3D Bioprinting for Engineered Tissue Constructs and Patient-Specific Models: Current Progress and Prospects in Clinical Applications

Abstract

Advancements in bioprinting technology are driving the creation of complex, functional tissue constructs for use in tissue engineering and regenerative medicine. Various methods, including extrusion, jetting, and light-based bioprinting, have their unique advantages and drawbacks. Over the years, researchers and industry leaders have made significant progress in enhancing bioprinting techniques and materials, resulting in the production of increasingly sophisticated tissue constructs. Despite this progress, challenges still need to be addressed in achieving clinically relevant, human-scale tissue constructs, presenting a hurdle to widespread clinical translation. However, with ongoing interdisciplinary research and collaboration, the field is rapidly evolving and holds promise for personalized medical interventions. Continued development and refinement of bioprinting technologies have the potential to address complex medical needs, enabling the development of functional, transplantable tissues and organs, as well as advanced in vitro tissue models.

1. Introduction

In the 2015 release of “The Avengers- The Age of Ultron,” a notable scene depicts scientist Dr. Helen Cho utilizing 3D printing technology to regenerate a body, leaving a lasting impression. The emergence of 3D printing technology, a method of 3D processing, gained prominence in various media outlets starting in 2012. Subsequently, 3D printing technology swiftly made a profound impact on numerous industries, especially in biomedical research. The allure of 3D printing lies in its revolutionary processing capabilities, enabling the relatively straightforward, layer-by-layer manufacturing of intricate structures. This attribute has positioned 3D printing as a transformative technology with far-reaching implications, particularly in tissue engineering and regenerative medicine applications.

The concept of contact printing plant cells was first described in 1990, followed by the use of animal cells in 1992.^[1] The integration of 3D printing technology to print biodegradable polylactide (PLA) scaffolds was described in 1996.^[2] The use of both cells and hydrogels sequentially for printing was reported in 2003.^[3] The 3D printing of viable tissue implanted in vivo, amnion stem cell-derived bone constructs, was first reported in 2006,^[4] leading others to explore the use of cells in bioprinting.^[5] This sequence of developments underscores the foundational progress in bioprinting, illustrating the transition from early cell printing techniques to the integration of 3D printing with viable tissue constructs, setting the stage for the field’s evolution toward more complex in vivo applications.

The primary objective of utilizing 3D printing technology in tissue engineering and regenerative medicine is to innovatively surpass the limitations of traditional manufacturing methods, thereby producing tissue constructs with practical clinical utility. Initially introduced in 1996, the technology focused on developing regular porous scaffolds using biodegradable polymers for cell culture. Despite advancements, attempts since 2003 to replicate human tissue by directly printing cells revealed significant limitations in creating stable, transplantable tissue constructs. Recognizing the need for more sophisticated and stable cell-laden tissue constructs, an Integrated Tissue and Organ Printing (ITOP) system was developed to address these challenges. Introduced at the Tissue Engineering and Regenerative Medicine International Society (TERMIS) in 2010, the ITOP system operates on the principle of simultaneously printing various biomaterials and cells to manufacture unified tissue structures.^[6] The printed polymeric material ensures the stability of the tissue structures, while specific cells composing the tissue are precisely printed and organized. In 2016, we successfully applied this technology to manufacture human-scale tissue constructs with closer relevance to clinical applications.^[7]

Tissue engineering seeks to restore damaged tissues or organs by creating biomaterial scaffolds with cells, shaping them into specific forms for implantation into patients. While various manufacturing methods have been developed for this purpose, conventional techniques need to be improved in producing intricate tissue constructs suitable for clinical applications. 3D printing technologies have emerged as a transformative solution, overcoming these limitations and expanding the potential for tissue engineering and regenerative medicine. The utilization of 3D printing allows for the reconstruction of complex tissue geometries, the fabrication of composite tissues with seamless connections between different tissue types, and the creation of microvasculature and functional units within solid organ structures. Moreover, these technologies enable the production of in vitro tissue models featuring microstructures that mimic specific organ functions. Such models serve as valuable devices for assessing drug toxicity and effectiveness. This review provides an overview of the current state and prospects of 3D bioprinting applications in tissue engineering and regenerative medicine.

2. 3D Bioprinting Technologies

3D printing, also referred to as additive manufacturing or rapid prototyping, has emerged as a groundbreaking technology with far-reaching implications across diverse industries. Its versatile potential has found applications in sectors such as manufacturing, electronics, architecture, automobiles, and healthcare. Diverging from traditional subtractive and formative manufacturing methods, 3D printing facilitates the creation of intricate objects through the flexible utilization of materials, from metals to polymers. This technology integrates computer-aided design (CAD) and computer-aided manufacturing (CAM), wherein the conversion of 3D CAD into 3D CAM software enables the fabrication of complex objects. This is achieved through dot, line, and layer-by-layer deposition guided by axis movement. While numerous 3D printing techniques have been developed for various purposes, only a few have been employed for 3D bioprinting applications. This review focuses specifically on extrusion, jetting, and vat photopolymerization (light) printing techniques, as these are the most commonly used in the field. It also explores advanced 3D printing methods designed to address the limitations of these major techniques.

2.1. 3D printing

Among the various 3D printing techniques, extrusion, jetting, and light-based are primary approaches for bioprinting. Extrusion-based bioprinting dispenses materials (100–500 μm) in a continuous string form. Pneumatic systems provide precise material control, while mechanical systems ensure spatial control. Jetting-based bioprinters deposit hydrogel “ink” in controlled volumes at specific locations (50–200 μm) using thermal or piezoelectric actuators. Light-based bioprinting enables the creation of high-resolution 3D models (25–50 μm) with precise energy control to minimize cell damage.

Extrusion-based printing employs mechanical deposition of materials through a nozzle, performing line-by-line deposition of materials with varying viscosities while the stage moves along x-y-z axes. This printing technique has been used in fused deposition modeling (FDM), pneumatic, and plunger dispensing (Figure 1A). However, extrusion printing faces challenges such as nozzle clogging and high shear stress due to relatively high viscous bioinks. The layer-by-layer stacking and micropatterning required in this technique involve thorough optimization of parameters like printing speed, extrudability, dynamic viscosities, and temperature, impacting the integrity of the printed structure.

Figure 1.

3D printing methods for cell-based bioprinting. Conventional 3D printing: (A) Extrusion, (B) jetting, and (C) light-based printing methods. Advanced 3D printing methods; (D) Embedded printing using supporting bath, (E) bead-jetting printing, and (F) volumetric printing by vat rotation and tomography-projection.

Jetting-based printing ejects low viscous materials through drop-by-drop dispensing using micro-nozzles for dropletization, spatial patterning, and gradient control (Figure 1B). Point-by-point positioning enables dynamically patterning drugs, adhesive reagents, and microparticles by selective crosslinking for high-throughput application. This method employs various jetting techniques, including binder, thermal, piezoelectric, and electrohydrodynamic jetting. However, the droplet is limited to layer stacking fidelity due to low viscosity and additional crosslinking time.

Light-based printing, specifically vat photo-polymerization of photoactive materials, utilizes light for crosslinking and enables layer-by-layer fabrication of complex geometries (Figure 1C). This method is employed in stereo-, soft-, multiphoton-lithography, and digital light processing (DLP). Lithography requires post-processing steps such as post-curing and washing the organic solution. Light-based printing needed to use photoactive materials with low viscous behavior to reduce the drag effect on layer stacking by the z-axis. In addition, microparticles in the photoactive materials hindered the printability due to light scattering and inhomogeneities caused by sedimentation. The usage of temperature-sensitive materials was limited due to the temperature gap between the printed materials and vat over time, and the cumulative light exposure during the vat polymerization process could be the cause of severe cell damage. Careful selection of the printing technique is essential based on the application and biomaterial properties.

2.2. Advanced bioprinting technologies

3D bioprinting techniques have evolved to address the limitations associated with conventional 3D printing methods. Extrusion-based bioprinting requires using relatively high-viscosity hydrogel bioink systems to uphold structural integrity during the printing process. However, obtaining such high-viscosity hydrogels often involves high concentrations, extensive crosslinking, or rapid crosslinking processes that may impact the cells within the printed constructs. In advancing extrusion bioprinting, embedded printing, also known as “printing into support baths,” offers a solution to achieve precise and intricate structures without the need for sacrificial layers (Figure 1D).^[8] This technique involves depositing a hydrogel precursor ink into temporary, thermoreversible, and biocompatible supporting bath materials, such as gelatin, alginate, Carbopol, and Pluronic F-127.^[9] Following the printing process, selective crosslinking of the structure occurs, and the completed structure is released from the support material through environmental changes, such as alterations in temperature. This fabrication strategy combines support and extruded bioink, incorporating a differential crosslinking mechanism. Freeform reversible embedding of suspended hydrogels (FRESH) technique represents the use of a thermoreversible support bath to allow the deposition of hydrogels into complex 3D biological structures.^[10] Conversely, the sacrificial writing into functional tissue (SWIFT) strategy involves printing sacrificial bioinks into a cell-containing matrix to create embedded vascular channels within organ building blocks.^[11]

Jetting-based bioprinting faces limitations in depositing high cell concentrations through micro-sized nozzles, achieving spatial precision by splashing the solution, experiencing slow crosslinking, and encountering challenges in effective scale-up while maintaining high accuracy and droplet integration. To address these limitations, using microfluidics, a modified approach applies jetting printing to sparsely deposited cell-containing beads (Figure 1E). This bead-jet printing technique allows for the reconstruction of volumetric muscle loss and skin augmentation, coupled with hair follicle regeneration, achieved through the in situ positioning of cell-containing Matrigel in a high-throughput manner.^[12] The process involves the in situ compartmentalization of droplets using air microfluidics, leveraging surface tension via Marangoni flow to encapsulate core spheroids within a hydrogel precursor solution for ionic crosslinking. To prevent nozzle clogging, a nozzle-free jet printing method using acoustic drop ejection has been developed.^[13] Furthermore, an approach called Alternating Viscous and Inertial Force Jetting (AVIFJ) has been introduced.^[14] This method employs automation with artificial image recognition and real-time cell distribution to develop a liver carcinoma model for drug screening. The technique controls endothelialization into hepatocyte spheroids, presenting the potential for advanced applications in bioprinting technologies.

Conventional light-based bioprinting faces limitations such as single-material vat polymerization, non-uniform cell distribution due to gravity, and challenges in achieving high printability due to light scattering in the presence of high cell concentrations. Moreover, light-based crosslinking poses risks of increased cell death and DNA mutation due to cumulative UV exposure. A novel volumetric printing method has been developed to mitigate the issues associated with cumulative UV exposure during vat polymerization, employing tomographic projection through the Radon transform algorithm (Figure 1F).^[15] This tomograph-based volumetric bioprinting approach allows for the rapid fabrication of cell-laden tissue constructs (>100 mm³) within seconds, achieving high viability (>85%). The application of this technology extends to the creation of trabecular bone constructs and the development of bone screws, utilizing CT data for precision. For multi-material volumetric printing, photo-click chemistry has been introduced to expedite complex shape engineering.^[16] This involves employing a differential crosslinking mechanism by altering resin vats and pre-filled layers of resins to fabricate channel and micropatterned structures. Additionally, acoustics has been integrated into the volumetric printing to allow for simultaneous reinforcement patterning and printing of the entire construct.^[17] These innovative approaches enhance the versatility and capabilities of light-based bioprinting technologies.

Several technology-driven approaches are being explored to address the limitations of current bioprinting methodologies. One such strategy involves the precise manipulation of various materials using microfluidic chips, enabling the creation of complex structures with multiple material structures^[18] Additionally, a novel multimaterial-multinozzle printing is being employed to accelerate the manufacturing of complex structures using a diverse range of materials.^[19] In parallel, the development of in situ tissue fabrication techniques has been investigated, allowing for the direct deposition of preformed-organoids into damaged tissues^[20] or the in situ photocuring of materials, thereby enhancing the applicability of bioprinting in intraoperative settings.^[21] These advanced approaches aim to improve the precision and effectiveness of bioprinting applications directly within surgical environments. The ultimate goal is to overcome the current technological challenges and develop more functional tissue constructs, benefiting a more significant number of patients.

2.3. Bioinks

To enable the bioprinting process, a crucial component is the utilization of a ‘bioink’ as the printing medium for biomaterials or cells.^[22] Bioinks, designed to accommodate living cells, predominantly consist of aqueous and hydrogel formulations where cells are suspended. It is important to note that cells alone lack printability. Therefore, the presence of hydrogel support is imperative to effectively print cells and construct intricate 3D tissue structures. Hydrogel refers to a 3D network of hydrophilic polymers capable of retaining substantial water while preserving its structure through the chemical or physical cross-linking of individual polymeric chains. Notably, the hydrogel’s high water content imparts flexibility like natural tissues. Furthermore, the mechanical properties of hydrogels are tunable, allowing for replication of tissue-specific extracellular matrices (ECMs).^[23]

Naturally derived materials, including proteins (e.g., collagen, gelatin, fibrinogen), polysaccharides (e.g., sodium alginate, gellan gum), and glycosaminoglycans [e.g., hyaluronic acid (HA)], serve as versatile bioinks in bioprinting. With exceptional biological properties and distinctive physical characteristics, these materials play a crucial role, offering diverse properties for tailored bioink formulations. And tissue-derived ECM components offer promise for tissue-specific bioprinting due to preserved ECM proteins and growth factors that mimic the tissue-specific microenvironment. Poly(ethylene glycol) diacrylate (PEGDA) is a prevalent synthetic hydrogel extensively employed in stereolithography-based bioprinting for creating high-resolution, cell-laden constructs. Despite its widespread use, the drawback of PEGDA-based constructs lies in their limited protein binding sites, resulting in suboptimal cell behaviors.

The advancement of bioink systems is intricately linked with the selected printing methods. The enhancement of bioink printability relies on hydrogel properties, cross-linking mechanisms, and printing parameters aligned with the specific printing technologies employed. For instance, in extrusion-based bioprinting, the optimal utilization of higher-viscosity hydrogels with strong shear-thinning behavior is observed. In contrast, stereolithography-based bioprinting necessitates the use of bioinks with relatively lower viscosity. This approach highlights the significance of modifying bioink formulations to match the specific demands of varied printing methodologies.

