Digital light processing bioprinting: bioink innovations and applications in tissue and organ regeneration

Abstract

Digital light processing (DLP) bioprinting offers superior spatial resolution, rapid printing speed, and gentle cell handling compared to extrusion- and inkjet-based methods. By enabling layer-by-layer photopolymerization through projected light patterns, DLP provides precise control over microscale architecture, making it highly suitable for fabricating tissue- and organ-specific constructs. However, the true potential of DLP bioprinting depends not only on optical precision but also on the development of photocurable bioinks that translate light-driven fabrication into biologically relevant outcomes. Recent advances in bioink development (including methacrylated natural polymers, decellularized extracellular matrix (dECM)-derived hydrogels, hybrid nanocomposites, and cell-laden formulations) have significantly expanded the biofabrication landscape. These materials enable the creation of constructs that closely mimic native tissues such as corneas, cartilage, liver, and skeletal muscle in terms of structural and biochemical complexity. Furthermore, strategies such as photoabsorber modulation, rheological tuning, and dual crosslinking mechanisms have improved print fidelity and cytocompatibility. Despite this progress, translational barriers remain. Limited vascularization, poor scalability, and insufficient functional longevity continue to restrict clinical adoption. This review critically examines how bioink design paradigms in DLP bioprinting redefine the interface between light-based engineering and biological function. We also discuss emerging trends in stimuli-responsive and cell-instructive bioinks, offering insights into how to bridge laboratory innovation with clinically relevant tissue regeneration.

Introduction

Organ failure represents one of the most pressing challenges in modern medicine. Millions of people suffer from life-threatening conditions, including liver diseases and end-stage renal disease (ESRD), which necessitate immediate transplantation [1–5]. This critical deficit has been exacerbated by an aging global population and rising chronic diseases. In the United States, over 100,000 people are on transplant waiting lists, with 90,000 awaiting kidney transplants, and 13–17 patients die daily while waiting for an organ. Globally, only 10% of the actual demand for organ transplants is met [1, 6–8]. Current treatments (dialysis for renal failure and allogeneic transplantation) face severe limitations: dialysis is associated with cardiovascular morbidity, infection risks, and reduced quality of life, while transplantation suffers from donor shortage, immune rejection, and lifelong immunosuppression [3, 8, 9]. These challenges underscore an urgent unmet clinical need for transformative approaches that transcend conventional treatments.

Tissue engineering has emerged as a promising paradigm, evolving from early scaffold-based strategies to advanced regenerative medicine that harnesses the body’s intrinsic repair capabilities [10–12]. Within this landscape, three-dimensional (3D) bioprinting has revolutionized the field by enabling precise, layer-by-layer fabrication of complex, patient-specific biological constructs using living cells, biomaterials (bioinks), and bioactive factors [1, 13, 14]. Several bioprinting modalities exist. Extrusion-based bioprinting deposits continuous filaments of high-viscosity bioinks, enabling fabrication of mechanically robust constructs, but suffers from low resolution (>100 μm), prolonged shear stress, and slow printing speeds for intricate geometries [1, 15, 16]. Inkjet bioprinting offers high resolution and excellent cell viability through precise ejection of picoliter-sized droplets, yet is restricted to low-viscosity bioinks and faces nozzle clogging [1, 17, 18]. Laser-assisted bioprinting (LIFT) achieves remarkable resolution (~20 μm) and high viability without mechanical contact, but is limited by high cost, complexity, and slow printing speeds [1, 19, 20]. Two-photon polymerization provides unprecedented sub-micrometer resolution via localized femtosecond laser polymerization, though it is extremely slow and expensive [21]. A comprehensive comparison of these techniques is presented in Table 1.

Table 1.

Comparison of common 3D bioprinting techniques (excluding DLP)

Technique	mechanism	Advantages	Disadvantages	Typical application	references
Extrusion-based bioprinting	Pressure-driven continuous filament deposition of bioink	-High viscosity bioinks compatible -Mechanically robust structures -High cell density possible	-Lower resolution (typically > 100 µm) -Shear stress on cells reduces viability -Slower printing speed for complex geometries	-Large tissue constructs -Scaffold fabrication for cartilage and bone -Organ-on-a-chip models	[1, 7, 9, 22]
Inkjet-based bioprinting	Thermal or piezoelectric ejection of picoliter bioink droplets	-High printing resolution and accuracy -Good cell viability due to minimal shear stress -Cost-effective and high-throughput for small constructs	-Restricted to low-viscosity bioinks -Nozzle clogging issues -Limited structural integrity for large constructs	-issue models for drug screening -Micro-patterns for cell studies -Personalized therapeutics	[1, 7, 22, 23]
Laser-assisted bioprinting	Laser-induced forward transfer (LIFT) of bioink	-Exceptional resolution (~20 µm) -High cell viability due to minimal mechanical stress -Precise deposition of multiple cell types	-High complexity and cost of equipment -Slower printing speed compared to extrusion -Limited to thin layers	-Micro-tissue engineering -Intricate cellular patterns -Nerve tissue regeneration	[1, 7, 22]
Volumetric bioprinting	Photo-crosslinking throughout a volume of bioink simultaneousy	-High speed for complex 3D structures -Enables printing within cell-laden hydrogels -Preserves cell viability by minimizing exposure time	-Limited range of suitable photo-crosslinkable bioinks -Challenging for multi-material printing -Still in early developmental stages	-Rapid fabrication of complex organoids 37 -Embedded bioprinting of vascular networks -Organ-on-a-chip models	[24]
Two-photon polymerization (2PP) bioprinting	Focused femtosecond laser for localized polymerization	-Sub-micrometer resolution -Fabrication of highly intricate microstructures -Enables precise control over cellular microenvironment	-Extremely slow printing speeds -Very high cost of equipment -Limited to small-scale constructs	-Micro-scale tissue models - Fine cellular patterning - High-resolution scaffolds for regenerative medicine	[25–27]
Stereolithogry (SLA) bioprinting	A focused UV laser beam scans the surface of a photopolymer bioink vat, solidifying the object point-by-point, layer-by-layer	-High resolution and good surface finish -Well-established and commercially available -Can create complex internal geometries	-Slower than DLP (point scanning vs. layer projection)\ -Potential for UV-induced cell damage -Limited to low-viscosity photopolymer bioinks	Acellular scaffolds for bone and cartilage; microfluidic devices; anatomical models for surgical planning	[19, 28, 29]

While these conventional methods have advanced the field, they face fundamental trade-offs among resolution, cell viability, printing speed, and the ability to create vascularized, macro-scale tissues. The “tyranny of diffusion” restricts viable tissue constructs to a few hundred micrometers in thickness, necessitating integrated perfusable vascular networks for nutrient and oxygen transport [1, 9, 28]. Digital Light Processing (DLP) bioprinting has recently emerged as a transformative technology that overcomes many of these limitations. DLP uses dynamic mask projection to photo-crosslink entire bioink layers rapidly, offering high printing speed (seconds per layer), high resolution (25–100 μm), and excellent cell viability (>85–95%) due to the absence of shear stress [10, 29]. These unique attributes position DLP as a particularly promising approach for fabricating complex, multi-scale, vascularized tissue constructs.

Despite its promise, DLP bioprinting faces a central challenge: the balance between resolution and cell viability is critically governed by bioink chemistry; specifically, photo-crosslinking kinetics, photo-initiator toxicity, optical attenuation, and rheological properties. Unlike previous reviews that have separately addressed DLP hardware engineering or bioink formulation, a critical gap remains: no existing work has systematically investigated how printing parameters (e.g., light intensity, exposure time, layer thickness) and bioink properties (e.g., viscosity, photo-initiator concentration, crosslinking kinetics) collectively determine print fidelity, cell viability, and construct functionality. This review fills this gap by providing an integrated analysis of this interplay, which sets it apart from previous works by offering a unified framework that bridges light-based printing parameters with photo-crosslinkable bioink design principles,an aspect that has been largely overlooked in the existing literature. We evaluate (1) the fundamentals of DLP technology and critical printing parameters, (2) design principles for photo-crosslinkable bioinks including natural and synthetic polymers, (3) strategies for balancing crosslinking density with cell viability, and (4) emerging applications in vascularized and complex tissue constructs. By bridging the gap between light-based printing and bioink design, this review aims to guide the rational development of next-generation DLP bioprinting systems toward clinical translation.

Digitial light processing (DLP) bioprinting

Digital Light Processing (DLP) is an advanced 3D bioprinting technology that has garnered significant attention in the field of regenerative medicine. Due to its high precision, rapid printing speed, and ability to produce complex structures, it has emerged as a powerful tool for fabricating bio-scaffolds and printed organs. The origins of DLP technology date back to 1987, when it was invented by Dr. Larry Hornbeck, and it was later commercialized by Texas Instruments in 1996 [20]. DLP 3D printers utilize light to cure photosensitive materials (inks) [21]. DLP bioprinting has emerged as a superior alternative to address these specific limitations. Unlike extrusion systems, DLP utilizes a digital light projector to polymerize photocurable bioinks [30, 31]. A key component of most DLP systems is the Digital Micromirror Device (DMD), a microelectromechanical array composed of thousands to millions of individually controllable mirrors. The DMD selectively reflects projected light to generate a dynamic two-dimensional pattern corresponding to each layer of the digital model. By projecting an entire layer simultaneously, the DMD enables rapid and high-resolution photopolymerization of bioinks. Unlike SLA, which scans a focused laser beam point-by-point, DLP employs a DMD-based projection system that cures an entire layer in a single exposure [32–34]. The primary distinction between DLP and other light-based methods, such as Stereolithography (SLA), lies in the light source and curing mechanism. While SLA employs a laser to trace and polymerize the bioink point-by-point, DLP projects an image of an entire cross-sectional layer simultaneously. This capability allows DLP to achieve significantly higher printing speeds regardless of the layer’s complexity, making it highly beneficial for rapid prototyping or producing large volumes of bio-scaffolds (Fig. 1) [30, 31]. This technology enables the fabrication of bio-scaffolds with controlled porosity and geometry, which is crucial for mimicking the natural cellular environment [35]. DLP can create scaffolds with microfluidic channels that facilitate nutrient and oxygen exchange between cells, promoting cell growth and proliferation. Additionally, it can produce multi-material structures, enabling the creation of complex environments that mimic various body tissues. Another significant advantage of DLP in regenerative medicine is its ability to produce engineered tissues with structures and functions similar to natural tissues. These engineered tissues can be used for studying diseases, developing new drugs, and implanting into patients. Furthermore, DLP can be used to create accurate anatomical models of organs and tissues for surgical planning and medical education. With recent advancements in biomaterials and design software, DLP is expected to play a crucial role in personalized medicine and tissue regeneration [35–37].

Fig. 1.

