (scale bar; i: 0.5 cm, ii: 1 cm). Reproduced with permission from Ref. [13]. Copyright 2024 Royal Society of Chemistry.

In addition to these advancements in printing techniques, recent developments in crosslinking strategies have focused on further enhancing the mechanical properties of dECM bioinks, including stretchability, toughness, and resiliency. Han et al. introduced a triple crosslinking strategy by incorporating thermally denatured gelatinized dECM (GeldECM) as a rheological modifier to enhance the mechanical strength of dECM bioinks without introducing exogenous materials that might compromise the bioactivity of the pristine dECM [13]. This approach utilized the physical entanglement properties of gelatin at low temperatures, along with Ru/SPS-mediated photocrosslinking and thermal crosslinking of the dECM bioinks, to significantly improve the printability and structural fidelity. The incorporation of GeldECM imparted more ductile, elastic, and tough characteristics compared to those of the dECM hydrogels alone, thus enabling the fabrication of robust, centimeter-scale tubular structures that are capable of withstanding external mechanical stress (Fig. 6C).

As well as enhancing the hydrogel stiffness, the phenol-mediated photocrosslinking system can also generate adhesive bioinks. Kim et al. demonstrated that GeldECM can produce L-3,4-dihydroxyphenylalanine (i.e., Dopa), which is a critical component of natural adhesives, via Ru/SPS-mediated oxidation of tyrosine [140]. The resulting photocrosslinked dECM-based hydrogel exhibits exceptional adhesive strength, comparable to that of commercial tissue glue. When used as a sealant, it facilitates corneal tissue remodeling by providing regenerative biofactors and adherable proteins.

In brief, phenol-mediated photocrosslinking offers a promising approach for dECM bioinks by leveraging the intrinsic phenol groups present in the amino acid tyrosine, thereby eliminating the need for additional chemical modifications. This approach not only reduces the production time but also preserves the bioactivity and native microenvironment of dECM. Nevertheless, the limited availability of tyrosine residues in collagen constrains the number of reactive sites for photocrosslinking [141].

3.2. Photoinitiators and their compatibility with dECM bioinks

Photoinitiators are compounds that trigger the photopolymerization process in bioinks. Upon exposure to light, these molecules absorb photons, transition to an excited state, and transfer their energy to a co-initiator, thereby generating reactive radical species. These radicals subsequently interact with double bonds within the polymer side chains to drive crosslink formation [32]. Based on their activation mechanisms, photoinitiators are generally classified into two main categories, designated as Type I and Type II, each with distinct activation wavelengths (Fig. 7A). Type I photoinitiators, such as 2-hydroxy-4′-(2-hydroxyethoxy)-2-methylpropiophenone (Irgacure 2959) and lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), function by absorbing incident photons and reaching an excited singlet or triplet state. This excitation induces bond cleavage within the molecule via a homolytic mechanism, thus resulting in the generation of free radicals that propagate the polymerization reaction (Fig. 7B) [142,143]. Irgacure 2959 is among the earliest Type I photoinitiators, and is activated under UV light, whereas LAP is a widely utilized photoinitiator in tissue engineering due to its activation under visible light [143]. Comparisons between these two commonly used Type I photoinitiators have been extensively conducted regarding cell viability and crosslinking efficiency. Xu et al. demonstrated that Irgacure 2959 exhibits significantly higher cytotoxicity than LAP at elevated concentrations [144]. Additionally, the crosslinking efficiency of Irgacure 2959 is lower than that of LAP, thus resulting in hydrogels with larger pore sizes, faster degradation rates, and higher swelling ratios.

Fig. 7.

Photoinitiators utilized for photopolymerization of dECM bioinks.

(A) Type I and II photoiniaitors and their activation wavelength. (B) Mechanism of Type I photoinitiator photopolymerization process. (C) Mechanism of Type II photoinitiator photopolymerization process.