The use of biomaterials in bioprinting is constrained by the need to satisfy both fundamental conditions (such as biocompatibility, biodegradability, and physical properties) essential for tissue engineering and regenerative medicine, as well as those specific to the employed printing methods. Consequently, biomaterials have been tailored to achieve the desired properties for the 3D printing process through various strategies, including formulations, nanomaterial incorporations, and chemical modifications (Figure 2). This thoughtful approach significantly contributes to improving the overall efficiency and effectiveness of the bioprinting process.

Figure 2.

Types of biomaterials as bioinks for cell-based bioprinting. Hydrogels, specifically, are customized to meet 3D printing requirements through formulations, nanomaterial incorporations, and chemical modifications.

2.3.1. Considerations of bioink systems for cell-based bioprinting

Successful 3D bioprinting relies on the crucial printability of hydrogel-based bioink, a factor influenced by various manufacturing techniques such as extrusion, jetting, and light-based printing.^[22] In extrusion-based bioprinting, hydrogel-based bioinks must meet specific criteria: (i) optimal viscosity for homogeneous cell suspension, (ii) strong shear-thinning behavior to minimize cell damage, and (iii) rapid crosslinking for constructing 3D tissue architecture.^{[5, 24]} A rapid crosslinking process is vital, as many hydrogels lack self-support during layer-by-layer deposition.^[25] Crosslinking, mostly physical crosslinking, before printing increases shear stress, risking cellular damage and nozzle clogging. Crosslinking after printing affects resolution as bioink may spread between extrusion and cross-linking, leading to incomplete cross-linking in large multilayered constructs.^[26] In the case of embedded printing, the support bath with both shear-thinning and self-recovery properties is additionally required. The self-recovery ensures that the bioink regains its structure after extrusion, maintaining the integrity of the printed shape within the bath.

The jetting-based bioprinting approach is gaining attention due to its remarkable advantages, including contactless printing, substrate functionalization, and drop-on-demand capabilities.^[27] Furthermore, material jetting allows for precise deposition patterns and volume control. The printing process involves two main phases: jetting and impacting, with the impacting phase further divided into nonpenetrative and penetrative droplet impacts. Both Newtonian and non-Newtonian fluids with suitable fluid properties can be utilized in jetting-based bioprinting. The dimensionless number Z is employed to describe droplet formation during the jetting phase. Z, the reciprocal of the Ohnesorge number (Oh = We^1/2/Re), evaluates droplet stability.^[28] Oh represents the ratio of the Reynolds number (Re) to the Weber number (We), reflecting the balance between surface energy and viscous dissipation dictating droplet formation. Stable jet formation typically occurs when the Weber number falls within the range of 2 < We_j < 25. However, there remains to be a greater understanding of the interaction between non-Newtonian materials like cell-laden hydrogels and substrates.^[27b] Understanding the interaction between jetted droplets and various substrates is crucial for enhancing printability. Substrate characteristics can be modified by adjusting factors such as surface energy, roughness, plasma treatment, and wettability. These modifications influence droplet spreading, adhesion, and overall print quality, underscoring the importance of substrate optimization in material jetting for advanced bioprinting applications.

In light-based bioprinting, hydrogel bioinks undergo photo-polymerization through a light pattern (365, 385, and 405 nm) to form 3D tissue structures.^[29] A comprehensive understanding of the photo-curing mechanism is crucial. Briefly, a photoinitiator, when exposed to light, generates free radicals (initiation). These radicals initiate a chain reaction, forming a strong polymer network through hydrogen abstraction from polymer double bonds (propagation). The process concludes with two free radical species reacting to create a stable, non-radical state (termination). Hydrogel-based bioinks must align with lithographic process requirements, distinct from those of extrusion-based bioprinting.^[30] Commonly used photo-crosslinkable materials include gelatin methacryloyl (GelMA) and PEGDA.^[31] Careful selection of a water-soluble, biocompatible photoinitiator with a wavelength of maximum absorbance matching the light source is essential. Additionally, light absorbers like color dyes can be added to prevent the over-curing of layers beyond the focal plane, enhancing printing resolution.^[32] Light-based bioprinting requires bioinks with relatively low viscosity to facilitate smooth relocation of the printing substrate. This lower viscosity allows for precise control during the curing process, minimizing the risks of over-curing or uneven layer formation. Conversely, volumetric printing necessitates high-viscosity bioinks to counteract the effects of gravity over time, as the rotation of the suspension bath may cause the printed pattern to collapse. The higher viscosity ensures structural stability, preventing deformation due to gravitational forces.

The success of cell-based bioprinting, aimed at constructing a 3D tissue construct with live cells, relies heavily on ensuring cell viability post-printing - a key criterion for bioink printability. To sustain cell survival, a biological microenvironment is essential, not only during printing but also in subsequent culture.^[33]

2.3.2. Bioink formulations

For hydrogels designed for extrusion-based printing cells, specific attributes such as a defined viscosity, high shear-thinning to minimize cell damage during printing, and rapid crosslinking to stabilize printed structures become essential. The continual development of novel bioink materials holds the key to unlocking further possibilities in tissue engineering applications. The bioinks for cell-based extrusion printing should fulfill the basic requirements discussed above.^{[24a, 34]} To achieve this goal, numerous hydrogel formulations have been developed. This approach for developing bioink systems involves formulating hydrogels by simply mixing multiple components. For instance, in our previous studies,^{[7, 35]} a bioink formulation was produced by combining cross-linkable hydrogel materials with various supporting hydrogel materials, including gelatin, HA, and glycerol. Gelatin was used due to its thermo-sensitive properties: it is a liquid form above 37°C and becomes a solid form below 25°C. HA functions to enhance dispensing uniformity, while glycerol prevents nozzle clogging. After cross-linking, the uncross-linked components (gelatin, HA, and glycerol) can be gradually washed out under the culture condition.

Light-based bioprinting commonly utilizes PEG-based bioink systems, known for their excellent printing outcomes.^{[32, 36]} However, PEG’s biological properties for cellular functions are limited. As an alternative, natural hydrogels like GelMA are frequently used in light-based bioprinting to overcome these limitations.^{[36b, 37]} Miller and colleagues employed PEG-based bioinks with varying molecular weights and concentrations to create intricate multivascular networks and functional intravascular topologies. These structures were utilized for fluid mixing, valve formation, intervascular transport, nutrient delivery, and host engraftment.^[36b] Wangpraseurt et al. conducted a notable study where they developed 3D bionic corals capable of cultivating microalgae at high spatial cell density.^[38] They employed a light-based bioprinting technique using a bioink formulation comprising PEGDA, GelMA, glycidal methacrylate-HA (GM-HA), and PLA. This bioprinted coral structure accurately replicates the morphological characteristics of living coral tissue and the underlying skeleton at a micron-level resolution while mimicking their optical and mechanical properties. As a result, a programmable synthetic microenvironment was established.

2.3.3. Nanomaterial incorporation

Nanomaterials, including carbon-based materials, nanoclays, and graphene oxide, have been incorporated into hydrogel formulations to enhance bioink printability and functionality. Laponite nanosilicates, a type of synthetic clay with compositions akin to bioactive glasses, possess nanodisk-shaped structures, averaging 25 nm in diameter and 1 nm in thickness.^[39] Laponites exhibit exceptional dispersibility in aqueous solutions, distinct from natural clays.^[40] Laponites possess both negatively and positively charged surfaces and edges, respectively. At low concentrations (~1.5 wt%), they preferentially exfoliate due to their ionic properties. However, at higher concentrations (~3 wt%), they form a structure resembling a ‘house of cards’ as a result of interplatelet edge-to-face ionic interactions.^[41] In 3D bioprinting, Laponites have emerged as crucial components for modulating the rheological and mechanical properties of various polymer hydrogels. Acting as cross-linkers, they facilitate the formation of 3D networks with polymer chains, thereby influencing the overall structural integrity.^{[39c, 42]}

Similarly, silica nanoparticles serve as effective candidates for modulating the rheological properties of nanocomposite hydrogels. The silanol groups (-Si-OH) on silica surfaces allow facile modification with alkoxide-containing reactive compounds, such as aminopropyl triethoxysilane (APTES).^[43] Further modifications with various polymers are feasible to impart additional functionality.^[44] The rheological and mechanical properties of silica nanoparticle-incorporated hydrogels can be fine-tuned through physical interactions, including host-guest interactions^[45] and electrostatic interactions.^[46] These advancements contribute to the enhanced versatility and performance of nanomaterial-enhanced hydrogels in various biomedical applications, including 3D bioprinting.

2.3.4. Chemical modification

Hydrogel should be rapidly crosslinked after the printing process or during the printing process. Hydrogels are generally fabricated through cross-linked networks formed by physical interactions or chemical bonds.^[47] Physically crosslinked hydrogels are characterized by the network being held together with weak noncovalent interactions, including hydrogen-bonding, hydrophobic interaction, ionic interaction, host-guest interaction, metal-ligand interaction, and π-π stacking interaction.^[48] These reversible interactions allow most physical hydrogels to exhibit good printability due to dynamic rheological properties, which are advantageous for extrusion-based printing processes.^[49]

While physically cross-linkable hydrogels have many advantages for extrusion bioprinting, the mechanical stability of the printed hydrogel-based constructs is usually low, and this weak resistance to deformation frequently requires post-treatment for stabilizing 3D tissue constructs. In contrast, chemical hydrogels cross-linked with a permanent covalent bond are mechanically more robust, but the network is irreversible once the cross-links are formed.^[47b] These nondynamic characteristics of chemical networks limit their applications as bioinks for 3D bioprinting. Current fabrication approaches of 3D tissue construct via chemical gelation usually are based on the extrusion of mixed solutions containing two-component complementary reactive polymeric species or light/temperature-mediated polymerizable compounds.^[50] Because the chemical networks form during the extrusion mixing state, it is critical to consider the various factors that affect the gelation behavior, including the homogeneity in the mixed solution as well as the mixing rate.^[51] Excessively rapid cross-linking results in the clotting of the nozzles, whereas excessively late cross-linking induces spreading after extrusion and thus cannot produce self-standing 3D constructs.

Various chemical modifications have been implemented to incorporate functional groups into hydrogel materials, facilitating light-induced crosslinking and the creation of photosensitive hydrogels. This cross-linking method is valued for its rapidity, surpassing alternative chemical crosslinking mechanisms, thus making it an advantageous option for diverse 3D printing applications. The use of light in 3D bioprinting is well-established, particularly in the preparation of crosslinked hydrogel-based bioinks containing viable cells.^{[32, 52]} A wide array of photoactivated groups exists, which, upon exposure to light, can form radicals or undergo internal electronic activation, enabling subsequent reactions.^[53] Radical polymerization-based crosslinking is prevalent in 3D bioprinting due to the stability of radicals in aqueous physiological conditions and their compatibility with hydrophilic compounds, unlike ionic species. Additional light-induced crosslinking strategies include cycloaddition reactions such as Diels-Alder reactions or 1,3-dipolar cycloaddition. Photo-crosslinking can proceed through either a chain-growth or step-growth process, depending on the polymerization mechanism involved.^[54]

Gelatin possesses modifiable functional groups such as hydroxyl, amino, and carboxylic groups. Crosslinking occurs through photo-polymerization of reactive functional groups, primarily methacryloyl, immobilized onto the gelatin backbone. GelMA, which is the most common component for cell-based bioprinting, is consistently prepared by reacting gelatin with methacrylic anhydride (MAA), resulting in both methacrylate and methacrylamide groups due to reactivity towards MAA.^[55] The methacrylamide groups offer superior hydrolytic stability compared to methacrylates. Sequential modification of amino and carboxylic groups enhances gelatin’s mechanical properties; Van Hoorick et al. converted primary amines to methacrylamides, then activated carboxylic groups to introduce additional methacrylates using amino ethoxy methacrylate (AEMA).^[30] This approach results in improved crosslinking kinetics, reduced swelling ratios, and heightened stiffness. Gelatin derivatives have also been utilized in step-growth crosslinking approaches, with thiol-ene photoinitiated click chemistry being a common choice. These derivatives feature alkene functionality (e.g., norbornene, vinyl ester)^[56] for crosslinking with thiols introduced onto gelatin via reactions like allyl glycidyl ether (AGE) with primary amines.^[57]

HA-based hydrogels are widely used in tissue engineering due to their remarkable biological properties, making them suitable for various tissue engineering applications. HA has emerged as a prominent biomaterial in 3D bioprinting applications due to its versatility. It contains several functional groups (primary and secondary hydroxyl, carboxylic groups) within its structure, enabling the introduction of photoreactive moieties such as methacrylates, vinyl esters, and norbornenes. The primary method of chemical modification involves methacrylation of primary hydroxyl groups, typically achieved through reactions with MAA^[58] or glycidyl methacrylate.^[59] These modifications have led to the development of numerous HA-based bioinks, such as HA-vinyl ester^[60] and HA-norbornene,^[61] each tailored for specific applications and compatible with various cell types. This strategic approach enhances the functionality of HA-based hydrogels and expands their applicability in 3D bioprinting applications.

PEG is extensively utilized as a synthetic hydrophilic polymer in various biomedical applications, including surface modification, drug delivery, tissue engineering, and 3D bioprinting. Unlike natural polymers, PEG can be precisely synthesized with targeted molecular weight and architecture, such as linear or branched structures and defined chain end groups. While PEG-based hydrogels inherently lack biological activity due to the antifouling nature of PEG chains, they can be customized for specific applications by incorporating elements like RGD peptides for enhancing cell adhesion or MMP-cleavable peptides for facilitating proteolytic degradation.^[62] Among the PEG-based bioinks used in light-based 3D bioprinting methods are PEG-diacrylate and dimethacrylate, PEG-divinyl ester, PEG-dithiol, PEG-norbornene, and PEG-vinyl sulfone.^[52] These PEG-based bioinks provide tunable mechanical properties through variations in molecular weight, number of arms, and crosslinking mesh sizes, making them suitable for diverse tissue engineering and regenerative medicine applications. Moreover, numerous hydrogel materials have undergone chemical modifications to enhance their biological properties or enable photo-crosslinking, resulting in various applications utilizing 3D bioprinting methods.