The main components of a DLP bioprinter

Slicing and control software translates the 3D design into printable layers and fine-tunes settings like exposure time and layer thickness. Some advanced DLP printers even include features like material switching systems or recoaters, which are especially useful when printing with multiple materials or thicker bioinks [38].

DLP is a nozzle-free technique, which eliminates issues such as nozzle clogging and shear stress factors that can negatively affect cell viability in extrusion printing. Moreover, it shows great compatibility with various photocurable bioinks and allows for multi-material fabrication, enabling the creation of complex tissue scaffolds with embedded stem cells and diverse extracellular matrix components [39]. Collectively, these features position DLP bioprinting as a leading candidate for the scalable, high-fidelity fabrication of complex, cell-laden tissues. However, despite the hardware advantages of DLP, its widespread clinical application is currently constrained by the material science of the bioinks. The development of bioinks that are simultaneously photocurable, biocompatible, and mechanically similar to native tissue remains a critical bottleneck. While existing literature has broadly covered bioprinting hardware, a comprehensive analysis focusing specifically on the interplay between DLP optical parameters and bioink chemistry is still lacking.

Bioinks for DLP bioprinting

Bioinks are fundamental components of digital light processing (DLP) bioprinting, enabling the fabrication of complex three-dimensional (3D) biological constructs for diverse applications, including tissue engineering and regenerative medicine [15, 31, 40, 41]. Typically, these materials are hydrogels that encapsulate living cells and provide them with a conductive microenvironment for survival, proliferation, and differentiation [42–45]. Photocurability, which allows bioinks to crosslink rapidly upon exposure to light, is a key requirement for DLP bioinks [40, 46, 47]. Smooth flow and rapid layer-by-layer fabrication without excessive shear stress are essential for efficient printing [40, 47, 48]. The printed layers must also be transparent to ensure uniform light penetration and polymerization [40, 47]. Bioinks must exhibit high cytocompatibility to minimize adverse effects on cell viability and function [42, 43]. Rheological properties, such as shear-thinning behavior and yield stress, allow the material to flow under printing pressure while maintaining structural integrity [49]. The combination of these properties defines a ‘printability window’ that balances flow during printing with structural integrity after deposition, ensuring both mechanical stability and biological activity [15].

The development of bioinks for DLP bioprinting has progressed significantly, from rudimentary formulations to sophisticated multi-component systems. Polymers such as poly(ethylene glycol) diacrylate (PEGDA) and modified natural polymers such as gelatin methacryloyl (GelMA) were the focus of early advancements in the mid-2010s [50]. In these initial bioinks, DLP showed its feasibility for creating 3D structures, although UV light for crosslinking can be cytotoxic. A critical shift towards visible light-based photopolymerization systems was observed post-2018, which leveraged photoinitiators compatible with longer wavelengths [51–53]. Biological applications and biomaterials were expanded by milder crosslinking conditions [53]. Research has focused on advanced natural and hybrid bioinks, integrating extracellular matrix (ECM) components and diverse functional groups to better mimic native tissue environments [54, 55].

Several types of biomaterials are used in DLP bioinks, including natural, synthetic, and hybrid types. In general, natural biomaterials exhibit excellent biocompatibility and often contain intrinsic biological cues that promote cell adhesion and differentiation [42, 54]. GelMA is a widely used photocrosslinkable derivative of gelatin with excellent cytocompatibility and tunable mechanical properties [43, 51]. Due to its derivative nature, GelMA lacks the triple-helix molecular structure of collagen, which may limit its ability to fully replicate native collagen’s biological functions. HAMA (Hyaluronic Acid Methacryloyl) is also a natural polymer used to enhance biocompatibility and cellular interaction, often mixed with GelMA to create functional volumetric soft tissues. Bioinks based on silk fibroin, including methacrylated silk (Sil-MA), are gaining attention due to their excellent mechanical properties, biocompatibility, and biodegradability [56, 57]. dECM-derived bioinks, such as dECM-methacryloyl (dECM-MA), mimic the native extracellular matrix composition, providing unique biological cues for cell behavior and promoting tissue regeneration [54]. In vat photopolymerization, dECM-MA bioresins functionalized with glycidyl methacrylate (GMA) from digested porcine liver dECM are promising [55]. Methyl acrylated collagen (CMA)-based bioinks are also being developed for precision DLP bioprinting for applications such as skin wound healing [58].

On the other hand, synthetic biomaterials offer greater control over mechanical properties, degradation rates, and chemical modifications [40]. DLP bioprinting often uses poly(ethylene glycol) diacrylate (PEGDA) because of its low viscosity and tunable crosslinking kinetics, although crosslinking typically occurs via chain-growth photopolymerization, which can result in heterogeneity in crosslinking density [48]. Another synthetic option is poly(vinyl alcohol) methacrylate (PVA-MA). Inks combining natural and synthetic materials aim to provide enhanced biological functionality and mechanical robustness. A composite scaffold utilizing GelMA and keratin methacryloyl (KerMA) has been developed for perforating tympanic membranes [59]. As a result, interpenetrating networks (IPNs) of GelMA and glycidyl methacrylate silk fibroin (SG) have been created to create reinforced bioinks for cartilage tissue engineering, which balances cell survival as well as the density of the crosslinking of the hydrogel. In addition, photocurable chitosan-based hybrid bioinks, often combined with acrylamide, have shown improved mechanical stability as well as increased biocompatibility for tissue engineering applications under favorable conditions [60, 61]. There is ongoing research aimed at creating novel hybrid compositions, such as poly(ethylene glycol)–norbornene (PEG-norbornene) photo-click bioinks, which offer mild and rapid reaction kinetics along with high cytocompatibility, making them an attractive alternative to traditional (meth)acrylated macromers [48]. There is a wide range of bioinks that is capable of allowing the creation of heterogeneous hydrogel constructs varying in mechanical properties and intricate architectures, including perfusable microfluidic chips, which enable high-fidelity tissue models to be created, advancing tissue engineering through tissue engineering [62, 63].

Biomaterials

Natural biomaterial

Natural biomaterials remain central to the development of photocrosslinkable bioinks; however, recent progress has introduced a new generation of natural-derived polymers designed specifically for high-resolution digital light processing (DLP) bioprinting. These materials offer improved photocuring efficiency, enhanced mechanical tunability, and superior biological performance compared with traditional gelatin-, alginate-, or chitosan-based systems [42, 60, 64].

One of the most impactful innovations is the development of methacrylated silk fibroin derivatives, including silk fibroin methacryloyl (Sil-MA) and silk fibroin–glycidyl methacrylate (silk-GMA). These materials exhibit excellent optical clarity, rapid photocrosslinking, and high structural fidelity, making them particularly suitable for DLP projection printing [65]. Kim et al. demonstrated that Sil-MA can be precisely printed into complex 3D architectures while maintaining high cell viability and mechanical robustness [66]. Further advancing this chemistry, Hong et al. used silk-GMA in DLP to fabricate chondrocyte-laden constructs that supported long-term viability, enhanced chondrogenic differentiation, and promoted cartilage regeneration in vivo [67]. Another important category includes GelMA-based composite bioinks, which remain the most widely adopted natural biomaterial in DLP bioprinting. Recent formulations incorporating nano-reinforcements (e.g., silica nanoparticles, nanoclay, or bioactive ceramics) have significantly improved photocuring depth, stiffness, and shape fidelity while retaining the inherent bioactivity of gelatin methacryloyl. Such composites continue to serve as benchmark natural-derived inks for DLP-based tissue engineering [68, 69]. Additionally, methacrylated hyaluronic acid (HAMA) has emerged as a versatile natural polymer for soft tissue–mimetic constructs. Its high water content, tunable viscosity, and compatibility with visible-light projection printing make HAMA an attractive option for fabricating cell-friendly hydrogel microarchitectures. Recent work has demonstrated its successful use in DLP systems for engineering muscle, cartilage, and neural tissues [70].

Another notable advancement is the development of plasma-derived methacryloyl hydrogels (Plasma-MA). Wu et al. synthesized Plasma-MA by methacrylating fresh frozen plasma and demonstrated that it can be processed using DLP to create customized constructs that retain endogenous growth factors and support favorable cell behavior [3]. Broader analyses of blood-derived biomaterials further reinforce the relevance of plasma-based inks for producing naturally bioactive, tissue-specific constructs [71]. Finally, the development of methacrylated decellularized extracellular matrix (dECM-MA) bioinks represents one of the most significant advances in biologically instructive DLP-compatible materials. These photocurable formulations retain a complex array of native ECM proteins and signaling cues while exhibiting improved print resolution and curing kinetics after methacrylation. dECM-MA has been successfully used for DLP printing of tissue-specific constructs, including cardiac, cartilage, liver, and adipose models [72]. These emerging natural biomaterials (including Sil-MA, silk-GMA, GelMA composites, HAMA, Plasma-MA, and dECM-MA) expand the capabilities of DLP bioprinting. They enable the fabrication of cell-responsive, structurally stable, and biologically functional constructs with micron-scale precision.

The cell-adaptive hydrogel represents a class of natural biomaterials that are specifically engineered for Digital Light Processing (DLP) bioprinting, utilizing their ability to reversibly reorganize structurally after being activated by the cells [73]. Using dynamic covalent bonds, such as hydrazones, boronate ester bonds, disulfide bonds, and imine bonds, or supramolecular interactions, these materials form stress-relaxing networks that respond to cellular traction forces [74, 75]. Due to this inherent dynamic nature, post-printing matrix remodeling can facilitate essential cellular processes such as migration, proliferation, and the deposition of native extracellular matrix (ECM) components [48, 73]. This adaptability enhances the spatiotemporal fidelity of DLP bioprinting. As a result, it enables better mimicry of native ECM dynamics and improves long-term cell viability and functional maturation compared with static hydrogels [76]. DLP-bioprinted constructs that are based on imine networks of chondroitin sulfate have been used for cartilage regeneration, vascular networks have been formed within hyaluronan-gelatin hydrogels that have been crosslinked with boronate, and soft tissue engineering has been enhanced using alginate–tyramine hydrogels with disulfide-mediated stress relaxation [77].

Synthetic polymers

In addition to natural polymers, synthetic biomaterials play a pivotal role in the development of DLP-compatible bioinks due to their highly tunable physicochemical properties, reproducibility, and structural stability [78, 79]. Unlike natural polymers, which often suffer from batch-to-batch variability and limited mechanical strength, synthetic polymers can be precisely engineered to achieve desired viscosity, degradation rates, and mechanical performance while maintaining optical clarity required for efficient light penetration in DLP systems [80].