By contrast, Type II photoinitiators generate free radicals through a multi-step reaction sequence that requires the presence of a co-initiator. Upon photon absorption and excitation to an activated triplet state, the Type II initiator transfers an electron to the co-initiator, followed by proton release to generate reactive radicals (Fig. 7C) [143,145]. Notably, several promising visible-light absorbing photoinitiators, such as Ru and riboflavin (vitamin B2), belong to this category [146]. These photoinitiators are particularly well suited for phenol-based redox polymerization processes [12,13,16,140,[147], [148], [149], [150], [151], [152]]. Recently, Ru-based photoinitiators have garnered attention for their capacity to function as visible light-harvesting molecules, efficiently absorbing within the 400–450 nm range. As part of a bimolecular system, Ru requires a co-initiator such as SPS or ammonium persulfate to initiate the photocrosslinking mechanism. Similarly, riboflavin can form covalent bonds between oxidized residues, including tyrosine, lysin, histidine, arginine, and methionine in collagen, under UVA irradiation [153,154]. Traditionally used for collagen crosslinking, this method has recently been extended to dECM bioinks [149,150,[155], [156], [157]]. Riboflavin-initiated photocrosslinking solidifies the dECM bioinks with high structural fidelity, and the mechanical properties can be easily adjusted by modulating the riboflavin concentration [149,150]. However, the underlying mechanism of UVA and riboflavin photocrosslinking remains a topic of ongoing investigation. Recent studies suggest that tyrosine residues play a central role as crosslinking sites, while riboflavin functions as both a catalyst and a competitive inhibitor in the process [153]. Type II photoinitiating systems, which rely on bimolecular interactions, are less efficient than Type I systems, which involve a single component. Thus, it is a complex matter to determine the optimal concentration ratio between the Type II photoinitiator and its co-initiator in order to maximize the crosslinking efficiency without inducing cell toxicity from either the photoinitiator itself or the byproducts generated during the photocrosslinking process [143].

Due to their high absorption coefficients in the spectral region of the excitation light, photosensitizers (dyes) such as flavin mononucleotide, rose bengal, and Eosin Y are commonly used to further catalyze the photopolymerization of tyrosine and tyramine functional groups [158]. Eosin Y is the most widely used photosensitizer, and is commonly employed in conjugation with triethanolamine (TEA) as a co-initiator and N-vinylpyrrolidone (NVP) as a co-monomer to enhance the photopolymerization kinetics [159,160]. Upon light irradiation, Eosin Y transitions from its ground state to a triplet energy state, which enables it to abstract hydrogen atoms from amine-functionalized co-initiators such as TEA, thereby generating TEA radicals that subsequently initiate the crosslinking. To address the slow nature of this reaction and expedite gelation, accelerants such as NVP are generally introduced. Notably, the activation of Eosin Y is visually indicated by a color shift from red to yellow [159]. Bejleri et al. reported that the Eosin Y system allows the most effective formation of structurally reliant and viable dECM-based patches compared to other photoinitiator systems, such as Ru/SPS and Irgacure 2959 [160]. This system exhibits its strongest absorption peak in the green light range (480–550 nm), which is advantageous for cytocompatibility during light exposure due to its visible light-activated crosslinking properties. Nonetheless, a significant challenge in using Eosin Y lies in determining the optimal concentrations of photoinitiator, co-initiator, and co-monomer to ensure effective crosslinking while maintaining non-cytotoxic thresholds [19]. This presents practical challenges for the broader application of Eosin-Y in bioprinting technologies. Given the potential cytotoxicity of light exposure and the free radicals generated during the process, the selection of an appropriate photoinitiator is critical for balancing crosslinking efficiency with biocompatibility in tissue engineering applications [146].

3.3. Light sources and tunable parameters for activation

The activation wavelength of a photoinitiator is dictated by both the photoinitiator itself and the photoresponsive material, which necessitates the optimization of the light source in order to achieve high photocrosslinking efficiency. Each photoinitiator has a distinct light absorption spectrum across a wide wavelength range. The probability of light absorption by the photoinitiator depends on its extinction coefficient (ε_λ), with higher ε values indicating greater absorption capacity at a given wavelength [161]. This facilitates photoinitiator activation, thereby improving the efficiency of dECM hydrogel crosslinking.

Among the available light wavelengths, UV light is widely utilized for photocrosslinking, largely due to its compatibility with methacrylate groups across various polymers. However, UV light poses potential limitations, including risk of cellular genetic damage [162,163]. Additionally, its short wavelength limits the penetration depth and therefore hinders the fabrication of volumetric 3D bioprinted constructs. By contrast, visible light offers longer wavelengths, thereby enabling deeper tissue penetration and improving the photocrosslinking efficiency while minimizing the risk of cellular genetic damage compared to UV exposure [123,164].

Beyond the choice of wavelength, several key parameters such as light intensity, exposure time, and the distance between the light source and the tissue play a critical role in determining the crosslinking kinetics and maintaining cell viability [19,131]. The degree of bioink crosslinking can be enhanced by increasing the light intensity, prolonging the irradiation time, and reducing the distance. However, excessive light exposure or prolonged irradiation can cause cell death. Moreover, if the light source is placed too close, the heat generated during illumination may burn the cells, thereby compromising their viability. Careful optimization of these parameters allows photocrosslinking to achieve precise spatiotemporal control over bioink polymerization, thus ensuring a balance between crosslinking efficiency and cell viability.