2.3.5. Tissue-derived ECM bioinks

Tissue-specific ECM-derived bioinks have emerged as a promising approach in 3D bioprinting.^[63] These ECM materials can be obtained from cell-secreted ECMs in vitro culture or directly from native tissues through decellularization, effectively removing cellular components to prevent adverse immunological responses.^[64] To utilize tissue-derived ECM materials as bioinks, the ECM-rich materials are solubilized to reformulate into a gel type.^[65] The ECM provides a structural architecture with adhesion sites for cell surface receptors^[66] and preserves tissue function through its mechanical and biochemical properties.^[67] The interaction between cells and the surrounding ECM regulates various physiological processes, including motility, migration, invasion, and proliferation,^[68] and also modulates signal transduction pathways.^[69] ECM hydrogels contain structural and functional molecules found in native tissue ECM, such as collagen, laminin, fibronectin, growth factors, glycosaminoglycans, glycoproteins, and proteoglycans,^[70] enabling bioinks derived from decellularized tissue-specific ECM to mimic natural ECM functions.^[71] However, ECM-based hydrogels typically exhibit low viscosity, leading to challenges in shape fidelity and structural stability. Therefore, efforts have been made to enhance their chemical and physical properties for 3D bioprinting process.^[72]

Various tissue-specific extracellular matrix (ECM)-derived bioinks have been developed for 3D bioprinting applications. For renal tissue bioprinting, a photo-crosslinkable bioink (KdECMMA) derived from decellularized porcine kidneys was synthesized through methacrylation.^[73] Bioprinted human kidney cells within the KdECMMA bioink exhibited high viability and maturation over time, replicating the structural and functional characteristics of native renal tissue. Similarly, a photo-crosslinkable dECM bioink derived from decellularized porcine skeletal muscles, chemically modified by methacrylation,^[74] facilitated the production of cell-laden structures with unique topographical cues for skeletal muscle tissue engineering. Myoblasts within the printed dECM structures showed alignment and differentiation, with increased gene-expression levels compared to GelMA-based structures. Additionally, a cartilage-specific photo-crosslinkable bioink (cdECMMA) was developed from decellularized porcine auricular cartilage,^[75] supporting the viability, proliferation, and production of cartilage ECM components by printed auricular chondrocytes. These studies demonstrate the potential of ECM-derived bioinks for tissue-specific 3D bioprinting, providing tissue-specific microenvironments in the bioprinted tissue constructs.

2.3.6. Synthetic hydrogels

Synthetic hydrogels employed as bioinks offer low cytotoxicity and tunable mechanical properties. However, their limited biological interactions with cells pose a challenge. Pluronic F127, a thermo-sensitive hydrogel, exhibits a phase transition at room temperature, transforming into a viscous substance.^[76] At concentrations of 25% w/v or higher, it can be dispensed with high printability. Despite achieving high-resolution printed structures, Pluronic F127 structures may collapse in culture conditions. To enhance mechanical stability, Pluronic F127 is chemically modified into a photo-crosslinkable hydrogel.^[77] Moreover, it is commonly used as a sacrificial bioink to support 3D architecture, given its ease of uniform printing and immediate washout post-printing.^[7]

PEG-based hydrogels find extensive use in 3D bioprinting due to their adaptability for chemical modification and functionalization. These modifications, such as introducing functional motifs at the terminal end of PEG or combining with other hydrogels, enhance their biological and biomechanical properties.^{[37b, 78]} To further achieve the structural integrity of PEG-based constructs, the incorporation of diacrylate (DA) or methacrylate (MA) has been employed, offering versatility for various applications. Consequently, the photopolymerization of PEG-based hydrogels enables the achievement of tunable mechanical properties in bioprinted constructs.

2.4. Printability and bioink development

Printability refers to the performance of a bioink, specifically during the printing process.^[79] Developing bioinks requires rigorous printability assessment methods and a deep understanding of how rheological factors impact printing outcomes. Despite its significance, the evaluation of bioink printability is often described qualitatively in research reports, lacking reproducible or reliable quantitative measures. Although a few measures have been developed, such as structural integrity,^[24a] overhang collapse,^[80] and Pr value,^[81] each of these measures only captures a limited aspect of bioink behavior. Since bioinks must be capable of creating diverse structures, there is a pressing need for a comprehensive and bioink-specific methodology to evaluate printability and aid in advancing bioink development.

Rheological measurement has emerged as a widely adopted approach to examine the influence of different rheological factors on bioink printability. Mainly, it has proven successful in assessing the extrudability of materials, where shear-thinning models have reliably predicted the pressure-flow rate relationships of various bioinks.^[82] Within the rheological framework, several parameters, including viscosity, storage modulus, loss modulus, tan(δ), yield stress, and recovery capabilities, have been linked to the final printing outcomes.^{[24a, 80–82, 83]} However, the existing studies have been limited in their scope, typically focusing on a single rheological parameter using a specific model bioink, often at different concentrations, and with a single measurement of printing outcome. This limited approach may restrict the generalizability of the findings to other bioinks. Moreover, the interrelationships between different rheological measures have further complicated the matters. For instance, increasing the hydrogel concentration often leads to higher values of viscosity, storage modulus, loss modulus, yield stress, and improved shape fidelity.

We previously investigated the relationship between rheological properties and extrusion printing outcomes.^[84] A specialized bioink artifact specifically designed to improve the quantification of printability assessment was used.^[84b] This bioink artifact adhered to established criteria from extrusion-based bioprinting approaches. Rheological analysis revealed that the high-performing bioinks exhibited notable characteristics such as high storage modulus, low tan(δ), high shear-thinning capabilities, high yield stress, and fast, near-complete recovery abilities. Although rheological data alone cannot fully explain printing outcomes, certain metrics like storage modulus and tan(δ) correlated well (R²>0.9) with specific printing outcomes, such as gap-spanning capability and turn accuracy. The results highlight the importance of considering the holistic view of bioink’s rheological properties and directly measuring printing outcomes. These findings emphasize the need to enhance bioink availability and establish standardized methods for assessing printability.

Moreover, it is essential to note that the bioinks used in practical applications are typically printed together with cell suspensions. Although this study did not include cell suspensions, we have previously investigated the impact of cell density on printing outcomes.^[85] Interestingly, our findings revealed no significant differences in printing outcomes for cell densities ranging from 0 to 40 × 10⁶ cells/mL. However, slight variations in rheological properties were observed with different cell densities, indicating a potential influence on printability. A standardized method to measure bioink printability is crucial for the FDA regulatory process, commercialization, and clinical applications.

3. Applications: 3D Bioprinted Transplantable Tissue Constructs

Tissue engineering aims to reconstruct damaged tissues or organs and restore their functionality by creating biomaterial scaffolds with cells in specific shapes suitable for implantation into patients. While various manufacturing methods have been developed to achieve this objective, conventional approaches exhibit limitations in producing complex, sophisticated tissue constructs suitable for clinical applications. The 3D bioprinting technology has substantially addressed these limitations. With the foundation of 3D bioprinting technology, it becomes feasible to reconstruct the intricate shapes of complex tissues. Furthermore, this technology enables the fabrication of composite tissues, establishing junctions between different tissues.

3.1. Design strategies

3D bioprinting strategies aim to create tissue constructs suitable for clinical applications, with the potential to engineer various tissue types. One of the primary considerations in manufacturing tissues or organs is the necessity for a design strategy that accurately mimics the structure and function of the target tissue or organ. This requires a deep understanding of the anatomical and functional aspects of the tissue or organ being replicated. The advantage of 3D bioprinting lies in its ability to effectively implement such a design strategy, allowing for the creation of precise tissue or organ models that closely resemble their natural equivalents.

Design strategies encompass various tissue types, including shape-based tissues like bone,^[86] cartilage,^[7] skin,^[87] and cornea;^[88] hollow structures such as blood vessels,^[89] urethras,^[90] and tracheas;^[91] organized tissues like skeletal muscle,^{[7, 35b, 92]} cardiac muscle,^[35a] and neural tissue; composite tissues like osteochondral (bone-cartilage) and musculotendinous (muscle-tendon)^[93] tissues; and whole organs like the kidney, liver, and heart (Figure 3). These organs necessitate intricate micro-vasculature and functional inner structures^{[76, 94]} to function effectively. Through the incorporation of biomaterials, cells, and biochemical and biophysical cues, 3D bioprinting offers the opportunity to reconstruct the structural and functional complexity of human tissues. This approach allows for the precise design of tissue shape, organization, structure, and integration, paving the way for advancements in regenerative medicine and tissue engineering.

Figure 3.

The 3D bioprinting technology enables the creation of constructs in various shapes and sizes, including human-scale bone,^[86] ear cartilage,^[7] (Copyright 2016, Nature) and multi-layered skin,^[87] (Copyright 2023, The Authors, published by AAAS) as well as hollow tubular structures such as trachea,^[91b] (Copyright 2019, IOP) urethra,^[90] (Copyright 2017, Elsevier) and blood vessel.^[89] (Copyright 2020, IOP) At the tissue organization level, precise cellular alignment is achievable, particularly in the construction of skeletal tissue^[7] (Copyright 2016, Nature) and cardiac muscle.^[35a] (Copyright 2018, Elsevier) Further innovation extends to the fabrication of composite tissues, including osteochondral (bone-cartilage) and musculotendinous (muscle-tendon) structures.^[93] Copyright 2015, IOP. As the field progresses, the integration of functional inner structures, including microvasculature^[76] (Copyright 2011, Wiley) and nephrons,^[94] Reproduced under the terms of the CC-BY license (Copyright 2016, Nature) is becoming increasingly essential to meet the demands of whole organ bioprinting.

3.1.1. Shapes

Clinically relevant simple cellular constructs of specific size, shape, and structural integrity have been successfully produced. For example, bioengineered skin constructs, comprising both epidermal and dermal layers, have been created by bioprinting layers of fibroblasts- and keratinocytes-laden hydrogels.^[95] Using jetting bioprinting, a 3D human skin construct was successfully fabricated, demonstrating the feasibility of bioprinting human skin with biological relevance.^[96] Additionally, in situ 3D bioprinting techniques have been utilized for treating large-scale skin wounds and burns in rodent and porcine models.^[97] Results showed that in situ skin bioprinting is able to deliver tissue elements directly onto the wound to achieve uniform coverage using cell-laden hydrogels. A recent study employed six primary human skin cell types to bioprint a trilayer skin construct comprising the epidermis, dermis, and hypodermis.^[87] Transplantation of this bioprinted skin, containing human cells, onto full-thickness wounds in mice resulted in rapid vascularization and the formation of epidermal rete ridges resembling those found in native human epidermis, accompanied by the presence of normal-looking ECMs. Numerous promising outcomes suggest the potential applicability of skin bioprinting technology for clinical use in humans.

Conventional additive manufacturing methods, like the fabrication of personalized metal implants, are already employed in clinical settings.^[98] 3D bioprinting presents a distinct and promising alternative, particularly for bone grafting, due to its ability to accommodate diverse anatomical variations, defect sizes, and patient-specific morphologies in bone pathologies.^[99] Medical imaging techniques such as CT and MRI enable the creation of personalized bone constructs using osteoconductive materials like hydroxyapatite and β-tricalcium phosphate (TCP), along with osteogenic cell types.^{[7, 100]} A recent study explored the efficacy of a specific construct design using geometric control.^[86] Improved bone regeneration was observed by incorporating a dense external layer (50-μm pore size) and a highly porous core structure (700-μm pore size). These findings highlight the potential of 3D-printed bone constructs to enhance bone regeneration by minimizing competition for fibrotic tissue formation within bony defects. This study underscores the ability to translate patient-specific anatomy into tailored 3D bioprinting strategies using medical imaging, thereby producing clinically relevant constructs.

Bioprinting strategies offer an appealing approach for fabricating cartilage tissue constructs, particularly in customizing the size and shape of individual lesions to match patient-specific needs.^[101] Bioprinted tissues exhibit biomechanical and biochemical properties similar to native cartilage, resulting in robust integration with surrounding tissue.^[102] In particular, bioprinting has achieved remarkable success in engineering cartilage for the external ear, leading to the development of a bionic ear capable of converting sound waves into electrical signals.^[103] This bioprinted ear cartilage construct comprised sodium alginate, silver nanoparticles, and chondrocytes arranged in an ear-shaped configuration. The extrusion-based bioprinting system has also been applied to fabricate a human-sized ear cartilage tissue construct.^[7] Upon implantation, the printed ear maintains its shape, facilitating cartilage tissue formation, as confirmed by staining for glycosaminoglycans (GAG) and collagen type II. The next frontier in bioprinting for cartilage regeneration lies in conducting translational studies.

3.1.2. Hollow structure

This cutting-edge approach enables the engineering of patient-specific tissue structures, encompassing hollow formations like the blood vessels, urethras, and tracheas. 3D bioprinting strategies were applied to fabricate bioengineered hollow tissue constructs.^[90] Skardal et al. demonstrated the bioprinting of a tubular vascular construct using a scaffold-based approach, showing cell proliferation and matrix remodeling in vitro.^[104] Similarly, Bertassoni et al. utilized cell-laden GelMA bioinks to fabricate vascular constructs.^{[104a, 105]} In contrast, scaffold-free constructs have been achieved through 3D bioprinting. For instance, Norotte et al. created a basic blood vessel, although the burst pressures fell below physiological ranges.^[106] Additionally, another study utilized extrusion bioprinting to craft a vessel-like structure with multiple fluidic channels at different levels.^[107] This involved extruding partially cross-linked hollow alginate filaments loaded with fibroblasts and smooth muscle cells through a coaxial nozzle and deposition along a rotating rod template. Vascular endothelial cells (ECs) were then seeded onto the inner wall, forming two-level fluidic channels due to fusion between adjacent hollow filaments.