Poly(ethylene glycol) diacrylate (PEGDA) and its derivatives represent one of the most widely utilized synthetic backbones for photopolymerizable hydrogels. PEGDA-based bioinks are hydrophilic, biocompatible, and mechanically stable, and their crosslinking density can be easily modulated by controlling molecular weight and degree of acrylation [81]. While PEGDA hydrogels inherently lack bioactive motifs necessary for cell adhesion and signaling, they can be functionalized with short peptides such as RGD or blended with natural polymers like gelatin or hyaluronic acid to enhance cellular interactions. Such hybrid systems have been successfully applied in vascularized tissue constructs, cartilage regeneration, and stem cell-laden scaffolds [82].

Poly (vinyl alcohol) (PVA) derivatives, including methacrylated PVA, have also demonstrated promise in DLP printing. Their high transparency and water solubility support precise photopolymerization, while modifications enable improved mechanical resilience and degradability [83, 84]. Similarly, poly(caprolactone) (PCL), a semicrystalline aliphatic polyester, has been incorporated into DLP systems to provide slow degradation and superior mechanical strength, particularly in load-bearing applications such as bone or dental tissue engineering. However, the hydrophobicity of PCL often necessitates blending with hydrophilic polymers to ensure favorable cell encapsulation and nutrient diffusion [85, 86].

More recently, synthetic polymers with dynamic or stimuli-responsive chemistries have been investigated to impart advanced functionalities to DLP bioinks. For example, polymers with reversible crosslinking bonds or shape-memory behavior allow for adaptive scaffolds that can better mimic the dynamic remodeling of native tissues. Biodegradable polyurethanes and poly(propylene fumarate) (PPF) have also attracted attention, offering tunable degradation profiles and mechanical properties tailored for musculoskeletal regeneration [87].

A concise comparison between natural and synthetic bioinks in hydrogel-based bioprinting, including DLP systems, highlights their complementary strengths and limitations. Natural-derived hydrogels, such as gelatin, hyaluronic acid, and silk-based matrices, provide excellent cytocompatibility and tissue-mimetic biochemical cues but often suffer from lower mechanical strength, slower or less controllable gelation, and batch-to-batch variability. In contrast, synthetic bioinks, particularly PEG-based and other engineered polymer networks, offer highly tunable and reproducible mechanical properties, rapid and uniform crosslinking, and superior shape fidelity, yet lack intrinsic bioactivity and typically require functionalization or blending with natural components to support cell adhesion and remodeling. Recent reviews on bioink design and hydrogel printability emphasize that many contemporary formulations adopt hybrid strategies. Combining natural and synthetic polymers balances printability with biological performance [88, 89].

Mechanisms of photopolymerization

Polymerization is the chemical process where monomer units link together to form large polymer chains or networks [90]. During the Digital Light Processing (DLP) [bioprinting process, light energy is specifically used to induce these reactions, enabling the rapid solidification of liquid bioinks into intricate, cell-laden three-dimensional structures. DLP bioprinting relies heavily on light-mediated crosslinking to achieve its high resolution and precision, allowing layer-by-layer fabrication of biomimetic scaffolds [27, 31].

DLP bioprinting primarily uses initiator-based and step-growth photopolymerization mechanisms. Photoinitiators absorb photons and produce reactive species [91]. When exposed to light, Type I photoinitiators undergo unimolecular cleavage, forming free radicals that initiate chain-growth polymerization [92]. It is possible to achieve rapid curing kinetics thanks to this mechanism, which contributes to faster fabrication and better structural integrity in high-speed DLP. However, free radicals are susceptible to oxygen inhibition, leading to uncured surfaces and heterogeneous crosslinking, while byproducts may compromise cytocompatibility [41]. The type II photoinitiators operate by bimolecular hydrogen abstraction, often utilizing visible light, which is more cytocompatible than UV light [93]. Although Type II systems are potentially slower, they offer better control over radical generation and improved cell viability, balancing the curing process with reduced cellular damage [94].

It has gained prominence for its biofabrication benefits through step-growth photopolymerization, in particular thiol-ene click reactions. In thiol-ene reactions, thiols are efficiently added to carbon-carbon double bonds, resulting in highly uniform networks [95, 96]. As a result of these reactions, the gelation kinetics and material properties can be precisely controlled and are less susceptible to oxygen inhibition. As a result of the mild reaction conditions and homogeneous crosslinking, thiol-ene systems make excellent cell-laden bioinks, promoting high cytocompatibility as well as mechanical stability when printed [97].

Photopolymerization is a process where light, particularly ultraviolet (UV) or visible light, is used to initiate chemical reactions, especially polymerization [98]. In Digital Light Processing (DLP) 3D bioprinting, this process plays a pivotal role in constructing three-dimensional layers. The light activates chemical initiators within the bioink, triggering the formation of chemical bonds between the bioink molecules, leading to solidification [99]. Understanding these photopolymerization and gelation mechanisms is essential for optimizing bioink formulations and printing parameters [48]. Generally, there are two primary types of photopolymerization mechanisms in DLP printers: initiator-based photopolymerization and initiator-free photopolymerization [100].

Initiator-based photopolymerization is a common method in DLP 3D bioprinting [48]. In this method, a light-sensitive material (bioink) contains chemical initiators. When exposed to light, these initiators generate free radicals. These radicals interact with the bioink monomers, initiating polymer chains. As this reaction progresses, the polymer chains grow and eventually bond together, forming a three-dimensional network [48, 101]. Advantages of this method include precise control over polymerization rate and the ability to use a wide range of bioinks. However, there is a possibility of residual initiators in the final product, which may affect its properties. Chemical initiators play a crucial role in this process. [102, 103].

In contrast to initiator-based photopolymerization, initiator-free photopolymerization is a process where light directly interacts with bioink monomers, breaking chemical bonds and generating free radicals. This approach is limited to monomers with inherent photosensitive moieties and is less commonly used in standard DLP systems due to higher energy requirements and lower efficiency and eliminates the need for added chemical initiators [104, 105]. Ultraviolet (UV) or visible light with sufficient energy strikes the chemical bonds in the bioink monomers, causing them to break and form free radicals. These radicals then react with each other or with other monomers, initiating polymer chains [106]. The specific mechanism depends on the monomer’s chemical structure. Certain monomers, due to their unique structures, are more susceptible to light absorption and free radical generation. This method is favored for its simplicity and the higher purity of the final product [105]. Additionally, it is a simpler process compared to initiator-based methods. However, this approach also has limitations [107]. Breaking chemical bonds in monomers requires higher light energy compared to other methods. Furthermore, only specific monomers can absorb light and generate free radicals [106]. As shown in Fig. 2, the fundamental differences between initiator-based and initiator-free photopolymerization systems are illustrated, highlighting how radical species are generated through photoinitiators or via auto-oxidation and chromophore-activated pathways, respectively.

Fig. 2.

Comparative schematic of initiator-based and initiator-free photopolymerization mechanisms. Initiator-based systems form radicals via photoinitiators, while initiator-free systems generate radicals through auto-oxidation or light-activated chromophores, leading to polymerization

In addition to photopolymerization mechanisms, gelation mechanisms play a crucial role in DLP printing. Gelation is a process where a liquid substance transforms into a solid or gel-like material. In DLP printers, gelation refers to the conversion of liquid bioink into a solid, three-dimensional structure. This process is directly linked to photopolymerization mechanisms. The quality of gelation significantly impacts the mechanical, chemical, and visual properties of the final part. If gelation is not carried out correctly, the final part may exhibit weaknesses, cracks, or deformation [48, 108]. Therefore, a deep understanding of crosslinking mechanisms is essential for optimizing bioink formulations, printing parameters, and achieving uniform network formation and high printing fidelity in DLP bioprinting. crosslinking mechanisms can be categorized based on the type of chemical reaction leading to the formation of a three-dimensional network:

• Free Radical Chain Polymerization: Free radical chain polymerization begins when heat or light decomposes an initiator into reactive radicals. These radicals attack monomers to start chain formation, which rapidly propagates through sequential monomer addition. Termination occurs when two growing chains combine or a radical is quenched, producing the final polymer. However, this mechanism is susceptible to oxygen inhibition and may lead to non-uniform crosslinking in thicker constructs, limiting its use in large-scale DLP-printed tissues (Fig. 3i) [109, 110].

• Thiol-Azide Click Reaction: The thiol-azide click reaction is a highly efficient and selective process involving the rapid formation of thioether (S-N) bonds between thiol (SH) and azide (N₃) functional groups. Due to its speed, selectivity, and compatibility with aqueous environments, this reaction is widely employed in click chemistry. In the context of DLP 3D bioprinting, bioinks containing thiol and azide groups can undergo click chemistry under UV or visible light irradiation, forming a three-dimensional network. This network solidifies the liquid bioink, allowing printed layers to adhere and create the final structure. A significant advantage of this method is the ability to design and synthesize bioinks with custom properties. Moreover, the reaction’s compatibility with aqueous environments makes it suitable for bioprinting applications. Owing to its speed and selectivity, this reaction is extensively used in the production of hydrogels. This step-growth approach offers excellent oxygen tolerance, uniform network formation, and superior cytocompatibility, making it highly suitable for cell-laden DLP bioinks (Fig. 3ii) [48, 111–113].

• Photoinduced Radical Polymerization: In this reaction, ultraviolet or visible light excites molecules containing phenolic groups. In photoinduced radical polymerization involving phenolic groups, radical generation typically occurs via photoinitiators or mediated oxidation pathways rather than direct excitation of the phenolic moieties alone. The generated free radicals then react with each other, forming new covalent bonds. These new bonds create a three-dimensional network, causing the liquid bioink to solidify. This mechanism is favored due to its simplicity and the possibility of using natural materials like phenolic bioinks. However, the reaction rate is slower compared to some other mechanisms, and it may require longer exposure times. Beyond the factors already mentioned, various elements such as bioink type, light intensity and wavelength, temperature, and initiator concentration influence the gelation process. While this mechanism enables milder conditions and compatibility with natural phenolic bioinks, it often exhibits slower reaction rates and remains partially dependent on initiator efficiency or light penetration (Fig. 3iii) [114–116].

Fig. 3.

Schematic illustration of three gelation mechanisms based on the type of chemical reaction that leads to the formation of a three-dimensional network: free radical chain polymerization, thiol-azide click reaction, Photoinduced radical polymerization

In summary, free radical polymerization offers fast curing and strong network formation but suffers from oxygen inhibition and non-uniform crosslinking in thicker constructs. Click-based step-growth reactions provide uniform, oxygen-tolerant crosslinking with good cytocompatibility, though they often yield lower stiffness and require precise functional group balance. Visible-light radical mechanisms support high resolution with reduced phototoxicity but depend on initiator efficiency and light penetration. Each mechanism presents trade-offs between curing speed, mechanical performance, crosslinking uniformity, and cell compatibility, making careful selection essential for DLP bioprinting.