4. Conclusions and future perspectives

The present review has highlighted the transformative potential of dECM as a foundational biomaterial in tissue engineering and 3D bioprinting technology. Its biochemical complexity and structural fidelity uniquely enable the reproduction of native tissue microenvironments to support critical cellular functions such as adhesion, proliferation, and differentiation. These attributes position dECM-based bioinks as indispensable tools for the fabrication of complex, multifunctional, and physiologically relevant tissue constructs, with far-reaching implications for regenerative medicine. However, despite its promise, the application of dECM bioinks faces significant challenges, including variability in production processes, insufficient mechanical properties, and inconsistent reproducibility. The absence of standardized protocols for dECM preparation and biofabrication further limits scalability and clinical translation, thereby necessitating strategic advancements in order to address these. Especially, for successful clinical translation, dECM bioinks must also meet several safety requirements, including the minimization of immunogenicity and tumorigenic potential. These concerns can be addressed by implementing GMP-compliant manufacturing processes and incorporating bioengineering approaches to ensure product consistency and patient safety [165]. To overcome these barriers, future research must focus on improving the manufacturing consistency, refining biofabrication strategies, and enhancing the structural fidelity and biological functionality of printed constructs. Among these technologies, photocrosslinking has shown particular promise, as it offers precise control over the gelation kinetics and mechanical properties, thus enabling the fabrication of geometrically intricate and biologically functional tissue architectures. However, photocrosslinking technology also presents inherent limitations that must be addressed. While providing high-energy short wavelengths, UV light sources pose the potential risk of genetic damage to cells, whereas long-wavelength light sources such as those in the visible spectrum offer improved cytocompatibility but often exhibit reduced reaction efficiency. Hence, advanced photoinitiators must be developed in order to bridge this gap. Both the selection and optimization of photoinitiators are critical factors, as they directly influence the uniformity of the crosslinked network, along with the reaction kinetics and structural stability of the resulting constructs. Additionally, residual photoinitiators and byproducts may exhibit cytotoxicity, thus requiring meticulous post-processing to ensure biocompatibility. The development of far-wavelength light sources and corresponding photoinitiators is particularly promising, as these systems minimize cellular damage while enabling the fabrication of volumetric tissue constructs with enhanced structural integrity. Despite these advancements, the inherent challenges associated with photoinitiation, including efficient light penetration, crosslinking heterogeneity, and the balance between cytocompatibility and crosslinking performance, remain critical barriers. Hence, it will be essential to address these issues via innovative strategies, and to undertake continued research in order to unlock the full potential of photocrosslinking systems in the scalable fabrication of complex, functional tissue constructs. Recent advancements include the development of volumetric bioprinting techniques, which significantly enhance fabrication speed and resolution. Bernal et al. introduced a volumetric bioprinting method that enables the creation of geometrically complex, centimeter-scale structures within seconds by projecting dynamic light patterns from multiple angles into a photocrosslinkable bioink [166]. Furthermore, to promote the fabrication of highly aligned tissue architectures, Liu et al. developed a filamented light bioprinting technique in which patterned filamented light beams are projected through photocrosslinkable bioinks, enabling the formation of highly aligned hydrogel microfilaments within macroscale constructs, resulting in anisotropic tissue organization [167]. Moreover, to mimic the heterogeneity of native tissues more accurately, recent efforts have focused on developing multimaterial and multicellular printing techniques. These include the use of multivat systems, microfluidics-integrated vats for automated bioink exchange and multi-wavelength strategies for spatially controlled printing of heterogeneous materials [168].

Although these technological innovations enhance fabrication capabilities, parallel progress in evaluating the functional performance of the resulting bioprinted constructs remains essential. As bioprinting evolves toward tissue-level constructs, assessing the tissue-level functionality remains challenging. Recently, biohybrid systems, which integrate living cells with engineered nonliving components, have emerged as promising tools for the real-time monitoring of biosignals and tissue function [169,170]. Such progress will bridge the gap between laboratory research and clinical application, thereby transforming regenerative medicine by providing innovative therapeutic solutions for a wide array of medical needs. Ultimately, the continued development of dECM bioinks promises to revolutionize the field and unlock new possibilities for translational medicine, ultimately paving the way for advanced engineered tissues that will be capable of addressing diverse healthcare challenges.

CRediT authorship contribution statement

Minji Kim: Writing – original draft, Visualization, Investigation, Data curation, Conceptualization. Dayoon Kang: Writing – original draft, Visualization, Investigation, Data curation, Conceptualization. Hohyeon Han: Data curation, Conceptualization. Jinah Jang: Writing – review & editing, Supervision, Funding acquisition.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Data availability

No data was used for the research described in the article.