Trachea reconstruction using tissue-engineered tubular constructs has made progress, yet the tubular configuration presents challenges such as mechanical rigidity and a lack of vascularization in the connective tissues between cartilages and epithelium. The native trachea’s structure, characterized by alternating cartilage and vascularized fibrous tissue rings, is crucial for its mechanical and physiological functions. To replicate this, 3D bioprinting strategies have been utilized. Different design strategies for trachea bioprinting have been explored.^{[91, 108]} For instance, a study successfully bioprinted tubular structures mimicking the native trachea’s alternating cartilage and vascularized fibrous tissue rings.^[109] These constructs achieved tight integration through enhanced interfacial bonding. This approach resulted in a functional trachea reconstruction with mechanical and physiological properties closely resembling those of the native trachea, due to its alternating stiff-to-soft tissue structure.

In urethra tissue engineering, a spiral construct design was created using a combination of poly(ε-caprolactone) (PCL) and poly(lactide-co-caprolactone) (PLCL) polymers through extrusion printing.^[90] This method aimed to mimic the structural and mechanical properties of native tissue. Subsequently, the scaffold was combined with a cell-laden fibrin-based bioink, resulting in the formation of a tubular construct comprising urothelial cells (UCs) and smooth muscle cells (SMCs). This innovative approach has the potential to be extended to the creation of other tubular structures, such as the ureter and vaginal tissue.

3.1.3. Tissue organization

Extrusion-based bioprinting has the capability to promote cellular orientation at the microscale level through unidirectional printing, a crucial feature necessary for the development of contractile tissues such as skeletal and cardiac muscles. Skeletal muscle, constituting about 40% of human body weight,^[110] is characterized by its highly aligned muscle fibers, critical for effective contraction and force generation.^[111] Leveraging the spatial organization capabilities of bioprinting, skeletal muscle constructs composed of well-oriented muscle-like bundles have successfully been engineered.^{[7, 74, 92–93]} These bioengineered constructs have demonstrated maturation into functional tissue, resulting in integration with host nerves when implanted in vivo.

Similarly, cardiac tissues exhibit a complex myocardial organization essential for contractility. Employing 3D bioprinting, a cardiac patch was created with spatially organized patterns of human mesenchymal stromal cells (hMSCs) and ECs on poly(ester urethane urea) (PEUU) substrate.^[112] Implantation of this bioprinted cardiac patch onto infarcted hearts in rats led to enhanced vascularization and improved cardiac function. Furthermore, human cardiac-derived cardiomyocyte progenitor cells (hCMPCs) were bioprinted, demonstrating high cell viability and cardiac tissue maturation.^[113] Additionally, a bioprinted cardiac tissue using neonatal rat ventricular cardiomyocytes was examined for drug response, including epinephrine and carbachol.^[35a] The bioprinted cardiac tissue showed a typical response to cardiac drugs. The classical androgen agonist epinephrine increased the beating frequency from 80 to 110 BPM, while the androgen antagonist carbachol decreased the beating frequency to 40 BPM. These effects were reversible when the drugs were removed from bioprinted cardiac tissues, indicating that bioprinted cardiac tissues could physiologically respond to cardiac drugs.

Human induced pluripotent stem cell (hiPSC)-derived cardiomyocytes represent a promising cell source for future clinical trials. These cardiomyocytes, when combined with SMCs and ECs within a well-organized ECM scaffold, have been utilized to fabricate hiPSC-derived cardiac muscle patches.^[114] In another study, a half-heart structure containing primary feline adult and H1 cardiomyocytes in alginate hydrogel was bioprinted using a modified jetting bioprinting method.^[115] This construct exhibited a porous structure, with the deposited cells retaining viability.

3.1.4. Composite tissues

While initial successes have been made in creating simple-shaped tissue constructs, there is a growing need for methods capable of generating more intricate, composite tissue constructs.^[116] To address this, combination approaches are being employed to create multi-layered architectures representing different phases, such as bone and cartilage, in osteochondral constructs. Demonstrating this capability, one study has successfully demonstrated complex osteochondral tissue constructs using bioprinting techniques.^[117] These constructs, anatomically accurate and composed of a composite of PCL and hydroxyapatite, have shown promise in regenerating entire articular surfaces of synovial joints in animal models, highlighting the potential for functional tissue reconstruction. In another study, bioprinted layers of articular cartilage and calcified cartilage were created using a combination of hMSCs and human articular chondrocytes. This was achieved alongside photo-crosslinkable hydrogel gradients, with the underlying calcified cartilage induced using TCP microparticles to generate a bi-phasic construct.^[118] In vivo analysis in rodent osteochondral defects revealed the formation of repaired articular cartilage, rich in tenascin and collagen type II, within 12 weeks.

Indeed, 3D bioprinting proves particularly advantageous in fabricating composite tissues like musculotendinous tissue. Four different tissue components were bioprinted to fabricate a single integrated muscle-tendon unit (MTU) construct.^[93] This MTU construct featured mechanically heterogeneous polymeric materials, exhibiting elasticity on the muscle side and stiffness on the tendon side. Furthermore, the construct displayed a tissue-specific distribution of cells, with myoblasts situated on the muscle side and fibroblasts on the tendon side. Results indicated high cell viability and cellular orientation, alongside an increase in musculotendinous junctional gene expression. This emphasizes the capability of 3D bioprinting to fabricate complex, region-specific tissues with both biological and biomechanical characteristics tailored to their intended function.

3.1.5. Solid organ bioprinting

The bioengineering of solid organs like the heart, liver, and kidney poses significant challenges due to their intricate complexity, which involves integrating complex vascular networks and functional units comprising multiple cell types.^[119] The challenge in solid organ printing lies in replicating the complex structural hierarchy of organs, which encompasses macrostructural, mesostructured, microstructural, and nanostructural levels.^[119a] For instance, the kidney consists of macrostructural elements such as the renal artery and vein, mesostructured elements like the renal pelvis and interlobular blood vessels, microstructural units including the nephron, and nanostructural components like the ECMs and diffusion channels. Each level exhibits unique features crucial for organ function and requires precise replication for successful tissue engineering.

Bioprinting offers a promising approach to address these complexities by enabling the creation of microchannels capable of housing layers of ECs.^[76] It is well-established that the limit of oxygen and nutrient diffusion for cells to survive in vivo is within the range of 100 – 200 μm.^[120] Despite numerous efforts to incorporate vasculature within 3D tissue constructs, this remains a significant technical challenge. Consequently, 3D bioprinting has emerged as a solution to create microchannels with EC coverage. One of the earliest approaches involved using sacrificial materials during printing, which served a structural role and were subsequently removed to form microchannels. However, methods such as vascular casting using carbohydrate glass as the sacrificial template have limitations in size and culture duration due to practical difficulties in direct perfusion.^[121]

Several advanced methods have been developed to print 3D tissue constructs prefabricated with vasculature, utilizing multiple cell types and ECM proteins.^[122] Human microvascular ECs, combined with printing biomaterials, have demonstrated self-alignment inside the microchannels, forming a confluent microvascular lining. The embedded bioprinting technique was employed to reconstruct various components of the human heart, including small capillaries and the entire organ, with unprecedented filament resolution using pH-driven gelation.^[123] This approach facilitated printing a porous microstructure, allowing rapid microvascularization and cell infiltration. Despite advancements in 3D bioprinting allowing for the creation of complex vascularized tissue constructs, reconnecting the vasculature to the host circulatory system remains challenging.^[124]

There are a few studies focused on bioprinting functional inner structures, such as renal tubules and hepatic lobules. A study reported the generation of 3D human renal proximal tubular structures containing proximal tubular epithelial cells.^[94] The tubule-like structure was circumscribed by proximal tubule epithelial cells and actively perfused through the open lumen, and the bioprinted epithelial barrier was disrupted with the introduction of nephrotoxin, Cycolosporine A. The perfusable 3D proximal tubules facilitated the development of a tissue-like epithelium with enhanced phenotypic and functional characteristics compared to cells cultured on 2D surfaces. Another study has demonstrated the preservation of hepatocellular function within bioprintable hydrogels. These hepatic tissue constructs were created using a photopatterning platform to embed cells within hydrogels, featuring multilayered, intricate architecture and supporting primary human hepatocytes through paracrine and juxtacrine signaling within the scaffold.^[125] These constructs exhibited sustained humanized liver function upon transplantation into mice, including human protein synthesis, human drug metabolism, drug-drug interactions, and drug-induced liver injury.^[126] Additionally, efforts to simplify complex fabrication methods and explore new bioprinting technologies for vasculature and functional inner structures are crucial for the fabrication of solid and complex organs.

3.2. Regulatory consideration

Additive manufacturing, including 3D bioprinting, is emerging as a valuable tool for printing biological products.^[127] According to a 2022 report, the global 3D bioprinting market is valued at $1.3 billion in 2022 and is projected to reach $3.3 billion by 2027, growing at a CAGR of 20.8% from 2022 to 2017.^[128] However, despite ongoing interactions between stakeholders and the U.S. Food and Drug Administration (FDA), biological products manufactured using additive manufacturing have yet to receive FDA approval or clearance.^[129] The FDA’s Center for Devices and Radiological Health (CDRH) oversees the regulation of firms engaged in medical device production, including manufacturing, repackaging, relabeling, and importation within the United States. This regulatory authority extends to devices produced using 3D bioprinting technology, which must adhere to the same regulatory standards as those manufactured using traditional methods.^[130] The differences in manufacturing techniques between 3D printing and conventional methods introduce distinct technical considerations that the FDA must address during its scientific evaluations. Medical products created through additive manufacturing undergo regulatory review and oversight via established pathways such as premarket notification [510(k)] and the new drug application (NDA) processes in the U.S. As more intricate additive manufacturing products, including biologics, become prevalent, additional regulatory pathways such as premarket approval (PMA) may also be utilized to ensure compliance and safety standards are met.^[129]

The FDA has published guidance documents and Quality Systems regulations to assist manufacturers, providing valuable information on regulatory requirements and best practices. The FDA’s Technical Considerations for Additive Manufactured Medical Devices document notably offers guidance specific to devices produced using additive manufacturing techniques, including 3D printing. However, it is essential to note that while this guidance addresses technical considerations for additive manufacturing, it does not explicitly cover incorporating biological, cellular, or tissue-engineered products into additive manufacturing processes.^[131] Therefore, additive manufacturing of cellular structures requires an established regulatory framework to ensure the safety, consistency, and long-term evaluation of individually printed products. These regulations must be distinct from those governing good manufacturing practice (GMP) facilities for cells and processed tissue constructs. As such, further regulatory clarity and guidance may be needed to ensure the safe and effective use of additive manufacturing for producing biological products.^[132]

3.3. Clinical bioprinting workflow

The clinical bioprinting workflow is a detailed process to leverage 3D bioprinting to create functional tissue constructs suitable for clinical use.^[133] Its primary goal is to generate reproducible and intricately structured tissue constructs that closely resemble native anatomy, facilitating their future clinical applications.^{[7, 134]}

3.3.1. Medical imaging and 3D CAD modeling

CAD/CAM processes are essential for advancing the clinical applications of 3D bioprinting, as they enable the automated replication of intricate tissue structures in three dimensions.^[134a] Typically, this process commences with patient scanning, utilizing medical imaging modalities to generate 3D volumetric data of the target object. These imaging tools capture cross-sectional slices of the body, which are then stored in the Digital Imaging and Communications in Medicine (DICOM) format, widely recognized as the standard for medical digital imaging. Subsequently, this data transforms a CAD model through the reverse engineering process.

The process is initiated by enhancing resolution and generating voxels from the measured data through interpolation of points within and between image slices. Subsequently, a CAD model is constructed by extracting localized volumetric data to form a surface model of the targeted tissue structure, requiring sophisticated reconstruction due to its complexity. Following this, a motion program is generated using a CAM system, which involves slicing, tool path generation, and motion program generation. Slicing consists in obtaining information on the sliced 2D shapes of an object for layer-by-layer printing. Tool path generation determines the path for the tool to follow in filling the cross-sectional space of each layer. Given the importance of accurately replicating tissue-specific architecture, a meticulously organized strategy for tool path generation is vital to ensure the proper inner functional structure, comprised of multiple cellular components, facilitating efficient tissue regeneration.

There is a pressing need for user-friendly and specialized software solutions for 3D bioprinting applications. While various commercial and open-source programs exist, they often lack the intuitive interface and streamlined workflows necessary for the efficient translation of medical imaging data into printable constructs. This inefficiency can lead to delays in surgical procedures and patient care. To address these challenges, it is crucial to develop software packages that are tailored to the specific requirements of bioprinting applications. These packages should offer intuitive interfaces and streamlined workflows to simplify the process of generating printable models from medical imaging data.

Furthermore, incorporating artificial intelligence (AI) into the bioprinting workflow offers significant value by automating intricate tasks and refining design strategies.^[135] AI-driven software has the potential to analyze medical imaging data comprehensively, facilitating the creation of precise and personalized models that account for unique patient anatomy and clinical needs. This augmentation not only enhances the efficiency of bioprinting processes but also elevates the functionality and effectiveness of the resulting constructs. By harnessing advanced software solutions, workflows can be streamlined, efficiencies enhanced, and innovative opportunities unlocked, thereby advancing personalized medicine and tissue regeneration.

3.3.2. Bioprinting process and operation

The 3D bioprinting workflow strategy from the medical image to the printed tissue constructs developed by the CAD/CAM process and automated printing of 3D shapes imitating target tissues or organs.^{[7, 134a]} This workflow consists of several essential components, each contributing to successfully fabricating complex tissue constructs for transplantation or therapeutic purposes (Figure 4). Firstly, medical imaging techniques such as MRI or CT scans are employed to obtain precise anatomical data of the target tissue or organ. This data is then processed using specialized software to generate accurate 3D CAD/CAM models, serving as blueprints for the bioprinting process. The software also generates printing code to guide the bioprinter in depositing bioink and cells layer by layer to create the desired tissue structure. Bioink selection is a crucial step for supporting cell growth, providing structural integrity, and mimicking the native tissue’s ECM. Construct design strategy involves planning the tissue’s architecture and organization to ensure optimal functionality and integration with surrounding tissues post-implantation. This includes considerations such as vascularization, cell-cell interactions, and mechanical properties. Cell manipulation techniques are utilized to prepare cells for bioprinting, including isolation, expansion, and encapsulation within the bioink.

Figure 4.

Clinical bioprinting workflow for manufacturing patient-specific tissue constructs.