Visible light initiators

Visible light initiators serve as photocatalysts in DLP 3D bioprinting processes. Upon absorbing visible light photons, these materials become excited and subsequently decompose into free radicals. The generated free radicals rapidly react with monomers present in the bioink, initiating polymer chains. As a result of this reaction, the liquid bioink gradually solidifies, forming the desired 3D structure [117]. The selection of initiator type and quantity depends on various factors including the type of bioink, the wavelength of the light source, the required polymerization rate, and the mechanical and chemical properties of the final part [115, 118]. Based on the photoinitiation process, they can be divided into two types, type I and type II photoinitiators. In type I photoinitiators, after light absorption, homolytic cleavage occurs and free radicals are produced, while in type II photoinitiators, non-cleavable molecules abstract hydrogen from hydrogen donors as co-initiators, generating free radicals [119]. Common commercial photoinitiators such as 2,4,6-trimethylbenzoyl diphenyl phosphine oxide (TPO) and phenyl bis(2,4,6-trimethylbenzoyl) phosphine oxide (BAPO) are usually type I [120]. The performance of visible light initiators is closely related to their chemical structure. These materials typically contain functional groups that enable them to absorb visible light and generate free radicals. Some of the most important types of photoinitiators include aromatic ketones, azo compounds, and triazinium salts. Each of these compounds has unique properties that make them suitable for specific applications [121, 122].

Photoinitiators play a pivotal role in DLP 3D bioprinting technology, especially in the field of bioprinting. These chemical substances, upon absorbing visible light energy, directly convert into free radicals. The generated free radicals, being highly reactive species, initiate covalent bonds with monomers present in the bioink, rapidly leading to polymerization [117]. The growth of polymer chains continues until two free radicals combine and terminate the chain. This process results in the localized solidification of the bioink in the irradiated areas, consequently producing thin and precise layers of solid material [123].

Aromatic ketones are one of the most important classes of photoinitiators used in DLP bioprinting processes. Due to their relative biocompatibility and high efficiency, they are considered a suitable option for creating complex and precise biological structures [35]. By appropriately selecting the aromatic ketone compound and adjusting the process parameters, structures with desired mechanical, chemical, and biological properties can be produced. When visible light is irradiated onto an aromatic ketone, the pi electrons in the benzene ring are excited to a higher energy state. This excited state is unstable and leads to the cleavage of the carbon-oxygen bond in the carbonyl group. As a result of this cleavage, two free radicals are formed: an aryl radical and an alkoxy radical. Due to its relative stability, the aryl radical can act as an efficient initiator in polymerization reactions [124, 125]. Some of the most common aromatic ketones used as initiators in DLP bioprinting include benzophenone, acetophenone, and anthraquinone [126, 127].

Azo compounds represent another significant class of photoinitiators employed in bioprinting processes. When visible light is irradiated onto an azo compound, the nitrogen-nitrogen (N = N) bond in the azo group cleaves, resulting in the formation of two nitrogen radicals. These radicals are highly reactive and quickly react with monomers in the bioink, initiating polymer chains [109, 128, 129]. This process leads to the localized solidification of the bioink in the irradiated areas, consequently producing thin and precise layers of solid material.

Triazinium salts, as a novel class of photoinitiators, have garnered significant attention in recent years. Upon exposure to visible light, the nitrogen-nitrogen (N = N) bond within the triazinium group cleaves, generating two nitrogen radicals. These radicals are highly reactive and promptly initiate polymerization reactions with bioink monomers. This process results in the localized solidification of the bioink in the irradiated regions, consequently producing thin and precise layers of solid material [104, 130, 131]. Compared to many other initiators, triazinium salts exhibit higher sensitivity to visible light, enabling the use of lower intensity light sources. Additionally, these materials induce high polymerization rates, reducing the time required to construct 3D structures. Triazinium salts are compatible with a broad spectrum of bioinks and can be utilized to fabricate various biological structures. Many triazinium salts possess lower toxicity compared to other initiators, making them more suitable for biological applications [121, 132].

Table 2 provides an overview comparison of each UV and Visible Light Photoinitiator Groups that can help in their selection for bioink applications, and Table 3 presents the most important photoinitiators used in inks used in DLP 3D bioprinting.

Table 2.

Comparison of common UV and visible light photoinitiator groups used in DLP 3D bioprinting. Cytocompatibility ratings are derived from standard cell viability assays reported in the cited references (primarily MTT, Alamar Blue, PrestoBlue, or Live/Dead fluorescence staining) performed on mammalian cells (e.g., fibroblasts, mesenchymal stem cells, endothelial cells, or chondrocytes) after exposure to the photoinitiator at relevant concentrations and light doses (typically 405 nm, 10–60 s); “good to very good” indicates cell viability > 85–95%, “moderate” to 60–85%, and “low to moderate” to <60% under typical DLP printing conditions.Toxicity ratings are based on the same viability assays and additional reports of cytotoxicity, genotoxicity, or in vivo biocompatibility; “low” corresponds to minimal adverse effects, “moderate” to acceptable levels with some concerns, and “high” to significant risks.Cost estimates (“low”, “Medium”, “high”) are qualitative assessments derived from commercial availability, synthesis complexity, and market pricing in research and industrial contexts as reported in the literature [130, 133–136]

Type of Photoinitiator	Common Examples	Optimal Wavelength (nm)	Cytocompatibility	Suitable Solvents	Decomposition Temperature (°C)	Polymerization Rate	Specific Applications	Toxicity	Cost	Advantages	Disadvantages	Chemical Comparison	Ref
Aromatic Ketones	Benzophenone, Darocure 1173, Anthraquinone, Benzil, Michler’s ketone	300–400 (UV-A)	Moderate to good	Alcohols, esters, ketones, aromatics	100–150	Medium to high	Coatings, adhesives, 3D bioprinting	Low to moderate	Medium	High sensitivity to light, high efficiency, structural diversity	Limited UV absorption, may cause yellowing	Simple structure, UV-Vis absorption	[35, 124–127]
Triazene Salts	Aromatic triazene salts, Aliphatic triazene salts, Triazene salts complexed with metals	380–500 (Visible)	Good to very good	Water, alcohols, polar organic solvents	100–200	Very high	3D bioprinting, coatings, adhesives, bioapplications	Low	High	High sensitivity to visible light, high polymerization rate, compatibility with various resins	High cost, sensitive to moisture	Complex structure, visible light absorption	[104, 121, 130–132]

Table 3.

Common photoinitiators used in DLP bioprinting

Photoinitiator	Maximum Absorption Wavelength (nm)	Primary Applications	Advantages	Disadvantages	Ref
Eosin Y	520–530	Bioprinting, light-sensitive cells, tissue engineering	High biocompatibility, sensitivity to green light	Oxygen sensitivity, moderate polymerization rate	[137, 138]
Ruthenium Bipyridine (Ru(bpy)₃Cl₂)	450–460	Visible light 3D bioprinting, cell scaffold fabrication	High quantum yield, tunable absorption wavelength	High cost, sensitivity to UV light	[139, 140]
Camphorquinone/Tertiary Amines	460–480	3D bioprinting of commercial resins, hydrogel production	Low cost, commercially available	Oxygen sensitivity, moderate polymerization rate	[141, 142]
BODIPY Derivatives	Tunable	Green and red light 3D bioprinting, fast polymerization	High polymerization rate, tunable absorption wavelength	High cost, complex synthesis	[143–146]
Azo Compounds	350–450	3D bioprinting of commercial resins, thin film production	Sensitivity to UV and visible light, low cost	Low thermal stability, toxicity of some compounds	[147–149]
Irgacure 784	470–550 nm	Fast polymerization, suitable for 3D bioprinting with commercial resins	High quantum yield, good thermal stability	Oxygen sensitivity	[133, 150]
Irgacure 819	385	Deep curing, suitable for thick parts	High penetration depth, good thermal stability	Oxygen sensitivity	[151]
TPO (Trimethylbenzoyl diphenyl phosphine oxide)	365–405	Fast polymerization, suitable for 3D bioprinting	High quantum yield, good thermal stability	Oxygen sensitivity	[141, 152]
BAPO (Bis phosphine oxide)	385–420	Deep curing, suitable for thick parts	High penetration depth, good thermal stability	Oxygen sensitivity	[153]
Benzophenone Acrylates	320–380	Fast polymerization, suitable for coatings and adhesives	High quantum yield, good thermal stability	Oxygen sensitivity	[154, 155]
Fluorescein	510–520	Bioprinting, light-sensitive cells	High biocompatibility, sensitivity to green light	Oxygen sensitivity, moderate polymerization rate	[156]
Rose Bengal	548	Bioprinting, thin film production	Sensitivity to green light, high quantum yield	Toxicity, light sensitivity	[157–159]
Lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP)	365–405	Visible light 3D bioprinting, cell scaffold fabrication, dental materials	High reactivity, good solubility in polar solvents	Sensitivity to moisture, potential toxicity	[134–136, 160]
Camphorquinone (CQ)	460–480	3D bioprinting, dental composites, adhesives	Low cost, widely available, good compatibility with amines	Oxygen sensitivity, moderate reactivity	[141, 161, 162]

Application of DLP bioprinting on regenerative medicine

To accurately mimic native tissue, the bioprinted tissue model requires various characteristics such as tunable stiffness (from rigid porous scaffolds for bone load-bearing to soft hydrogels for cartilage, liver, muscle, vessels, etc.), structural resolution, shape stability, and support for the physiological microenvironment [163, 164]. However, the extrusion-based bioprinted model exhibits poor printing resolution (typically less than 200 μm) as well as permanent deformation due to the change in crosslinking rate during printing. On the other hand, bioprinting technology has recently been used to integrate multifunctional rigid and soft biomaterials in the field of tissue engineering, which enables the printing of tissue models with complex structures and functions. In this regard, the extrusion-based bioprinting strategy faces many limitations that make its application challenging [165, 166]. To address these shortcomings, the DLP-based bioprinting strategy has been used as a novel, effective, and practical approach by providing the infrastructure to create complex tissue scaffolds with the ability to encapsulate various types of (stem) cells. In addition, the stiffness of the printed model can be tuned by adjusting the duration of light irradiation, and scaffolds with the desired stiffness can be created [56, 167, 168].

Wound regeneration

As the largest organ of the body, the skin is highly susceptible to injuries such as burns, chronic skin diseases, lacerations, and others. Today, skin tissue engineering has emerged as a revolutionary new strategy for skin regeneration [169]. The development of biomimetic skin substitutes requires multilayered tissue architecture, precise control over cellular organization, and angiogenesis. DLP-based bioprinting creates functional epidermal and dermal structures through layer-by-layer deposition of endothelial cells, fibroblasts, keratinocytes, and melanocytes in hydrogels. Recent studies have demonstrated the potential of DLP-printed skin structures in accelerating wound healing, reducing scarring, and providing topical treatments for diseases such as diabetic ulcers [170–172]. In one study, a GelMA/chitosan oligosaccharide (COS)-based bioink was developed via the DLP technique for printing skin substitutes. The GelMA/COS bioink showed good mechanical properties, swelling behavior, biodegradability, and porosity. Furthermore, the addition of COS enhanced the antibacterial effects of the bioink and also significantly reduced the expression of fibrosis-related genes such as fibronectin I, collagen I, and collagen III. After culturing human dermal fibroblasts (HDFs), the GelMA/COS skin scaffold facilitated the expansion and proliferation of HDFs, exhibited high cell viability, and showed excellent shape fidelity. Moreover, in vivo studies showed that the skin scaffold reduced inflammation, accelerated wound closure, and promoted wound healing (Fig. 4i) [173]. In another study, DLP-printed skin hydrogels of methacrylated silk (Silk-GMA) and gelatin (Gel-GMA) mixed with cells from the epidermal and dermal layers (fibroblasts, keratinocytes, and vascular endothelial cells) were developed for skin tissue engineering. Histological analysis and immunofluorescence staining showed promotion of fibroblast and keratinocyte proliferation as well as increased expression of CD31, phalloidin, and cytokeratin 13. Additionally, in vivo studies in full-thickness wound models demonstrated improved wound healing in the dermis and epidermis layers. DLP-printed skin hydrogels exhibited good mechanical stability and biocompatibility over 4 weeks, confirming their potential in skin tissue engineering (Fig. 4ii) [174].

Fig. 4.

DLP 3D printing for skin modeling. i) HDFs loaded dermal model with GelMa.COS bioink for skin defect model. ii) Skin multilayer model with GelGMA.SGMA bioink

Recently, Xiao et al. designed dimethyl acrylated collagen/dimethylphenylphosphonate (CMA/DPPA)-based bioinks for skin tissue engineering via DLP technology. This strategy significantly enhanced the mechanical strength, resistance to enzymatic degradation, and printing accuracy of 3D structures. The CMA-DPPA bioink also significantly improved cell adhesion, proliferation, and migration, and showed excellent biocompatibility. Furthermore, the results of in vivo studies in a rat model showed that the CMA-DPPA bioink increased collagen deposition, facilitated wound closure, and improved epidermal regeneration. These findings confirm that the CMA-DPPA bioink has promising potential in the future of skin tissue engineering [58]. In another study, DLP-printed bioinks based on CMA and dihydromyricetin (DHM) were developed to produce biomimetic skin substitutes for diabetic wound regeneration. The results showed that the CMA-DHM bioink offered high micro-scale printing accuracy, low swelling, high cell viability, and structural stability (231% increase in mechanical strength, 161% increase in stiffness, and 208% increase in biodegradability). The bioink was also shown to increase HDF cell proliferation, modulate macrophage polarization, reduce oxidative stress, and exhibit high anti-inflammatory and antioxidant properties. Furthermore, the CMA-DHM bioink enhanced angiogenesis, facilitated collagen deposition and re-epithelialization, prevented wound formation, and accelerated effective tissue regeneration. These findings confirm that the CMA-DHM bioink is a promising scalable system for the management and treatment of diabetic wounds [91]. DLP-printed skin constructs have shown enhanced fibroblast and keratinocyte activity, reduced inflammation, and improved angiogenesis, supporting faster and more organized wound healing. Despite these promising outcomes, challenges remain in achieving long-term mechanical stability, reproducing the complex barrier functions of skin, and ensuring adequate hydration and elasticity under dynamic wound conditions. Future research should focus on developing bioinks with improved antibacterial and antioxidant properties, enhanced stromal–epidermal integration, and tailored mechanical behavior to better replicate the layered structure and regenerative capacity of native skin.

Cartilage and bone regeneration

Effective bone and cartilage regeneration requires addressing tissue engineering, cell biology, biomechanics, angiogenesis, and clinical translation, and the limited regenerative capacity of chondrocytes is only part of a multifactorial challenge, while it is a key factor for articular cartilage repair [175]. A suitable solution must integrate with the host tissue, restore structure and function, and be feasible and safe in the clinical phase. To address this issue, DLP-based 3D bioprinting technology has been investigated as an effective approach for fabricating tissue-engineered scaffolds that mimic host microenvironmental conditions [176]. In one study, biocompatible hydrogels based on methacrylated gelatin/polymethacrylic acid (GelMA/PMAA) were prepared for bone tissue engineering by DLP technology. The presence of PAMA in the hydrogel significantly increased viscoelasticity, and the use of DLP technology also showed significant effectiveness in iron ion adsorption, which led to the activation of the HIF-1α signaling pathway, facilitated angiogenesis, promoted chondrogenic differentiation, and improved bone regeneration. Therefore, DLP technology has significant implications for endochondral bone formation due to its iron ion adsorption property (Fig. 5i) [177]. In another study, GelMA/dextran-based hydrogels mixed with bone marrow stem cells (BMSCs) were designed for cranial bone regeneration via DLP technology. The results showed that these hydrogels could enhance the spreading, proliferation, and migration of BMSCs, and also improve the osteogenic differentiation of BMSCs by activating the YAP signaling pathway. Furthermore, the results of in vivo studies confirmed that the bioprinted hydrogels significantly promoted osteogenesis. These findings indicate that these hydrogels are suitable candidates for bone regeneration (Fig. 5ii) [178]. DLP bioprinting has enabled the fabrication of patient-specific cellular scaffolds for treating bone defects and disorders. Kumari et al. designed 3D methacrylate-κ-carrageenan composite hydrogels with bioactive silica nanoparticles (MA-κ-CA/BSNPs) via DLP technology. The 3D printing of complex bone structures, such as gyroids, was confirmed with high resolution and precision. In vitro studies of the composite encapsulated with MC3T3–E1 pre-osteoblast cells showed no cytotoxicity, high cell viability and proliferation over 21 days, and significant osteogenic differentiation. Furthermore, no inflammation was observed after a 21-day in vivo study in Wistar rats, indicating the potential of these composites for bone tissue engineering [179]. Hong et al. developed glycidyl methacrylate-modified silk fibroin (Silk-GMA) bioinks using DLP for cartilage tissue engineering. The results showed that the DLP system produced a 3D product with a uniform cell distribution, resulting from photopolymerization and a fast printing rate. In vitro cell culture demonstrated the viability, proliferation, and differentiation of cells encapsulated in Silk-GMA for up to 4 weeks. Additionally, in vivo experiments on rabbit tracheal defects confirmed the growth of epithelium and cartilage tissue around Silk-GMA. These findings confirm the promising potential of Silk-GMA bioinks in cartilage tissue engineering (Fig. 5iii) [67]. In another study, bioinks based on cartilage microtissues (composed of microtia chondrocytes with a cartilage cell-free ECM) and GelMA were prepared using DLP technology for treating microtia. The results showed that the presence of microtissues increases cell-cell connections and simulates the microenvironment of cell growth. On the other hand, the use of DLP technology resulted in the production of 3D earlobe structures with high printing accuracy, favorable elasticity, low swelling ratio, and high cell viability. In vitro and in vivo studies also showed that microtissue bioinks improved cell growth and proliferation, significantly increased ECM deposition, and showed mature cartilage regeneration after 12 weeks of implantation in mice. These results confirm the promising potential of microtissue bioinks for the effective treatment of microtia [180]. Rajput et al. developed biocompatible and biodegradable bioinks based on methacrylated silk fibroin (SF-MA) for bone tissue engineering. SF-MA bioinks printed by DLP with compressive modulus of 12–96 kPa mimicked the viscoelastic behavior similar to bone tissue. These bioinks also showed a favorable degradation profile (48–91% in 21 days). The use of the DLP technique enabled the printing of complex structures with high precision. Moreover, bioinks encapsulated with pre-osteoblast cells showed high cell viability and effectively supported cell proliferation, cytoskeleton organization, and osteogenesis stimulation. These findings demonstrate the potential of SF-MA bioinks for bone tissue engineering [181].

Fig. 5.

DLP 3D printing for cartilage and bone regeneration. i) GelMA-PMAA scaffold for bone regeneration. ii) Regenerate cranial bone defect with BMSCs-GelMA-Dextran void forming hydrogel. iii) Silk-GMA 3D printed scaffold for rabbit tracheal defect

In another study, digital light processing (DLP) bioprinting was employed to fabricate bone constructs using GelMA bioinks functionalized with nanomaterials [182]. The aim was to directly compare the impact of calcium phosphate (CaP) nanoparticles and graphene oxide (GO) on the printing process, mechanical performance, and biological activity of human mesenchymal stem cells (hMSCs). Results showed that GelMA alone supported high cell viability and proliferation, outperforming some commercial bioinks. The incorporation of CaP nanoparticles (~150 nm) provided a strong osteoinductive effect, evidenced by enhanced osteogenic gene expression and calcium deposition, thus promoting hMSC differentiation toward the osteogenic lineage. In contrast, GO-modified GelMA significantly improved the mechanical stiffness of the constructs, increasing the Young’s modulus, which influenced cell morphology but did not compromise cell viability. Importantly, neither CaP nor GO interfered with the DLP printing process, although minor modifications in microstructural organization were observed. These findings demonstrate that GelMA-based bioinks can be effectively tailored by nanomaterial incorporation, with CaP enhancing biological differentiation and GO providing mechanical reinforcement. Such approaches may help overcome key challenges in regenerating critical-sized bone defects, where both structural stability and early osteogenic signaling are essential [182]. DLP-printed bioinks consistently support osteogenic and chondrogenic activity, demonstrating high cell viability, enhanced ECM deposition, and promising in vivo regeneration of both bone and cartilage defects. Despite these advances, several barriers still hinder full functional repair. Matching the mechanical strength of load-bearing tissues, ensuring long-term scaffold stability, and achieving seamless integration with native skeletal tissue remain key challenges. Future progress will rely on developing stronger hybrid bioinks, incorporating bioactive nanomaterials, and refining DLP curing strategies to better reproduce the hierarchical architecture and mechanical gradients of native bone and cartilage.