The bioprinting process entails precise deposition of bioink and cells according to the CAD/CAM models, layer-by-layer, using specialized equipment. Quality control assessments follow printing to confirm structural integrity, cellular viability, and functionality, involving techniques like histological staining, immunohistochemistry, and functional assays. Feedback from operators and patients is collected to assess the effectiveness of the bioprinted tissue constructs and identify areas for workflow improvement. This iterative refinement process is crucial for advancing clinical bioprinting in regenerative medicine and tissue engineering.

3.4. Clinical trials

In 2022, two Ukrainian patients suffered severe skull injuries due to regional circumstances. Dr. Warren Dorlac stepped forward to undertake the skull bone reconstruction, seeking assistance from the Wake Forest Institute of Regenerative Medicine (WFIRM).^[136] To support the production of 3D-printed patient-specific implants, Dr. Dorlac collaborated with T&R Biofab, a company based in South Korea known for its advanced 3D manufacturing capabilities and a partner of WFIRM. The collaboration led to the successful development and delivery of the implants within a remarkable 10-day timeframe, enabling prompt surgical intervention in Poland. Follow-up CT scans conducted over the course of a year revealed that the transplanted artificial tissues remained well-maintained, with no reported adverse effects. Today, the use of 3D printing technology in real clinical settings has become a practical reality (Table 1).

Table 1.

Clinical trials utilizing 3D printed patient-specific constructs.

Title	PI	Country	Method	Cell	Application	Status	ID #	Sponsor
A dermo-epidermal autologous skin substitute for further therapeutic use (Biopskin)	Dominique Casanova, MD, PhD	France	Unknown	Keratinocytes and dermal fibroblasts	Patient-derived dermo-epidermal substitute	Unknown	NCT04925323	Assistance Publique Hopitaux De Marseille
Patient-customized bioprinting technology for practical regeneration of the respiratory tract	Ja Seong Bae, MD, PhD	South Korea	Extrusion	Stem cells derived from the nasal cavity and septum cartilage	PCL bellow scaffold with decellularized ECM	Active/Not recruiting/Phase 2	NCT06051747	Seoul St. Mary’s Hospital
AuriNovo for auricular reconstruction	John Reinisch, MD	USA	Extrusion	Chondrocytes	3D printed auricle with collagen hydrogel scaffold for microtia	Terminated/Phase 2	NCT04399239	3D Bio Therapeutics
Evaluation of using 3D printed PEEK facial implants in repairing maxillofacial deformities	Hekmat Yacoub, PhD	Syria	Extrusion	None	PEEK of 3D printed facial implant for maxillofacial deformities patients	Completed	NCT05348434	Tishreen University
Regenerative medicine approach to nasal reconstruction	Brittany Howard, MD	USA	Unknown	Diced cartilage autologous graft	3D printed nasal tip for nasal reconstruction	Recruiting	NCT05273060	Mayo Clinic
Clinical application of personal designed 3D printing implants in bone defect restoration	Fang Guofang, MD	China	Powder bed fusion	None	Laser beam melting printing of titanium implant for bone tissue grafting	Unknown	NCT03166917	Shenzhen Hospital of Southern Medical University
Safety study of 3D printing personalized biodegradable implant for breast reconstruction	Ju Liang Zhang, MD	China	Extrusion	None	3D printing of PCL implants based on MR imaging data for breast cancer patients	Recruiting/ Phase 1	NCT03348293	Xijing Hospital
A clinical study to evaluate the safety and efficacy of surgical material for nasal septoplasty	Sung Won Kim, MD Se-Hwan Hwang, MD	South Korea	Extrusion	None	3D printed PCL mesh for nasal reconstruction.	Recruiting	NCT03088618	T&R Biofab Co., Ltd.
Mitigation of tracheobronchomalacia with 3D-printed personalized medical devices in pediatric patients[137]	Glenn E. Green, MD	USA	Extrusion	None	3D printed PCL/hydroxyapatite (96/4) blended construct as external airway splints	3 patients/completed	Institutional Review Board (IRB) and the FDA under the medical device emergency use exemption	University of Michigan

Although clinical trials using 3D printing technologies have increased, the application of cell-based bioprinted tissue products remains limited. Scaffold-only products, which do not contain cells, generally require only device clearance or approval. In contrast, cell-based products often require biologics or combination product approval, which involves more extensive testing and regulatory review. Additionally, as previously mentioned, the regulatory framework for 3D printing technologies is not yet fully established. Therefore, this section addresses both cell-based bioprinting and scaffold-only 3D printing in clinical applications.

As of January 2024, there is significant exploration of 3D printing technologies in clinical trials, with 240 studies related to 3D printing listed on ClinicalTrials.gov, managed by the National Library of Medicine. One key advantage of 3D printing technology is its capacity to fabricate structures tailored to individual patients based on their 3D clinical data. This capability allows for the creation of personalized structures designed to meet the unique needs of each patient. In particular, biomaterials capable of supporting tissue transplantation are being employed in 3D printing to reconstruct surgical sites in a patient-specific manner, with various products and technologies undergoing clinical trials. For instance, there are endeavors to develop acellular porous titanium bone structures using laser beam printing for bone defect reconstruction, as well as the utilization of biodegradable PCL implants for partial or total mastectomy reconstruction in breast cancer patients over two years. Polyetheretherketone (PEEK) printing is also being investigated to address maxillofacial deformities in patients.

While cell-based bioprinting applications are still limited in clinical trials, the use of 3D printing technologies with patient-derived cells holds significant potential for advancing personalized medicine. Autologous cells, sourced directly from the patient, could improve immunocompatibility, increase safety, and enhance tissue regeneration and engraftment rates.

4. Applications: 3D Bioprinted In Vitro Tissue Models

The drug discovery process is increasingly burdened by rising costs and delays, despite technological advancements. Eroom’s Law exemplifies this trend, revealing that the inflation-adjusted cost of developing a new drug nearly doubles every nine years.^[138] Adding to this challenge is the high failure rate of drugs beyond pre-clinical and clinical trials.^[139] While 12% of drugs show promise in pre-clinical animal studies, 88.3% fail during clinical trials, mainly due to drug toxicity or lack of efficacy in humans. This raises questions about the necessity of pre-clinical animal testing. If animal toxicity studies could more accurately predict human responses, harmful drugs might be eliminated earlier, potentially boosting success rates to 56%. Recognizing this, the Senate passed the FDA Modernization Act 2.0 (https://www.congress.gov/bill/117th-congress/senate-bill/5002), which allows alternative methods, such as cell-based assays and computational models, for evaluating drug safety and efficacy. This legislation marks a significant shift, emphasizing the need for in vitro tissue and organ models using human cells over traditional animal-based studies.

In vitro tissue models have been designed to replicate the biological, structural, or physiological functions of tissues and organs by integrating human cells and ECM components. These biomimetic platforms have diverse applications, such as drug testing, toxicity assessment, and disease modeling, and hold promise for revolutionizing drug discovery processes. Traditional pharmaceutical approaches face challenges like high costs, lengthy timelines, and limited predictability of preclinical animal testing. In vitro tissue modeling technology seeks to overcome these obstacles by offering more accurate and efficient platforms for drug development.

In conventional in vitro tissue modeling, cells are typically cultured under 2D conditions, such as in tissue culture plates, which represent a basic approach. However, certain cell types may lose their natural characteristics and functionality under such conditions. To address this issue and provide cells with a more lifelike environment, alternative methods are employed (Figure 5). Advanced techniques include culturing cells as aggregates (spheroids) or organoids, utilizing 3D bioprinting to create intricate tissue/organ structures, and integrating microfluidics platforms to support dynamic microenvironments. While both 2D and 3D cell culture systems offer high-throughput capabilities for screening, they may fall short in mimicking the complexity and function of native tissues. In contrast, microfluidics-based tissue chips excel in replicating tissue intricacies but may need to be more suitable for high-throughput applications.

Figure 5.

In vitro tissue modeling systems, including 2D cell culture, 3D organoids, bioprinted tissue constructs, and microfluidics-driven tissue models.

The utilization of 3D bioprinting in in vitro model systems is essential for recreating the intricate structures and functions of tissue and organs, supporting predictive clinical models for personalized therapies, and enabling high-throughput drug screening and disease modeling with clinical relevance.^[140] While traditional fabrication methods for in vitro tissue models involve lithography and molding processes with materials like polydimethylsiloxane (PDMS), PCL, and glass, these methods are often labor-intensive and hinder uniform production. In contrast, 3D bioprinting offers automated manufacturing systems and precise spatial control, overcoming these challenges and advancing the development of in vitro tissue modeling technology.^[140–141]

Given the inherent complexity of tissues and organs, characterized by diverse cell types and specific geometric arrangements, 3D bioprinting is essential for improving the accuracy of in vitro models. To achieve tissue models that closely replicate native functions and phenotypes, they must undergo a maturation process. This involves dynamic culture conditions, such as perfusion environments, mechanical and electrical stimulation, and co-culture of multiple cell types. Figure 6 illustrates a functional classification of various tissue models, including neurophysiological, muscle function, metabolic disease, skeletal and hematopoietic, infection, and multi-organ interaction systems. Current research focuses on refining these models to better address specific tissue characteristics and applications. Advances in technology are driving progress in 3D bioprinted in vitro tissue systems, tailored to a range of tissue and organ models (Table 2).

Figure 6.

Bioprinted in vitro tissue model systems that replicate tissue or organ-specific biological, biochemical, and biomechanical functions and organ-to-organ interactions for drug discovery and precision medicine.

Table 2.

Representative in vitro tissue models fabricated by 3D bioprinting technologies.

Organ	Design concept	Method	Bioinks (2D/3D)	Cell types	Frame	Outcomes
Adipose tissue	Pathophysiology in obese adipose tissue with in-bath printing[142]	Extrusion	Alginate, dECM (3D)	Preadipocyte	PEVA	Recapitulation of inflammation and insulin resistance
Bone	Spatiotemporal mimics of osteon in well plate for high throughput screening[143]	Light	dECM, collagen (3D)	Mouse osteocyte like-cells, mouse pre-osteoblasts	PDMS	Horizontal bone tissue maturation and osteoporosis drug testing of deep learning
Neural tissue	Oxygen-gradient chip with patient-derived tumor and brain-derived ECM for patient-specific response[144]	Extrusion	dECM (3D)	Patient-derived primary glioblastoma, ECs	PDMS, glass cover	Oxygen gradient control in brain cancer incubation increased patient-specific chemotherapy response
	Spatially-controlled vascularized cerebral organoid with perfusion[145]	Light	Matrigel (3D)	hiPSC-derived vascular cells and cerebral organoids	Dental SG resin	Dextran (40 kDa) and red fluorescence bead (1 μm) perfusable vascularized brain organoid
Cardiac tissue	One-step fabrication of physio-mimetic laminar cardiac tissue with sensor by multi-material printing[146]	Extrusion	Fibronectin, Matrigel (3D)	hiPSC-cardiomyocytes, neonatal rat ventricular myocytes	PDMS, Ag:Pa ink versamid 973 PU	Multilayer cantilevers, strain sensor-embedded anisotropic laminar cardiomyocyte tissues
Hepatic tissue	Oxygen-gradient hepatic metabolic zonation with differentiated hepatocytes[147]	Light	Matrigel (2D)	hESCs	PDMS	Oxygen gradient liver zonation and zonation specific functionalities
Intestinal tissue	Spontaneous intestinal morphogenesis and 3D epithelial structure[148]	Light	Matrigel, rat-tail collagen (3D)	Caco-2, human colonic organoids	PDMS. Polyester membrane	Spontaneous intestinal morphogenesis and 3D epithelial structure with crypt–villus characteristics
Renal tissue	Compartmentalization of proximal tubule and endothelial channel[149]	Extrusion	Pluronic, alginate, collagen (3D)	Primary renal proximal tubule epithelial cells and glomerular ECs	PEVA	Drug-induced nephrotoxicity modeling
Lung	Alveolar epithelium chip for radiation-induced lung injury for radiotherapy[150]	Light	Collagen, laminin (2D)	lung microvascular ECs, alveolar epithelial cells	PDMS	Human physiology recapitulation of radiation therapy and evaluation of radiation countermeasure drugs effect
Skin	HSV infection of immuno-competent skin model with endothelial network[151]	Light/injection modeling	Collagen (3D)	Primary dermal fibroblasts, primary epidermal keratinocytes, dermal microvascular ECs	PDMS, plexiglass	Host immune response and antiviral treatment of vascularized skin model
Pancreatic tissue	Glucose metabolism for type 2 diabetes by different adipose tissue origin ECM[152]	Extrusion	dECM (3D)	Adipocyte, HepG2, THP-1, HUVEC, human primary beta cells	PCL, PDMS cover	Subcutaneous and visceral adipose ECM induced differential glucose metabolism

ECM: extracellular matrix; dECM: ECM derived decellularized tissue; PEVA: poly(ethylene-co-vinyl acetate); PDMS: polydimethylsiloxane; PU: polyurethane; hESCs: human embryonic stem cells; PCL: poly(ε-caprolactone).

4.1. In vitro adipose tissue models

Adipose tissue, constituting around 20–25% of total body mass, ranks as the second largest organ, following the skin. It consists of adipocytes and stromal vascular fraction (SVFs), including ECs, pericytes, fibroblasts, and MSCs, and serves as an endocrine regulator of energy homeostasis and glucose metabolism through the secretion of adipokines.^[153] To replicate an in vivo environment, adipocytes play a crucial role as primary cell components in adipose tissue. However, the challenge lies in the limited size of lipid-filled adipocytes, typically ranging from 10 to 30 μm in in vitro culture. To address this limitation and achieve sizes comparable to those found in lean and obese adults, micropatterning technology and 3D spheroid culture techniques emerged, offering enhanced physiological relevance and resembling in vivo patterns of mature adipocytes.^[154] Additionally, a scaffold-free 3D adipocyte culture platform was utilized to differentiate adipocytes, demonstrating enhanced physiological conditions.^[155] Culturing adipocytes in 3D spheroids showed promising results, with mature adipocytes exhibiting patterns resembling those found in vivo, as evidenced by multi-omics profiling of over 1000 lipid species.