Vascular tissue constructs

The proper physiology and function of human organs require an efficient vascular system to facilitate nutrient transport, oxygen supply, and removal of waste products. Various strategies have been explored to fabricate vascular systems that can mimic the specific, complex, and unique structure and function of blood vessels. Tube molding, sheet rolling, and direct scaffolding of biomaterials are common techniques used to create vascular networks. However, these techniques face limitations such as fabricating only macroscale blood vessels and lack of control over the structure of the construct [183, 184]. With advances in materials engineering and microstructure fabrication, control over the cellular microenvironment and cellular aspects such as patterns, alignment, and phenotypes has improved. Meanwhile, 3D bioprinting has become a promising approach for printing vascular constructs due to its scalability, reproducibility, and control over vascularization [185]. In one study, gyroid-shaped tissue constructs based on polyethylene glycol (PEG)/GelMA, mixed with human hepatocytes (HepG2), were produced via DLP technology and then covered with interconnected vascular channels formed by human endothelial cells (ECs). The constructs demonstrated good structural stability over 30 days and maintained cell viability of more than 85%. Functional studies showed that hepatocytes produced significant levels of bilirubin and albumin, mimicking the function of native liver tissue. The ECs’ coating also significantly covered the vascular channels, while viable hepatocyte cells were accessible within the construct. These results confirm the significant progress in the development of vascular liver tissue models [186]. In another study, GelMA-based bone constructs modified with graphene oxide (GO) and calcium phosphate (CaP) were developed by the DLP technique. The results showed that GelMA biomaterials exhibited significantly higher cell viability and proliferation compared to other commercial biomaterials. On the other hand, the addition of CaP enhanced the differentiation of human mesenchymal stem cells (hMSCs) by increasing the expression of osteogenic genes and calcium deposition, as well as increasing vascularization. In contrast, the presence of GO led to an increase in Young’s modulus and also affected cell morphology [182]. Chen et al. printed thick high cell density (HCD) tissues with vascular networks using DLP technology. Using DLP, they achieved a resolution and precision of fifty micrometers with a cell density of 0.1 billion/ml. The printed tissues also showed improved angiogenesis and endothelialization after 14 days, confirming the efficient potential of bioprinting for vascular tissue/organ production [187]. Zhou et al. designed small, tunable, small-diameter blood vessels based on GelMA/PEG diacrylate/alginate/lyase loaded with vascular smooth muscle cells (VSMCs) as the vessel wall and vascular endothelial cells (VECs) in the lumen. The results showed that this design is favorable for nutrient and oxygen exchange while simultaneously supporting continuous cell growth and proliferation. Furthermore, fluid dynamic simulations confirmed that the novel blood vessels presented excellent mechanical and perfusable properties. These findings indicated that printed blood vessels could be a promising candidate as a replacement for small-diameter blood vessels [188].

In another study, DLP bioprinting enabled the encapsulation of endothelial progenitor cells (EPCs) in photocurable hydrogels, where the cells rapidly self-organized into branched, capillary-like tubes with diameters of 10–100 μm [189]. These lumenized structures, stabilized by tight junctions, remained viable for several weeks and could even be perfused with fluorescent dyes, highlighting their potential as functional vascular units. Unlike the conventional endothelial tube formation assay, which typically collapses within 24 h, this method produced long-lasting and predefined microvascular networks. Optimization of the bioink composition combining gelatin, collagen I, hyaluronic acid, and fibrin was critical for providing the mechanical stiffness and biochemical cues required for EPC organization across multiple donors. Furthermore, the constructs were recoverable in alginate patches, enabling their physical manipulation and integration into stacked multicellular tissues [189].

In another study, a novel gelatin–norbornene (GelNB) bioink was introduced for DLP bioprinting of vascularized constructs [190]. Norbornene-based macromers provide advantages over conventional (meth)acrylated polymers due to their mild and rapid crosslinking kinetics, high cytocompatibility, and potential for post-printing functionalization. Using low GelNB concentrations (2–5 wt.%) in combination with poly(ethylene glycol)-tetra-thiol (PEG4SH), researchers successfully fabricated soft hydrogels with tunable stiffness (120–4000 Pa) and perfusable vascular channels. Importantly, encapsulated human umbilical vein endothelial cells (HUVECs) remained viable and self-organized into interconnected microvascular networks with lumen structures. A key feature of the GelNB system was its ability to support both in situ tethering and secondary conjugation of bioactive motifs, such as the QK peptide, a vascular endothelial growth factor (VEGF)-mimetic sequence. Incorporation of QK peptide significantly enhanced endothelialization and vasculogenesis, demonstrating the bioink’s capacity to guide vascular morphogenesis. This work highlights GelNB as a versatile bioink platform that not only supports cytocompatibility and structural fidelity during DLP printing but also allows for post-printing biochemical functionalization, expanding its utility in vascular tissue engineering, wound healing, and regenerative medicine [190].

Despite these advances, challenges remain in establishing robust perfusion systems that can deliver oxygen and nutrients under physiological shear stress and in ensuring long-term scaffold stability. Nevertheless, these findings demonstrate how DLP-guided biofabrication can overcome major limitations of in vitro angiogenesis models and provide reproducible, modular vascular constructs for tissue engineering, organ-on-chip systems, and regenerative medicine.

DLP bioprinting supports the fabrication of vascular structures with precise geometries, stable lumens, and organized endothelial networks. Current bioinks show high viability, angiogenic activity, and the capacity to form perfusable channels. However, replicating physiological shear resistance, long-term perfusion, and mechanical stability remains challenging. Future improvements will require bioinks with better mechanical–biochemical balance and integrating pro-angiogenic and microfluidic strategies to enhance vascular function.

Corneal regeneration

Corneal diseases represent one of the leading causes of blindness worldwide, affecting millions of patients annually. Current clinical treatments, including corneal transplantation, are limited by donor shortages, risk of graft rejection, and post-surgical complications. As a result, there is an urgent need for alternative strategies that can restore corneal transparency, mechanical integrity, and biological function [191]. In this context, three-dimensional (3D) bioprinting has emerged as a transformative technology, enabling the fabrication of patient-specific corneal substitutes with high precision and tunable properties. Among the different bioprinting modalities, digital light processing (DLP) bioprinting offers unique advantages such as submicron resolution, smooth surface generation, and the ability to produce curved geometries without support structures features that are particularly critical for replicating the native corneal architecture.

Recent advances in DLP bioprinting have enabled the fabrication of artificial corneas with high precision, smooth curvature, and customizable geometry. In one study, researchers developed a composite decellularized corneal extracellular matrix–gelatin methacryloyl (CECM-GelMA) bioink specifically tailored for corneal applications [192]. This hybrid hydrogel system combined the biochemical cues of native CECM with the tunable mechanical and optical properties of GelMA, yielding constructs that preserved high transparency and exhibited adjustable stiffness suitable for corneal tissue engineering. The DLP-printed CECM-GelMA hydrogels supported the encapsulation of human corneal fibroblasts (hCFs) with excellent viability, while promoting the expression of cornea-specific matrix proteins. Importantly, in vivo implantation demonstrated the capacity of these constructs to facilitate epithelial regeneration, restore transparency, and maintain stromal alignment, key features for functional corneal repair. The integration of cornea-derived ECM was shown to play a critical role in guiding cellular behavior and enhancing bioactivity, while the DLP platform allowed fabrication of corneal substitutes with smooth surfaces and programmable double curvature, closely mimicking native anatomy. Collectively, these findings highlight the promise of CECM-GelMA bioinks in producing customizable, biocompatible, and optically transparent corneal equivalents. DLP bioprinting thus provides a powerful platform for developing patient-specific corneal implants, advancing both in vitro disease modeling and clinical regenerative therapies for vision restoration [192].

In another study, researchers employed a dual-network bioink composed of methacrylated hyaluronic acid (HA-MA) and gelatin methacryloyl (GelMA), crosslinked with ruthenium salt and sodium persulfate while tartrazine served as a photoabsorber [193]. Using DLP bioprinting, they successfully fabricated corneal lenticules with dimensions and mechanical performance comparable to native human cornea. The bioprinted lenticules exhibited excellent optical transparency (≈85%), adequate compressive modulus and adhesive strength, and high puncture resistance. Importantly, they supported re-epithelialization and stromal regeneration when seeded with human corneal stromal cells, confirming their potential as donor-independent corneal implants [193].

More recently, an interpenetrating GelMA-OxiCMC hydrogel was developed, leveraging dual crosslinking (Schiff base and photocrosslinking) to achieve a compressive modulus closely matching that of the native cornea (≈106 kPa) [194]. The constructs exhibited high shape fidelity, enhanced porosity for nutrient diffusion, excellent optical transparency after tartrazine removal, and supported > 93% keratocyte viability and proliferation. This system eliminated the need for synthetic polymers or toxic crosslinkers, marking an important step toward clinically relevant corneal stromal equivalents [194]. These advances demonstrate that DLP bioprinting allows for precise, donor-independent fabrication of patient-specific corneal substitutes with tunable mechanical, biological, and optical properties. While in vitro performance is highly promising, future directions will need to address full-layer corneal replication, long-term in vivo biocompatibility, and clinical translation.

The advances highlighted in this section show that DLP-printed corneal constructs can achieve the transparency, stromal alignment, and mechanical properties needed for functional repair, while supporting epithelial and stromal regeneration. Despite these advances, long-term optical stability, full-layer reconstruction, and extended biocompatibility remain unresolved. Future work should prioritize multilayered designs, improved transparency retention, and rigorous preclinical evaluation to move DLP-printed corneal substitutes closer to clinical application.

Uterus regeneration

Replicating the endometrium’s multi-layered structure and intrinsic vascular networks remains a significant challenge, primarily due to the lack of suitable bioinks and strategies for multi-material integration. Gelatin methacryloyl (GelMA) has been widely explored in tissue engineering due to its biocompatibility and cell-responsive features, but further chemical modifications, such as glycidyl methacrylate functionalization (GelMAGMA), have been reported to enhance printability, mechanical stability, and long-term functionality in bioprinted constructs. Combining GelMAGMA-based bioinks with endometrial stem cells, stromal cells, and endothelial cells has been shown to mimic both the basal and functional layers of the endometrium while integrating perfusable vascular channels [195]. Furthermore, the incorporation of perfusion culture systems and hormonal stimulation (e.g., estrogen) has been demonstrated to validate the physiological responsiveness of engineered tissues. Altogether, these developments highlight the potential of DLP-based bioprinting platforms for creating implantable endometrial substitutes and establishing advanced disease models in reproductive medicine [195]. DLP-based strategies can recreate multilayered uterine structures and perfusable vascular channels, especially using GelMAGMA bioinks combined with endometrial and endothelial cells. These systems also respond to hormonal cues such as estrogen. Remaining challenges include achieving long-term mechanical stability and fully replicating native uterine remodeling. Future work should refine bioinks and improve vascular and hormonal integration to move closer to functional, implantable constructs.

Liver regeneration

Liver transplantation remains the most effective treatment for end-stage liver disease, yet the persistent shortage of donor organs and tissues of limited graft integration underscore the urgent need for bioengineered alternatives [196]. In this context, 3D bioprinting and particularly digital light processing (DLP) has gained increasing attention for fabricating liver tissue substitutes that can restore hepatic functions.