Despite these advancements, the in vitro culture of adipocytes still faces challenges primarily associated with the 3D fabrication process. Issues such as buoyancy, fragility, and de-differentiation due to the characteristics of lipid-filled cells limit the effectiveness of traditional culture methods. To address these challenges, the utilization of primarily isolated human adipocytes was investigated.^[156] Techniques such as developing PDMS microchannels and employing BODIPY staining enabled real-time monitoring of fatty acid release, offering valuable insights into adipocyte behavior and function.

Adipose tissue is richly supplied with vessels formed by SVFs, highlighting the importance of vascularization in enhancing the physiological conditions of in vitro adipose tissue models. However, existing methods for culturing adipocytes and ECs rely heavily on tissue maturation medium, presenting challenges in achieving a balance between adipogenic differentiation and vessel formation. The presence of 3-isobutyl-1-methylxanthine (IBMX) in adipogenic induction medium was found to promote lipid accumulation while hindering vessel structure formation.^[157] Therefore, modifications to the co-culture medium are necessary to facilitate the vascularization of adipose tissue.

Dysregulated lipolysis contributes to ectopic lipid deposition and organ dysfunction, including nonalcoholic fatty liver disease and type 2 diabetes. In obesity, malfunctioning adipose tissue exacerbates proinflammatory M1 macrophage infiltration.^[158] Crown-like structure (CLS) morphology, observed in vivo during adipocyte-macrophage interactions in pathological adipose tissue, underscores the need to replicate such pathological environments accurately.^[159] To address this, an iPSC-derived adipocyte and macrophage microphysiology system was developed, enabling the creation of patient-specific preclinical models.^[160] This system allowed real-time observation of macrophage translocation, revealing insights into insulin resistance and dysregulated lipolysis in an obese environment. Moreover, the use of decellularized ECM derived from different obese states of adipose tissue was instrumental in modulating breast cancer adhesion and immune response, highlighting its potential for studying obesity-related pathophysiology and its interactions with other diseases.

Extrusion bioprinting has emerged as a viable technique for fabricating 3D adipose tissue. To preserve the microphysiology of adipose tissue, adipose tissue fragmentation was utilized by extrusion bioprinting using alginate and nanocellulose.^[161] This approach maintained a heterogeneous cell population like native tissue, with comparable proteomic profiles. Moreover, the printed adipose tissues maintained their volumetric structure for up to 150 days. In addition, embedded printing has been employed. In this method, preadipocytes were embedded in an alginate bioink within a calcium-containing adipose tissue-derived ECM bath to create densely packed adipose tissue.^[162] Over a period of 38 days, the printed adipose tissue was induced into an obesity and chronic inflammation state, leading to the production of pro-inflammatory cytokines that induced M2 macrophage monocyte migration. Thus, the 3D bioprinting strategy offers a broader selection of biomaterials, including patient-derived tissues with volumetric structures and long-term maintenance, entirely associated with metabolic activities and pathological behaviors, for in vitro adipose tissue modeling.

4.2. In vitro bone models

Bone tissue primarily consists of collagen I and a calcified matrix featuring both mineralized and non-mineralized components along with noncollagenous proteins. Structurally, bone tissue comprises a stiff mineralized outer layer and bone marrow situated within trabecular cavities, which supports hematopoiesis, producing immune cells, bone, and blood. Utilizing a microfluidic system is a recognized method for dynamic bone tissue formation, as the shear stress induced by fluid flow aids in the activation of primary cilia, mechanosensory organelles crucial in microfluidic culture.^[163] Recent advancements have aimed to replicate the microenvironment of bone tissue to better model bone-related pathologies.

Inflammation of bone tissue has been observed in conditions such as rheumatoid arthritis and osteoarthritis.^[164] This inflammatory environment can enhance the resorbing activity of osteoclasts and involves crosstalk with neuronal and vascular components in vivo.^[165] Consequently, an inflammatory bone model was developed to incorporate innervation and vascular migration dynamics.^[166] The effectiveness of this model was demonstrated through anti-inflammatory therapy using ibuprofen-loaded nanoparticles. Furthermore, the vascularization mechanism of bone tissue was a subject of investigation for both pathological understanding and tissue regeneration purposes. Conditions like avascular necrosis of the femoral head, linked to factors such as frequent alcohol consumption, cancer therapy, transplantation, and unknown etiologies, have highlighted the importance of vascular health in bone integrity.^[167] Especially glucocorticoid intake was associated with an increased risk of osteonecrosis due to impairment of bone microvascular ECs (BMECs).^[168] To study this phenomenon, an osteonecrosis model was developed to assess the therapeutic potential of DNA aptamers in attenuating glucocorticoid-induced apoptosis.^[169] Moreover, advancements in image-based artificial intelligence (AI) facilitated the analysis of bone disease models. For instance, in osteoporosis studies, AI algorithms were employed to quantify β-catenin expression, particularly in the context of romosozumab treatment.^[170]

To investigate the bone maturation process, a miniaturized in vitro bone model was developed using pre-differentiated adipose-derived stem cells (ADSCs)^[171] and patient-derived bone marrow stromal cells (BMSCs).^[172] This model was inspired by the stabilization of HIF-1 with high reactive oxygen species (ROS) production during osteogenesis, leading to the fabrication of an oxygen variant chip equipped with catalase and hydrogen peroxide. The mechanism behind the gradient expression of osteo-chondrogenesis from BMSCs was discovered by regulating oxygen levels in the osteogenesis model. Despite the prevalent use of cell-laden hydrogels in in vitro bone studies, existing methods relying on these hydrogels often result in non-mineralized structures or are simply combined with calcium and phosphate micro/nanoparticles. To address this limitation and achieve mineralization within microchannels, a hydroxyapatite-based system was fabricated using stereolithography-based ceramic manufacturing.^[173] This system also proved instrumental in facilitating osteogenic differentiation and establishing a gradient-controlled drug screening platform.^[174]

The bone marrow, nestled within trabecular cavities of bones, serves as the hub for hematopoiesis, generating immune cells, bone components, and blood. This stromal environment, consisting of ECMs, cell-secreting EVs, and cytokines, plays a critical role in conditions like myelofibrosis^[175] and leukemia.^[176] To replicate this stromal environment, a vascularized endosteal bone marrow niche was fabricated in a 96-well plate, aiming to study hematopoietic stem cell (HSC) behaviors in radiation-induced pathologies.^[177] Furthermore, an in vitro model of hematopoietic dysfunction was developed using a vascularized bone marrow strategy to mimic neutrophil abnormalities in patients with Shwachman-Diamond syndrome. This model successfully sustained CD34-positive progenitor cells for 4 weeks, aided by perfusion and vascularization, while responding to clinically relevant chemotherapy doses such as AZD2811 and 5-FU against myeloid leukemia.^[178] In addition, a leukemia-on-a-chip model replicated the leukemia microenvironment to understand therapy challenges.^[179] This chip was applied to further enhance the efficacy of CAR T immunotherapy.

The bone marrow microenvironment significantly influences the occurrence of osteosarcoma.^[180] To exert spatial control over the osteosarcoma matrix, decellularized osteosarcoma ECM was integrated into bioprinting alongside human MSC-derived EVs.^[181] Utilizing a PDMS mold with a PMMA structure and tubing, a dynamic culture of the 3D bioprinted osteosarcoma structure was achieved.^[179] Accordingly, the perfusion system enhanced pathological processes in osteosarcoma, mirroring patient and xenograft responses via the CXCL12/CXCR4/PI3K/AKT pathway. Overall, in vitro bone models represent a valuable tool for advancing our understanding of bone-related diseases and hematopoietic disorders, as well as for developing novel therapeutic strategies and personalized treatment approaches.

4.3. In vitro neural tissue models

The brain vasculature comprises neuronal and perivascular cells, forming the unique blood-brain barrier (BBB). The BBB plays a crucial role in maintaining central nervous system (CNS) homeostasis and protecting against pathogens through selective permeability. Tight junctions between ECs physically restrict the transportation of hydrophilic biochemicals while allowing essential hydrophobic molecules to pass through. Advancements in the physiology and maturation of in vitro BBB models have been achieved with microfluidic-based cultures compared to traditional 2D static cultures. These models enabled the study of BBB physiology in the context of interactions between brain-derived ECs, astrocytes, and pericytes.^[182] Particularly, astrocytes and pericytes in the BBB were highly responsive to neuroinflammation (TNF-a) in dynamic 3D cultures, demonstrating mature structural reconstitution compared to 2D static Transwell cultures.^[183] Furthermore, specific molecular systems regulating CNS angiogenesis, such as the influx transporter glucose transporter 1 (GLUT-1) and the efflux transporter P-gp, were controlled by the Wnt-β signaling pathway,^[184] which is crucial for barrier-specific properties. Utilizing soft lithography, BBB models were developed to study CNS angiogenesis and BBB maturation.^[185] These models facilitated the migration of human brain-derived ECs and pericytes into astrocyte-containing environments, demonstrating functional efflux transporter activity following an angiogenic gradient established by fibroblasts.

In recent developments, in vitro neural tissue models incorporating neurovascular structures were utilized to assess clinically relevant pathologies.^[186] These models assessed the metabolic interplay between the microvasculature and brain neurons, revealing the direct utilization of vascular metabolites such as glutamate and gamma-aminobutyric acid by neurons.^[187] This finding indicates the importance of brain vasculature in neurometabolism and neurodegeneration. Furthermore, the penetration of microbes into the brain was simulated using brain model platforms with neurovascular structures. Through real-time monitoring, these models elucidated the neurotropism and BBB penetration mechanisms of fungal meningitis pathogens.^[188] They revealed the process of transcytosis by remodeling the EC layer, a phenomenon not observed in traditional Transwell studies.

The choroid plexus, a secretory tissue located in the brain ventricles responsible for producing cerebrospinal fluid (CSF) via ependymal cells, has garnered attention.^[189] Clinical studies demonstrated that dynamic CSF flow could influence the efficacy of intrathecal chemotherapy for CNS-derived tumors.^[190] To mimic the dynamic oscillatory flow of CSF, the light-based printing technique was employed to fabricate choroid plexus models, optimizing the frequency of the rocking system.^[191] These models successfully recapitulated the function of the capillary-epithelium barrier in the choroid plexus through dynamic culture, thereby demonstrating enhanced chemotherapy response and immune reactions to cancer metastasis, akin to in vivo conditions.

One of the most prevalent malignant brain tumors is glioblastoma multiforme (GBM), associated with a high mortality rate of approximately 6 months.^[192] Therefore, there is a significant need for clinically relevant platforms to facilitate rapid drug screening and treatment prediction for selecting chemotherapy candidates. Traditional animal models, such as cell-derived xenografts, have been utilized to substitute human stromal tissue for murine tissue. However, this approach is time-consuming, taking up to six months to generate disease models.^[193] Interestingly, an in vitro GBM model was developed to recreate physiological relevance by controlling hypoxic environments, providing support for the ECM, and facilitating interaction with stromal tissue. This model reconstructed GBM physiology using patient-derived GBM cells and stromal cells incorporated into decellularized brain ECM via a multi-dispensing extrusion bioprinting technique.^[144] The gas permeability of this bioprinted GBM model allowed for the control of spatial pathological progression. Remarkably, this model exhibited patient-specific responses to irradiation and chemotherapy. Furthermore, a vascularized GBM model was generated through the extrusion bioprinting of Pluronic F127 as a sacrificial bioink for microchannels and fibrin-based bioink for the cancer growth matrix.^[194] The fibrin bioink accurately replicated the dormancy phenomenon observed in human GBM, a feature not observed in SCID mice or 2D culture systems. Genomic analysis of these bioprinted GBMs, conducted via RNA sequencing, demonstrated a closer resemblance to in vivo conditions compared to 2D culture systems, particularly in angiogenesis hallmarks, the JAK-STAT signaling pathway, and interferon-g response.

4.4. In vitro cardiac tissue models

The heart is a vital organ responsible for pumping blood throughout the body to facilitate oxygen exchange. It mainly comprises cardiomyocytes, which orchestrate the systolic (contraction) and diastolic (relaxation) processes crucial for its function. Cardiomyocytes derived from iPSCs and neonatal tissue have found widespread use in heart tissue engineering due to their ability to exhibit spontaneous beating behavior.^[195] Critical characteristics of functional cardiomyocytes, such as their electrophysiology and mechanical forces, serve as representative markers for assessing toxicology and tissue maturation. Despite passing clinical evaluations, certain drugs have been withdrawn from the market due to their cardiotoxic effects. For instance, Rofecoxib, an anti-inflammatory drug approved by the FDA, was associated with over 100,000 heart attacks.^[196] Similarly, Linsitinib, a tyrosine kinase inhibitor used in cancer treatment, has been linked to an increased risk of atrial fibrillation.^[197] Consequently, there has been a growing emphasis on preclinical in vitro cardiac models to enhance the efficacy of drug evaluations. In a cardiac model, the failure of Linsitinib in the context of bone cancer was recapitulated, exhibiting delayed calcium transient in cardiomyocytes and proarrhythmic events consistent with clinical data.^[197]

Long QT syndrome (LQTS) is characterized by abnormalities in the heart’s electrical system, potentially leading to life-threatening arrhythmias. LQTS patients often present mutations in the KCNH2 gene, reducing the rapid delayed rectifier (IKr) channel crucial for cardiac action potential repolarization.^[198] An LQTS model utilized iPSCs-derived cardiomyocytes carrying the amino acid (R531W)-mutated KCNH2 gene, generated using CRISPR-Cas9 technology.^[199] This LQTS model was employed to validate pharmaceutical interventions aimed at correcting abnormal calcium transients.

Myocardial infarction (MI) frequently arises from inadequate blood supply. An ischemia model was developed using lithography techniques, illustrating the detrimental effects of hypoxia on synchronous contraction, which is fundamental to the study of ischemic cardiac tissue.^[200] Mainly, spatiotemporal control of oxygen levels was achieved through normoxic and hypoxic dual channels in the MI model system.^[201] The oxygen gradient-regulated chip exhibited robust activation of inflammatory responses, including interleukin-6 (IL-6), IL-17, and mitogen-activated protein kinase signaling, as observed through RNA-seq analysis, compared to uniform normoxia and hypoxia conditions.