Recent studies have demonstrated the potential of liver-specific bioinks, formulated by combining photocurable gelatin methacryloyl (GelMA) with liver-derived decellularized extracellular matrix (dECM) [197]. The inclusion of dECM not only enhances printability and structural fidelity but also provides organ-specific biochemical cues that support hepatocyte survival and function. For instance, when human-induced hepatocytes (hiHep) were encapsulated within GelMA/dECM bioinks, the printed constructs showed improved albumin secretion and urea synthesis compared to GelMA-only counterparts, highlighting the role of dECM in maintaining hepatocyte-specific activity. Furthermore, DLP bioprinting enables the fabrication of microtissues with intricate geometries, such as gear-like internal structures, which significantly increase surface area and promote mass transport within engineered constructs. These advances are essential for replicating the highly metabolic nature of the liver. Despite these achievements, challenges remain, including the recreation of the multicellular complexity of native liver tissue, which involves hepatocytes, stellate cells, Kupffer cells, and endothelial networks. Future directions are expected to focus on co-culture strategies, vascularization approaches, and in vivo validation of engineered liver constructs for long-term functionality and integration [197]. DLP-printed liver constructs based on GelMA/dECM bioinks show enhanced hepatocyte function and improved structural fidelity, offering a promising route toward donor-independent liver substitutes. The ability to fabricate intricate geometries also supports better nutrient transport in metabolically active tissues.However, replicating the full multicellular architecture of native liver and achieving stable long-term function remain significant challenges. Future work should emphasize co-culture systems, vascular integration, and extended in vivo evaluation to advance engineered liver tissues toward clinical relevance.

Muscle regeneration

Skeletal muscle disorders, including myopathies and muscular dystrophies, remain among the most challenging conditions to treat, primarily due to the tissue’s structural complexity and functional demands [198, 199]. Conventional therapeutic approaches offer limited regenerative potential, which has fueled significant interest in tissue-engineered muscle constructs for both disease modeling and regenerative medicine.

Recent progress in extrusion- and light-based bioprinting, particularly DLP, has facilitated the fabrication of muscle-mimetic scaffolds with precise microscale resolution [200]. However, the development of bioinks that not only support myogenic cell survival but also promote alignment, differentiation, and contractility remains a critical bottleneck. One promising strategy involves the chemical modification of bioactive molecules to improve bioink functionality. For instance, methacrylation of taurine has been explored to covalently integrate this bioactive molecule into GelMA networks, creating a photocrosslinkable hydrogel with enhanced structural fidelity and prolonged retention of taurine’s biological activity. Compared with physically blended counterparts, such covalently bound hydrogels exhibit improved crosslinking control, finer porosity, and sustained bioactive signaling [200]. Importantly, DLP-printed constructs derived from taurine-modified GelMA bioinks have demonstrated enhanced myoblast differentiation, increased myosin heavy chain expression, and more extensive myotube formation in vitro, indicating their potential as biologically active scaffolds for skeletal muscle tissue engineering. Looking forward, integrating such chemically functionalized bioinks with alignment-inducing printing strategies and vascularization approaches could pave the way for the generation of implantable, contractile muscle tissues suitable for both disease modeling and therapeutic applications.

DLP-based bioprinting enables the fabrication of muscle-like scaffolds with microscale precision and supports the incorporation of bioactive modifications, such as taurine-functionalized GelMA, which enhance myoblast differentiation and myotube formation. These advances highlight the potential of chemically tailored bioinks for engineering biologically active muscle constructs. Remaining challenges include achieving aligned, contractile architectures and integrating vascular support for thicker tissues. Future work should combine functionalized bioinks with alignment-guiding designs and vascularization strategies to move closer to implantable, therapeutically relevant muscle tissues.

Kidney regeneration

Kidney diseases, including chronic kidney disease and acute kidney injury, represent significant global health challenges with limited therapeutic options beyond dialysis and transplantation. The shortage of donor organs has driven the exploration of bioengineered renal tissues as alternatives for transplantation, drug testing, and disease modeling [201]. Three-dimensional (3D) bioprinting has emerged as a promising approach in renal tissue engineering, offering precise spatial control to fabricate complex, biomimetic structures. However, the development of organ-specific bioinks that replicate the native kidney microenvironment remains a critical challenge. Decellularized extracellular matrix (dECM)-based bioinks are particularly advantageous because they preserve tissue-specific biochemical cues that support cell viability, organization, and maturation [202].

In a study, a kidney-specific, photocrosslinkable bioink derived from methacrylated decellularized porcine kidney ECM (KdMA) was developed [203]. Porcine kidneys were decellularized using chemical detergents, enzymatic treatments, and sterilization steps to remove cellular components while preserving structural and functional ECM constituents. The resulting dECM was methacrylated to enable light-mediated crosslinking, forming a bioink with tunable stiffness and rapid polymerization kinetics suitable for both digital light processing (DLP) and extrusion-based bioprinting. Human embryonic kidney (HEK) cells encapsulated within the KdMA bioink demonstrated high viability, proliferation, and the ability to form multicellular spheroids, highlighting the cytocompatibility and organ-specific functionality of the bioink.Multilayered kidney constructs fabricated with the KdMA bioink exhibited high structural fidelity, maintaining mechanical integrity and microarchitecture critical for tissue-specific function. Rheological analyses confirmed that the bioink could be tuned to match the mechanical properties of native renal tissue, while morphological assessments demonstrated a conducive environment for cell organization and tissue maturation. These results indicate that kidney-specific dECM bioinks provide both structural support and biochemical cues necessary for engineering functional renal tissues in vitro [203].

Kidney-specific dECM bioinks demonstrate strong potential for renal tissue engineering by providing organ-matched biochemical cues, tunable mechanics, and high structural fidelity. Encapsulated HEK cells show high viability and organization, indicating suitability for constructing biomimetic renal microarchitectures. However, recreating the full nephron complexity and establishing functional filtration and perfusion remain major challenges. Future work should integrate multicellular co-culture systems, microfluidic flow, and long-term functional evaluation to move engineered renal tissues toward translational applications.

Dental regeneration

Digital light processing (DLP) bioprinting offers high-resolution fabrication of hydrogel-based constructs, yet achieving high-performance hydrogels that maintain both structural fidelity and cellular functionality remains challenging. A practical strategy involves the use of gelatin methacryloyl (GelMA) emulsions containing dextran microdroplets, which act as decoys to aggregate stem cells and facilitate the in-situ formation of spheroids within the printed constructs. The GelMA matrix provides mechanical support, while the controlled leaching of dextran after printing generates smooth curved surfaces that enhance cell aggregation and differentiation potential. This approach enables the fabrication of hydrogel constructs with complex geometries while maintaining high cell viability [204].

Dental pulp stem cells (DPSCs) encapsulated in the GelMA/dextran emulsion were observed to form spheroids in situ, promoting enhanced regenerative capacity [204]. The bioink was characterized in terms of microdroplet morphology, emulsion stability, and compressive mechanical properties, using confocal microscopy and mechanical testing. The dextran microdroplet size and distribution were optimized to support uniform spheroid formation, while the GelMA matrix maintained structural integrity during and after DLP bioprinting. The resulting hydrogel constructs demonstrated high cell viability, enhanced aggregation of stem cells into spheroids, and improved differentiation potential. Curved microarchitectures and high cell density within the dextran phase contributed to spheroid formation, which in turn promoted regenerative dentin formation and vascularization in vivo. This approach represents a feasible and effective strategy for engineering dental tissue constructs and highlights the potential of cell-concentrated bioinks in regenerative medicine applications [204]. DLP bioprinting has shown strong potential for engineering dental constructs by enabling controlled cell aggregation, high cell viability, and enhanced differentiation of dental pulp stem cells key factors that support dentin formation and regeneration. However, long-term mechanical stability and accurate reconstruction of the vascular dentin interface remain major challenges. Future research should focus on improving microarchitectural control, promoting angiogenesis, and refining printing strategies to advance DLP-printed dental tissues toward clinical applications.

Organ-on-a-chip and drug testing

Organs-on-a-chip (OOC) or microphysiological systems (MPSs) are microfluidic devices composed of parts of natural or engineered tissue and organs, allowing for the mimicking and simulation of the physiological structure and function of native organs. They also create a dedicated microenvironment to control and maintain cell and tissue activity, providing accurate feedback on complex biological processes. In addition, these systems are powerful tools for disease modeling, drug testing, personalized medicine, and other applications. Recent advances in 3D bioprinting and biomaterials for the production of OOCs have led to the creation of complex constructs with the ability to mimic the function of native tissue/organs [205, 206]. In one study, researchers optimized PEG diacrylate/GelMA-based composite hydrogel bioinks for rapid prototyping of microfluidic chips using DLP technology. The results showed that the mechanical properties of the chips depended on the PEGDA:GelMA ratio. After encapsulation of human-derived tumor cells, the chips were shown to provide a cell-friendly environment, exhibit good bioactivity, and have significant efficacy for vascular microtissues. Therefore, this approach provides a platform for organ-on-a-chip modeling and drug testing with high throughput [89]. In another study, a 3D culture model based on GelMA hydrogel embedded with human induced pluripotent stem cell-derived liver cells (hiPSC-HPCs), human umbilical vein endothelial cells (HUVECs), and adipose-derived stem cells (ADSCs) was developed for the study of liver disease and drug screening. The results confirmed that this 3D model showed increased expression levels of liver-specific genes, increased induction of cytochrome P450, increased secretion of metabolic products, and cellular morphological organization over weeks compared with other models. These findings indicate that DLP bioprinting technology can be used to design a 3D biomimetic liver model and also enable disease modeling and drug screening [207]. Ma et al. designed hydrogel chips based on decellularized liver extracellular matrix (dECM) to study the disease progression of hepatocellular carcinoma (HCC). The dECM chips exhibited mechanical properties similar to those of cirrhotic liver tissue. After encapsulating HepG2 cells in the engineered platform with cirrhotic stiffness, they showed increased expression of invasion markers, decreased cell growth, and were able to simulate HepG2 stromal invasion from the cirrhotic stiffness node. These results represent a significant advance in the design of 3D platforms for in vitro disease modeling [208]. In another study, an in vitro animal-derived breast cancer model for drug screening and testing was developed using DLP technology. The use of DLP enabled rapid printing of a system with vessel-like channels to replicate the tumor microenvironment. The bioprinted model showed good cellular compatibility while maintaining the cellular phenotype. Furthermore, the replicated microenvironment in the construct via vessel-like channels in the bioprinted model showed a significant reduction in drug chemoresistance. Therefore, the researchers confirmed that this model is a promising approach for designing drug testing platforms [209]. To provide a comprehensive overview of the applications of digital light processing (DLP) bioprinting across various tissue and organ systems, Table 4 summarizes the most recent advances, highlighting bioink compositions, cell types, structural features, and functional outcomes reported in different studies.

Table 4.