As both electrophysiology and chemical stimuli influence heart tissue, sensors have been developed to detect cardiomyocyte activity in real-time. Oxygen sensors integrated within heart models enabled the measurement of metabolic changes in situ, facilitating the detection of cardiomyocyte maturation through electrical stimulation.^[202] Moreover, resistive sensors, capable of detecting static and cyclic strain, were employed to monitor spontaneous contractions using direct laser writing technology.^[203] However, integrating sensors into the in vitro cardiac tissue models has presented challenges such as manual assembly, hindering insulation of culture medium flow, aseptic processing, and high throughput screening. To address these limitations, multimaterial 3D bioprinting methods were utilized to fabricate instrumental cardiac microphysiology devices for laminar cardiac tissue.^[204] These devices incorporated features like cantilevers, strain gauge wires, and microfilaments, which were one-step printed within PDMS frames to measure cardiac tissue contraction and guide tissue alignment.^[205] These sensors were applied to investigate interactions between human ECs and cardiomyocytes, where the restoration of contractile function by endothelial EVs was assessed using thin-film cantilever bending to measure strain gauge resistance. Furthermore, flexible 3D micropillar electrodes and microwires, directly printed with conductive polymers, enabled electrophysiological assessments of extracellular field potential and contractile force measurement.^[206] Additionally, an in situ monitoring system employing electrodes tracked cardiomyocyte behavior in contractile dynamics during chronotropic drug dosing^[207] and hypoxia cultivation.^[208] Recent advancements include the application of two-photon direct laser writing to develop miniaturized hearts.^[209] Microfluidic cardiac chambers, powered by iPSCs-derived cardiomyocytes, mimic ventricular function as miniaturized unidirectional pumps with involuntary control.

4.5. In vitro hepatic tissue models

The liver plays a crucial role in maintaining homeostasis, including regulating drugs, glucose, and albumin levels in the body. It is comprised of parenchymal cells, primarily hepatocytes, and non-parenchymal cells such as liver sinusoidal ECs (LSECs), hepatic stellate cells, and Kupffer cells. The liver’s structure consists of hepatic lobules and sinusoids. Within the sinusoids, there is a gradient-dependent zonation: zone 1 receives ample oxygen and nutrients from the artery, while zone 3 contains metabolic waste destined for the central vein.^[210] Various factors, including viral infections, fatty diet, alcohol consumption, and exposure to toxic chemicals, can initiate the progression of liver disease. Damage to parenchymal cells triggers the activation of non-parenchymal stem cells. Specifically, injured hepatocytes release ROS and pro-inflammatory cytokines, which activate LSECs and Kupffer cells. Additionally, hepatic stellate cells secrete collagen, leading to increased tissue stiffness and fibrosis.

Non-alcoholic fatty liver disease (NAFLD) is a prevalent chronic liver condition associated with hepatic steatosis, cirrhosis, cancer, and cardiovascular diseases.^[211] Liver transplantation remains the only treatment option, with no commercialized drugs available. Therefore, it is crucial to replicate the microenvironment of NAFLD to evaluate potential drug candidates. Given the strong correlation between NAFLD and obesity, as well as type II diabetes, the effects of metabolites derived from adipose tissue in hepatic models have been investigated. Treatment with free fatty acids (FFAs) resulted in an increase in NAFLD-like gene expression networks, which were analyzed using single-cell sequencing in a gut-liver-axis (GLA) model.^[212] Additionally, white adipose tissue was co-cultured in a hepatic model.^[213] Proinflammatory cytokines released from the adipose tissue chamber led to increased lipid storage in hepatocytes. Metformin, a drug used to treat type II diabetes, significantly affected steatosis by reducing lipid accumulation. Furthermore, to model primary sclerosing cholangitis, a vascularized bile duct model was fabricated using patient-derived organoids and soft lithography techniques.^[214] Cholangiocytes derived from patients exhibited increased production of inflammatory cytokines, which recruited immune cells from the vascular channel into the bile duct channel.

To mitigate the risk of frequent metastasis of breast cancer to the liver, reprogramming technology was employed in the development of a bioprinted in vitro hepatic model.^[215] Breast cancer cells were reprogrammed through lentiviral expression of transcription factors, including HNF4, FOXA2, FOXA3, ATF5, PROX1, and HNF1, converting them into induced hepatocytes. Reprogrammed hepatocytes exhibited enhanced functionality, as evidenced by increased albumin and urea secretion compared to traditional 2D and static culture methods. The bioprinted heptic structure, combined with GelMA-liver EV bioink, effectively remodels liver microphysiology. This study highlights the potential of the bioprinted hepatic model as a platform for gene-editing tests aimed at preventing metastasis.

When biochemicals circulate in the bloodstream, cytochrome P450 enzymes play a crucial role in detoxification. Therefore, the evaluation of newly developed drugs must consider liver toxicity, which can lead to chronic liver damage. However, the poor prediction of drug-induced liver injury has hindered the translation of preclinically tested drugs to clinical applications.^[216] Particularly, there is a gap in metabolism between animals and humans that has not been adequately addressed.^[217] Several liver models have been developed to overcome this challenge. These models recapitulated conditions such as steatosis and fibrosis and included species-specific Kupffer cells and stellate cells. Furthermore, they unveiled idiosyncratic toxicities, like the one seen with a G protein-coupled receptor 40 agonist for diabetes, which exacerbated liver injury during Phase 3 trials.^[218] A vascularized liver model showed enhanced expression of CYP450 enzymes in the presence of ECs, providing a more accurate representation of drug metabolism compared to 3D liver spheroids.^[219] This model was tested with nicotine, which resulted in junctional irregularities in the microvasculature due to the downregulation of GTP cyclohydrolase 1. A “Space of Disse” model was developed to evaluate toxin-induced veno-occlusive disease, which obstructs the liver sinusoids.^[220] With a perfusable lumen structure, this model simulated azathioprine’s effects on EC regression. Additionally, computational fluid dynamics simulations were used to assess the impact of donepezil, an Alzheimer’s disease drug, on CYP3A4 enzyme activity.^[221]

3D bio-dot printing has emerged to deposit primary hepatocytes in a way that promotes in situ spheroid formation.^[222] This innovative approach addressed the issue of diminishing hepatic functionalities observed in isolated primary hepatocytes due to the loss of cell-to-cell and ECM interactions, which significantly affect albumin, urea, and CYP enzyme activities. The bioprinted hepatocyte spheroids were utilized to investigate distance-dependent hepatocyte-EC interactions and for hepatotoxic drug testing using liver-derived ECM.^[223] Additionally, an automated multi-dispensing system was implemented to create miniaturized liver lobule structures for drug screening purposes.^[224] EC patterning within these structures simulated hepatoprotective effects in cases such as acetaminophen (APAP) toxicity. Leveraging pre-set syringe printing, a model using endothelialized hepatic spheroids was developed for high-throughput drug screening.^[225] Furthermore, extrusion bioprinting facilitated the one-step fabrication of liver models integrated with a perfusion system. The perfusion system was refined to allow for precise deposition of materials such as poly(ethylene-co-vinyl acetate) (PEVA)^[226] and PCL,^[227] with the former used for framing and the latter for containing bioinks intended for drug toxicity testing. These one-step processes utilizing extrusion bioprinting techniques have the potential to create 3D bioprinted tissue models, offering a promising avenue for accurately studying in vivo biological processes in a high-throughput and reproducible manner.

4.6. In vitro intestinal tissue models

The intestine is responsible for both acting as a barrier and facilitating nutrient absorption. Its unique crypt-villus structure maintains a continuous epithelium, preserving the stem cell niche within the basal crypt to shield against exogenous antigens and metabolites from the microbiome.^[228] Gut epithelial responses to microbial exposure encompass a range of processes, including mutation accumulation, effects on proliferation, and disruption of barrier integrity.^[229] Furthermore, microbial exposure triggers the activation of Toll-like receptors (TLRs), leading to subsequent cytokine release and recruitment of immune cells. Similarly, an in vitro gut model was successful in recapitulating the degradation and invasion of intestinal tissue by pathogens like Entamoeba histolytica and the colonization process of Shigella flexneri.^[230] Conversely, the therapeutic potential of probiotics was demonstrated in gut models, where their barrier-enhancing effects were observed in damaged epithelial layers exposed to lipopolysaccharide.^[231] Moreover, the role of immune mediators in inflammatory stimulation was evaluated through the analysis of reactive oxygen and inflammatory cytokines in immune-responsive human microbiota-intestine axis models.^[232]

The polarization of intestinal epithelial cells typically requires 3 weeks in static culture, often resulting in inadequate differentiation and functional characteristics of these cells.^[233] However, enhanced epithelial cell differentiation was achieved through dynamic flow culture, leading to morphological changes in villi height.^[234] A hybrid model, insertable into Transwell membranes, was developed to mimic the crypt-villus structure.^[235] Control of apical and basolateral flow within this microfluidic chip induced 3D morphogenesis of Caco-2 cells and intestinal organoids, resulting in lumen heights of 120–150 μm. Utilizing PDMS as a negative mold, collagen scaffolds were prepared, and intestinal organoids were seeded, spontaneously reconstituting spatially differentiated crypt-plateau-villi structures through shear stress and stromal-epithelial cell interactions.

Advancements in bioprinting technologies have enabled the self-organization of centimeter-scale gastrointestinal tracts, complete with lumens, vasculature, and tubular intestinal epithelia.^[236] Scaling up intestinal models via bioprinting techniques has proven beneficial for remodeling injury models, particularly for conditions such as intestinal mucosal damage caused by ischemia/reperfusion injury (IRI).^[237] Inspired by native intestinal villi, an artificial vessel-embedded intestinal organoid model was developed using coaxial extrusion methods.^[238] Hollow fibers were fabricated using pH-sensitive zeolitic imidazolate framework-8 (ZIF-8) alginate, allowing for the simulation of vessel-oriented hypoxia and reoxygenation. Treatment with UV damage and hydrogen peroxide-containing medium induced the secretion of proinflammatory cytokines from vessels. The IRI chip investigated OLFM4 protein, related to NF-kappa B signaling in intestinal organoids, as a potential therapeutic target for conditions such as celiac disease or inflammatory bowel disease, aiming to reduce intestinal inflammation. The in vitro intestine models provide a versatile platform for investigating various aspects of intestinal biology, disease pathology, and therapeutic interventions.

4.7. In vitro renal tissue models

The kidneys serve a critical role in the body by filtering drugs and metabolites through the vasculature and nephron, maintaining fluid and electrolyte balance. The nephron, the basic structural and functional unit of the kidney, consists of various components, including podocytes in glomerular capillaries, epithelial cells in proximal tubules, loops of Henle, and distal convoluted tubules.^[239] Efforts to replicate the intricate functional and structural complexity of native nephrons have accelerated the clinical translation of in vitro renal tissue models for toxicology studies. Renal tubuloid cultures, established using primary cells derived from adult kidney tissue or urine, enabled high-throughput screening of drug efficacy.^[240] Furthermore, the maturation of kidney organoids was facilitated by fluid shear stress, which enhanced the expression of multidrug and toxic compound extrusion (MATE) genes.^[241] Patient-derived organoids hold promise as clinical tools for identifying therapeutic targets and testing new drugs.

Autosomal recessive polycystic kidney disease, a rare genetic disorder characterized by the formation of fluid-filled kidney cysts due to mutations in the polycystic kidney and hepatic disease 1 (PKHD1) gene, has been studied using kidney organoids derived from patient-derived iPSCs.^[242] Unlike normal kidney organoids, PKHD1^−/− organoids demonstrated cyst formation in the distal nephron when cultured under fluidic conditions, driven by overexpression of mechanosensing genes.^[243] Treatment of the PKHD1 model with targeted therapy using the Rho family of GTPase proteins successfully suppressed cyst formation.

Renal dysfunction affects approximately 15% of the U.S. population, contributing significantly to the disease burden and necessitating transplantation for many individuals.^[244] To evaluate nephrotoxicity, proximal tubule models have been extensively investigated. These models replicated polarized renal epithelium and assessed the response of basal-apical uptake mechanisms. Mechanisms of receptor-mediated endocytosis in the proximal tubule model were elucidated, revealing alterations in basolateral uptake^[245] mediated by epidermal growth factor (EGF)^[246] using antisense oligonucleotides, which have been clinically associated with proximal tube accumulation. Furthermore, vascularized human kidney spheroids incorporating proximal tubule cells and microvascular ECs were developed to study the nephrotoxic effects of immunosuppressive drug cyclosporine and cancer chemotherapy drug cisplatin.^[247] Integration of a vascularized human kidney model with tissue-embedded microsensors could enable real-time assessment of cellular metabolism.^[247] Amperometric four-analyte sensors detected drug-induced metabolite accumulation in glucose metabolism, complemented by immunostaining for lipid formation. Additionally, patients experiencing nephrotoxicity due to treatment with the sodium-glucose cotransporter-2 (SGLT2) inhibitor empagliflozin demonstrated significant reductions in creatinine, uric acid, and lactate dehydrogenase levels.

Bioprinting technologies have significantly advanced the maturation of proximal tubule models and improved the accuracy of high-throughput screening. A renal tubule model was developed using extrusion bioprinting technique to simulate insulin barrier and albumin reabsorption functions based on polarized apical and basal structures.^[248] Kidney organoids can also be directly printed onto Transwell membranes, facilitating the controlled self-organization of glomerular and distal nephron formation through bioprinted dots and lines.^[249]

The embedded printing was also utilized to fabricate perfusable channels, enabling the segmentation of glomerulus microvascular endothelial channels and proximal tubules and facilitating real-time analysis of renal absorption.^[250] Similarly, the one-step fabrication of a pumpless vascularized proximal tubule model was achieved through a 3D multi-dispensing bioprinting system, aiming to model renal dysfunction induced by cisplatin-induced nephrotoxicity.^[149] This fabrication process involved melting extrusion of PEVA for the pumpless chip and coaxial printing of sacrificial Pluronic F127 bioink with human primary renal proximal tubule epithelial cells to create perfusable channels. These advancements in bioprinting technologies show great potential for the development of clinically relevant in vitro renal tissue models. These models can offer valuable insights into kidney physiology and pathology, and they hold promise for applications in drug discovery and personalized medicine.

4.8. In vitro lung models

The lungs serve as vital organs in the respiratory system, facilitating gas exchange between air and blood. Therefore, lung tissue is susceptible to various infections, toxins, and diseases such as pneumonia, tuberculosis, and cancers.^[251] With the emergence of contagious infection through the respiratory system, a lung infection model by viruses and bacteria has been developed to investigate penetration mechanisms and to screen drug candidates.^[252] High-throughput screening of drugs by infection variants was especially required for global health to cope with the pandemic.^[253]

The alveolar barrier in the lungs consists of a thin, multilayered structure comprising both epithelial and endothelial layers.^[254] To replicate the alveolar-capillary interface, the importance of employing fluidic shear stress culture and cyclic stress resembling breathing patterns was highlighted for tissue maturation.^[255] A lung model, inspired by breathing mechanics, was created with the aid of a cyclic vacuum in the alveolar channel.^[256] This model successfully mimicked IL-2-induced pulmonary edema, characterized by vascular leakage and extensive fibrin deposition within the alveolar airspace.

In efforts to replicate tumor growth and metastasis, lung cancer has been implanted into orthotopic sites in animal models. However, the limitations of xenograft models, including challenges related to stromal-cancer interactions, the tumor microenvironment, and the absence of dynamic fluidic culture systems characteristic of humans, have been well-documented. Non-small-cell lung cancer (NSCLC), comprising approximately 40% of adenocarcinomas, often arises at the bronchiolar-alveolar border and is highly dependent on its microenvironment.^[257] In response to these challenges, an NSCLC model was developed to replicate tumor growth dynamics within a microenvironment that included epithelial and endothelial layers alongside a vacuum-assisted breathing mechanism.^[258] This model revealed that mechanical strain experienced during breathing suppressed the activation of tyrosine-protein kinase c-Met, contributing to chemotherapy resistance, consistent with clinical observations.

Barrier permeability is a critical function of the lungs. A pulmonary edema model demonstrated similar barrier permeability to that observed in ex vivo mouse disease models induced by IL-2. During cancer radiation therapy, lung injury often occurs, resulting in pneumonitis and fibrosis. In addition, a radiation-induced lung injury (RILI) model was developed to simulate the effects of radiation damage on lung epithelium and endothelium.^[150] This model increased DNA fragmentation and inflammatory cytokine production while weakening barrier functions. The RILI model allowed the testing of FDA-approved drugs such as lovastatin and prednisolone, which target heme oxygenase expression and suppress inflammatory reactions during radiation exposure.

Traditional methods of measuring permeability, such as transendothelial electrical resistance (TEER), are limited in providing accurate readings due to minor deformations, even in well-formed tight junctions. To address this, antibody-based electrochemical sensors were developed to substitute TEER-dependent evaluations by detecting large analyte penetration through disrupted alveolar barriers.^[259] Additionally, a bubble-inspired visualization technique was utilized to diagnose breathing-induced deformation. This technique detected rhythmic contraction and expansion in idiopathic pulmonary fibrosis (IPF) models.^[260] This model applied cyclic deformation from airflow to sense mechanical stretching induced by airflow (30 kPa), leading to a fibrotic reaction in pulmonary alveoli accompanied by ECM accumulation.

The air-liquid interface plays a critical role in facilitating epithelial polarization and maturation within lung tissue. A multi-drop dispensing system was utilized to automate tissue fabrication and foster air-liquid interface maturation.^[261] This approach allowed human-derived primary small airway epithelial cells to mature and facilitated high-throughput screening of epithelial barrier function responsiveness to pro-inflammatory cytokines. Moreover, direct extrusion bioprinting into a Transwell insert was employed to develop vascularized airways through layer-by-layer stacking.^[262] However, conventional tissue fabrication technologies were limited to achieving the necessary resolution for remodeling thin lung tissue, typically ranging from 50 to 200 mm.^[263] To address this challenge, jetting printing was utilized to construct ultra-thin multilayered alveolar barriers with a thickness as low as 8 mm. This technology supported the precise patterning of layers containing ECs, fibroblasts, and alveolar cells onto a Transwell insert using high-resolution drop-on-demand (DOD) inkjet printing, followed by UV light crosslinking to ensure stability.^[264] The integration of a perfusion system with the insert was shown to enhance the expression of epithelial junction-related genes compared to traditional 2D culture methods. This advancement highlights the importance of incorporating dynamic microenvironments to better mimic physiological conditions.

4.9. In vitro skin models

The skin, the body’s largest organ, encompasses immune functions, vascularization, sensory perception, and hair follicles. As a physical barrier, it regulates exposure to external environments, pathogens, and nanoparticles. Structurally, the skin consists of the hypodermis, dermis, and epidermis layers, each with distinct roles. Various dynamic culture methods have been developed to mature in vitro skin models, including air-liquid interface culture for epithelial maturation, cyclic stretching for stratification, and microfluidic culture with immune and vascular cell infiltration.

In recent advancements, paper microfluidics were utilized to develop in vitro epidermis models, enabling air-liquid culture with differentiated organelles and the formation of involucrin-positive layers with tight junctions.^[265] Physical stimulation techniques, such as cyclic stretching, enhanced keratinocyte stratification, and increased epidermal thickness. Moreover, 3D bioprinting has emerged as a suitable fabrication technology for skin equivalents, offering improved uniformity in epidermal layer fabrication compared to manual seeding methods. Extrusion bioprinting allowed for the segmentation of epidermis-dermis layers, resulting in thicker epidermal layers with enhanced properties like TEER and restricted dextran diffusion.^[266]

Furthermore, bioprinted in vitro skin models have facilitated the perfusion of pre-vascularized endothelial structures within skin layers, offering opportunities to recapitulate the pathophysiology of skin diseases like diabetic ulcers.^[267] These models supported the simulation of vascular dysfunction, adipocyte hypertrophy, and inflammatory responses in hyperglycemic conditions, allowing for the study of delayed re-epithelialization and the effects of diabetic drugs. Additionally, embedded bioprinting techniques were used to create 3D hair follicle spheroids within skin models, promoting the maturation of follicle-like structures through keratinocyte and melanocyte migration.^[268]

In viral infection models, embedded printing with functional microvascular structures demonstrated efficacy in inhibiting herpes simplex virus infection by facilitating the transmigration of neutrophils and the delivery of antiviral drugs like acyclovir through printed microvasculature.^[151] These improvements highlight the potential of advanced fabrication methods in developing sophisticated skin models for various applications, including drug testing and disease modeling.

4.10. In vitro pancreatic tissue models

The pancreas plays a vital role in regulating nutrient digestion through its exocrine function and glucose metabolism hormones through its endocrine function. Diabetes, a chronic condition, manifests as inadequate insulin production (Type I) or the body’s resistance to insulin (Type II), affecting over 10% of the population based on the National Health Survey in the US. For Type I diabetes treatment, 10,000 iEQ/kg of primary pancreatic islets are required, but donor health quality and functional loss pose challenges to successful implantation, necessitating pre-transplantation assessment of islet graft function. Electrophysiology of islets, correlated with the CHIP-score (clinical trial number NCT03067324), utilizing a multi-electrode array chip, offers a method to predict islet functionality before transplantation.

Type II diabetes is associated with complications such as diabetic retinopathy and ulcers due to damaged blood vessels from hyperglycemic environments. To address glucose regulation for healthy longevity, a hollow microfiber-assembled endocrine pancreas model was developed, incorporating a vascularized structure with pancreatic islet cells (β-TC-6) and islet-derived ECs (MS1).^[269] This model allowed in vitro safety testing of sugar substitutes through insulin and glucagon secretion. A novel living pancreas slicing technique also preserved exocrine and endocrine structures, offering a paradigm shift from conventional isolation methods and allowing for dynamic glucose-stimulated insulin secretion (GSIS) monitoring.^[270]

In the context of insulin resistance associated with obesity, chronic inflammatory secretion from adipose tissue triggers pathological macrophage reactions, contributing to diabetes pathology. While murine-derived beta cell lines like Min 6 secrete insulin, high-passage cell numbers may not validate GSIS due to reduced ATP contents. Thus, enzymatically isolated primary pancreatic islets were utilized for diabetic studies, with functional restoration demonstrated in co-culture microfluidic systems with primary rat liver hepatocytes.^[271] However, species-dependent cytochrome enzyme activation limited the recapitulation of human drug response in this model. To address these challenges, a multi-material dispensing system was employed to fabricate a glucose metabolism-related pancreas-adipose tissue-liver axis, enhancing tissue-specific functionalities and enabling the study of type 2 diabetes pathology.^[152] This model successfully demonstrated insulin secretion and glucose uptake recovery in a time-dependent manner with drugs like metformin and tolbutamide.

5. Summary and Outlook

Since the inception of tissue engineering in 1993,^[272] there have been high hopes for swift advancements and significant benefits for patients. However, the journey proved to be more challenging than anticipated, with numerous trials and setbacks highlighting substantial limitations in clinical applications, contrary to initial optimism. Despite these hurdles, this process has greatly expanded our knowledge and understanding of biology. The integration of current 3D bioprinting technologies represents a promising step forward, bringing us closer to developing more realistic tissue constructs for implants and advanced in vitro tissue models. It is important to realize that bioprinting is not a magical solution capable of creating any tissue structure at will. Success in clinical applications relies on the creation of meticulously designed tissue constructs based on a deep understanding of biology. Setting achievable goals in a step-by-step manner is crucial, considering the current limitations of available technology. By setting incremental technical objectives and planning clinical applications accordingly, we can maintain optimism for progress.

Despite the rapid advancements in 3D bioprinting techniques globally, the fabrication of fully functional tissues or organs still needs to be discovered. Numerous challenges persist in producing implantable tissue and organ constructs suitable for clinical applications and therapeutic in vitro tissue models. Achieving vascularization and innervation of bioengineered tissues are crucial milestones for constructing and engrafting functional constructs. While several strategies have been explored for vascularization, finding a solution for the vascularization of volumetric or human-scale tissue constructs remains challenging. Additionally, limited progress has been made in designing innervation. Moreover, a comprehensive understanding of the role of the immune system and the foreign body response is needed to ensure functional engraftment upon the implantation of tissue constructs. Recreation of whole organs requires a detailed biological understanding of tissue-specific cell populations and phenotypes to replicate the anatomy and physiology of the organ, including cell-to-cell and cell-to-ECM interactions, as well as morphogenesis.

Efforts to advance bioprinting techniques have primarily concentrated on extrusion, jetting, and light-based methods, each possessing unique advantages and drawbacks. However, to achieve clinically relevant tissue construct fabrication, combining these approaches into a unified multifunctional bioprinter may be necessary. This integration allows for the fulfillment of varied requirements with enhanced efficiency and precision. Several research groups have demonstrated the benefits of such combination bioprinters, leveraging the strengths of different techniques. For example, combining high-viscosity hydrogels with extrusion methods enhances structural stability, while coaxial extrusion enables the creation of vascular networks.^[273] Furthermore, the combination of jetting-extrusion^[274] facilitates precise cell placement and material deposition, while integrating light-extrusion bioprinting^[275] allows for high-resolution microstructure fabrication and spatially controlled crosslinking. Developing tissue-specific bioprinters tailored to distinct tissue types can further enhance the versatility and application of these technologies.

The clinical workflow of 3D bioprinting involves several crucial components, as outlined previously. Especially, significant advancements have been made in the development of novel bioink systems to enhance printability while maintaining high-resolution capability and structural integrity. These bioinks, composed of advanced biomaterials such as hydrogels and polymers, serve as both cell delivery vehicles and supportive structures, offering biological properties as well as mechanical and structural support essential for successful bioprinting. Furthermore, ongoing progress in biomaterials tailored to different bioprinting mechanisms is crucial for the sustained advancement of tissue engineering applications. A recent innovative approach involves utilizing decellularized ECMs, which provide tissue-specific microenvironments for cells. These ECM-based bioinks closely mimic in vivo conditions and offer critical cues for targeted cell engraftment, survival, and tissue formation, making them highly promising for various applications. Further advancements in technology and the development of new bioinks may eventually lead to the restoration of complex organs with microvasculature and precise internal structures.

Successful bioprinting for clinical trials relies on the large-scale production of patient-derived primary cells. Automated bioreactor systems are now crucial for scaling cell expansion, minimizing human error, and maintaining consistent culture conditions through even distribution of nutrients and oxygen for uniform cell growth. Additionally, standardized, chemically defined, and xeno-free cell culture media are essential for clinical applications. Efforts are underway to develop such media tailored to cells from each of the three embryonic germ layers, using synthetic, recombinant, and human-sourced molecules to create well-defined formulations. Adoption by the clinical manufacturing community could streamline FDA approval processes.^[276]

Bioprinted tissue constructs must mature to replicate tissue-specific functions. Tissue-specific bioreactors are instrumental in this process, especially for clinical applications, as they mimic physiological conditions by delivering nutrients, removing waste, supplying oxygen, and maintaining pH in a sterile environment. These systems play a crucial role in developing optimal tissue microarchitecture, functionality, and durability through targeted stimulation and conditioning processes, such as mechanical and electrical stimulation. Real-time monitoring of key chemical components (e.g., pH, O₂, glucose/lactate) enables automatic control of the environment, ensuring safe and reproducible tissue culture. Evaluating functional tissue parameters during maturation is essential to confirm the construct’s suitability for patient implantation.^[277]

The successful bioprinting of fully functional tissues or organs relies on collaboration across various disciplines, including engineering, materials science, biology, medicine, and business administration. Each field brings unique expertise, contributing to the advancement of bioprinting technologies and their translation into clinical applications. Despite the complexities involved, the biofabrication and biomanufacturing community is making significant strides in developing tissue constructs for preclinical models and is now focused on scaling up these technologies to create human-scale tissues. Continued advancements in bioprinting technologies and collaborative efforts across disciplines bring the goal of clinical translation closer to reality. Looking ahead, tissues such as skin, bone, cartilage, vascular patches, cardiac patches, and peripheral nerve grafts are positioned to lead the way toward clinical implementation. With ongoing innovation and collaboration, the prospect of bioprinting fully functional tissues and organs holds promise for transforming regenerative medicine and improving patient outcomes.