Summary of recent advances in digital light processing (DLP) bioprinting for regenerative medicine across different tissue and organ systems

Application	Bioink/Modification	Cells	Key Features	Outcomes/Applications	Ref.
Wound Healing	GelMA + COS	Fibroblasts	Antibacterial, antioxidative	Accelerated healing, reduced infection	[173]
	Silk-GMA	Keratinocytes, Fibroblasts	Mechanical strength, elasticity	Reduced scarring, enhanced regeneration	[174]
	CMA + DPPA	Endothelial cells	Angiogenic, pro-healing	Vascularization, wound closure	[58]
	CMA-DHM	Fibroblasts	Anti-fibrosis, sustained activity	Improved remodeling, minimized scar tissue	[91]
Bone & Cartilage	GelMA + PMAA	BMSCs	High viscoelasticity, osteoinduction	Osteogenesis, bone repair	[177]
	GelMA + Dextran	Chondrocytes	Emulsion, mechanical fidelity	Cartilage regeneration	[178]
	MA-κ-CA + BSNPs	hMSCs	Enhanced stiffness, bioactivity	Bone/cartilage scaffolds	[179]
	Silk-GMA	MC3T3–E1	ECM mimicry	Osteoblast proliferation	[180]
	SF-MA	MC3T3–E1	ECM mimicry	Osteoblast proliferation	[181]
	GelMA + CaP/GO	hMSCs	High print precision	Patient-specific scaffolds	[182]
Vascular Tissue	PEGDA + GelMA	HUVECs	Perfusable channels	Angiogenesis, vascularization	[186]
	GelMA + GO + CaP	VECs	Tunable stiffness	Endothelialization	[182]
	GelNB	EPCs	Photo-crosslinkable	Microvascular networks	[188]
	GelMA + PEG + Alginate	VSMCs	High resolution	Functional vascular tissue	[190]
Corneal Regeneration	CECM-GelMA	hCFs	Transparency, smooth curvature	Corneal stromal regeneration	[192]
	HA-MA + GelMA	Stromal cells	Tunable stiffness	Optical clarity	[193]
	GelMA-OxiCMC	Corneal stromal cells	Donor-independent, high fidelity	Corneal repair, re-epithelialization	[194]
Uterus Regeneration	GelMAGMA-based bioinks	Endometrial stem, stromal, endothelial cells	Multilayered endometrium mimicry, vascular channels	Implantable substitutes, disease modeling	[195]
Liver Regeneration	GelMA/dECM	hiHep cells	Organ-specific cues, high fidelity	Hepatic function restoration, liver modeling	[197]
Muscle Regeneration	Taurine-modified GelMA	Myoblasts	Alignment, differentiation, contractility	Myotube formation, muscle tissue engineering	[200]
Kidney Regeneration	KdMA (kidney ECM bioink)	HEK cells, proximal tubule, iPSC-organoids	Tunable stiffness, multilayered structures	Functional renal tissue, drug screening	[203]
Dental Regeneration	GelMA/dextran emulsion	DPSCs	In-situ spheroid formation, curved architecture	Dentin regeneration, vascularization	[204]
Organs-on-a Chip	PEGDA/GelMA	Tumor cells	Rapid prototyping, tunable mechanics	Vascular microtissues, drug testing	[89]
	GelMA composite	hiPSC-HPCs, HUVECs, ADSCs	3D liver-like model	Disease modeling, drug screening	[207]
	dECM-based chips	HepG2 cells	Cirrhotic stiffness mimicry	HCC progression modeling	[208]
	GelMA with vessel-channels	Breast cancer cells	Tumor microenvironment replication	Reduced drug resistance	[209]

DLP bioprinting enables the rapid fabrication of microfluidic and multicellular platforms that more accurately replicate native tissue physiology, offering improved performance for disease modeling and drug testing compared with traditional culture systems. These printed constructs support organized cell behavior, relevant metabolic activity, and controlled microenvironmental features. Key challenges include sustaining long-term perfusion, recreating full organ-level complexity, and standardizing functional outputs. Future efforts should focus on integrating vascularized architectures, enhancing dynamic microfluidic control, and refining tissue–device interfaces to advance DLP-based organ-on-a-chip systems as reliable preclinical testing tools.

Challenges and future perspective

The development of tissue-specific bioinks and DLP-based bioprinting are advancing rapidly, but several critical barriers remain in the way of clinical implementation. There is a significant challenge in reproducing the native microenvironment, including heterogeneous mechanical gradients, anisotropic ECM organization, and perfusable vascular networks. Oxygen and nutrient transport are often restricted in thick or multilayered constructs, resulting in hypoxic regions and reduced long-term viability of cells. Furthermore, due to limitations in printing resolution, curing depth, and cell-specific phototoxicity, achieving precise spatial placement of multiple cell types is a technical challenge. There are additional constraints associated with the development of bioinks. In spite of the fact that methacrylated hydrogels, dECM-based inks, and functionalized polymers provide tunable mechanics and cytocompatibility, their batch variability, limited long-term stability, and insufficient recapitulation of biochemical cues prevent them from providing full tissue maturation. Cytotoxicity of photoinitiators, suboptimal light penetration, and reactive oxygen species (ROS) generated during photocuring pose additional risks to encapsulated cells, underscoring the need for safer initiation systems and wavelength-shifted curing techniques. Moreover, current bioinks often lack dynamic or adaptive behavior, preventing them from responding to mechanical loading or biochemical changes in the surrounding environment after implantation. These limitations can be addressed through several mitigation strategies in the future. By using oxygen-tolerant photochemistry, visible-light initiator systems, and low-ROS photoinitiators, cytotoxicity associated with curing may be significantly reduced. By incorporating dynamic polymers (e.g., thermoresponsive, enzymatically degradable, or mechano-adaptive networks), printed constructs can be better integrated and matured. It is possible to reproduce complex mechanical gradients, hierarchical ECM structures, and spatially organized multi-cellular environments using multi-material DLP and hybrid bioprinting platforms. Furthermore, embedding pre-vascularized networks, growth factor-releasing microsystems, or microfluidic perfusion channels can enhance mass transport and sustain long-term viability. Finally, the integration of real-time imaging, biosensors, and AI-assisted feedback control may enable continuous monitoring and dynamic adjustment of printing parameters, which may accelerate the development of clinically relevant tissues. Taking these strategies together, they outline realistic ways to overcome current limitations and advance DLP-based tissue engineering toward scalable, functional, and translational applications.

Conclusion

Digital light processing (DLP) bioprinting, combined with organ- and tissue-specific bioinks, has demonstrated strong potential for generating anatomically precise and functionally relevant tissue constructs. Recent advances in hydrogel chemistry, decellularized ECM formulations, and multicellular patterning have enabled the fabrication of preclinical models for corneal, endometrial, hepatic, renal, muscular, and dental tissues. These constructs exhibit high cell viability, appropriate mechanical or optical properties, and tissue-specific differentiation. These findings highlight the growing relevance of DLP platforms for regenerative medicine, disease modeling, and drug screening.

However, several limitations still prevent clinical translation. Recreating physiological vascularization, long-term functional stability, and multiscale tissue heterogeneity remains challenging. In addition, issues such as phototoxicity, limited curing depths, and insufficient biochemical adaptability of current bioinks underscore the need for safer photochemistry and more sophisticated biomaterials.

Looking forward, meaningful progress will require the integration of dynamic and stimuli-responsive hydrogels with oxygen-tolerant or visible-light photoinitiators. Multimaterial DLP strategies and perfusion-enabled architectures that support thick, metabolically active tissues are also needed. In the near future, advances in microfluidics, real-time biosensing, and artificial intelligence-based control may further enhance the fidelity and scalability of printed constructs. Together, these innovations outline a realistic pathway toward clinically actionable DLP-printed tissue substitutes and robust precision medicine platforms.

Abbreviations

3D

Three-Dimensional

ADSCs

Adipose-Derived Stem Cells

AVN

Azobisvaleronitrile

BAPO

Phenyl bis(2,4,6-trimethylbenzoyl) phosphine oxide

BMSCs

Bone Marrow Stem Cells

CaP

Calcium Phosphate

CECM

Corneal Extracellular Matrix

CMA

Dimethyl Acrylated Collagen

COS

Chitosan Oligosaccharide

CQ

Camphorquinone

dECM

Decellularized Extracellular Matrix

DHM

Dihydromyricetin

DLP

Digital Light Processing

DMD

Digital Micromirror Device

DPPA

Dimethylphenylphosphonate

DPSCs

Dental Pulp Stem Cells

ECM

Extracellular Matrix

EPCs

Endothelial Progenitor Cells

ESRD

End-Stage Renal Disease

GelMA

Gelatin Methacryloyl

GelNB

Gelatin–Norbornene

HA-MA

Methacrylated Hyaluronic Acid

HAMA

Methacrylated Hyaluronic Acid

HCC

Hepatocellular Carcinoma

hCFs

Human Corneal Fibroblasts

HDFs

Human Dermal Fibroblasts

HEK

Human Embryonic Kidney

HepG2

Human Hepatocytes

hiPSC-HPCs

Human Induced Pluripotent Stem Cell-Derived Liver Cells

hMSCs

Human Mesenchymal Stem Cells

HUVECs

Human Umbilical Vein Endothelial Cells

KdMA

Methacrylated Decellularized Porcine Kidney ECM

LAP

Lithium Phenyl-2,4,6-trimethylbenzoylphosphinate

MA-κ-CA

Methacrylate-κ-Carrageenan

MC3T3-E1

Pre-osteoblast Cell Line

MPSs

Microphysiological Systems

NIR

Near-Infrared

OOC

Organs-on-a-Chip

PCL

Poly(caprolactone)

PEG

Polyethylene Glycol

PEGDA

Poly(ethylene glycol) Diacrylate

PMAA

Polymethacrylic Acid

PPF

Poly(propylene Fumarate)

PVA

Poly(vinyl alcohol)

QK

VEGF-Mimetic Peptide

RGD

Arginine-Glycine-Aspartic Acid

Ru(bpy)₃Cl₂

Ruthenium Bipyridine

SF-MA

Methacrylated Silk Fibroin

Sil-MA

Silk Fibroin Methacryloyl

Silk-GMA

Silk Fibroin–Glycidyl Methacrylate

TPO

2,4,6-trimethylbenzoyl diphenyl phosphine oxide

UV

Ultraviolet

VEGF

Vascular Endothelial Growth Factor

VSMCs

Vascular Smooth Muscle Cells

Author contributions

HR, MH, MS, DD, MN, NA and SA designed the review study and contributed to writing the manuscript draft. HR, MH, DD, MN, NA and SA searched the literature and contributed to writing the manuscript. MS contributed to the literature search and edited the manuscript. All authors have confirmed the final version of the manuscript and are accountable for the contents of all parts of the work. All authors read and approved the final manuscript.

Funding

This research received no external funding.

Data availability

No datasets were generated or analysed during the current study.

Declarations

Ethics approval and consent to participate

Not applicable.

Consent to publish

Not applicable.

Aritifical intelligence (AI) statement

The authors declare that no AI-assisted technologies were used in the generation of this manuscript.

Competing interests

The authors declare no competing interests.