Three-Dimensional Printing/Bioprinting and Cellular Therapies for Regenerative Medicine: Current Advances
Nowadays, regenerative medicine is considered an emerging field of research worldwide, with the potential to revolutionise healthcare (improving patient outcomes and quality of life) in the 21st century . This field is focused on the replacement, engineering and regeneration of human cells, tissues and organs. Its aim is to repair, restore, supplement or replace the normal function of a biological system following treatment with autologous, allogeneic stem and stromal cells [ , , ]. In recent years, two promising approaches have gained attention: namely, three-dimensional (3D) printing and cellular therapies. 3D printing has revolutionized tissue engineering through the manufacture of complex, patient-specific materials whilst providing precise control over the spatial distribution of cells and biomaterials. On the other hand, cellular therapies encompass the use of living cells to replace, repair, or regenerate injured tissues. The integration of 3D printing with cell therapies provides personalised, cell-laden constructs that mimic the complex architecture and function of native tissues. This approach stimulates cell–cell interactions, cell–matrix interactions and the creation of functional tissue-like constructs . This review presents an overview of the current advances in the combination of 3D printing/bioprinting and cell therapies in the field of regenerative medicine. Firstly, the principles and applications of 3D printing/bioprinting and cell therapies are addressed individually. Subsequently, we explore recent advances in the integration of these technologies to create 3D-printed cellular structures. The limitations and challenges associated with these approaches are also mentioned, and the current state of the field is analysed. Finally, we present a future perspective on the potential impact of these technologies in the field of regenerative medicine and conclude with a summary of the main findings. By providing a detailed understanding of the current state of the field and future directions, this review seeks to contribute to the growing body of knowledge in this challenging and rapidly evolving field. 

Three-dimensional (3D) printing (alternatively termed additive manufacturing (AM)) is a burgeoning technique for the rapid prototyping of structures in diverse fields, including regenerative medicine [ , , , , ]. In the past two decades, 3D printing has been widely researched due to its simplicity and the fact that its highly flexible manufacturing provides unlimited possibilities for creating complex structures . Several examples of 3D printed medical devices include the following: instrumentation, including guides to facilitate the correct surgical placement of devices; implants, such as hip joints; and external prostheses, such as bionic hands . This advanced manufacturing technology, pioneered by Charles Hull in the early 1980s, enables the creation of customized and intricate structures with high precision and accuracy . The technology operates by using computer-aided design (CAD) software to create a virtual model of the structure, which is subsequently converted into a physical object through the use of various 3D printing techniques . The International Standard Organization (ISO) and the American Society for Testing and Materials (ASTM)  have established a classification system comprising seven overarching categories to categorize AM methods based on part manufacturing approaches, as outlined in . The subsequent sections provide a succinct overview of AM processes. 2.1. Vat Photopolymerization This method is characterized by the principle of selectively solidifying a liquid photopolymer resin by exposing it to specific light sources or patterns. In the procedure, a vat contains the liquid resin, which is selectively cured layer by layer in the desired regions, gradually constructing the intended object. The laser, guided by predetermined digital coordinates, provides the necessary information to cure the resin at precise locations within each layer. To facilitate this curing process, a photoinitiator is introduced into the resin material. The photoinitiator absorbs the incident ultraviolet (UV) radiation and generates active species that initiate the photopolymerization reaction . The vat photopolymerization technique includes various commonly utilised processes, such as stereolithography (SLA), continuous light interface production (CLIP), digital light processing (DLP) and two-photon polymerization (2PP) . These methods involve SLA, which uses a laser or digital light projector to cure the resin layer by layer, whereas DLP employs a digital light projector to cure an entire layer at once. CLIP employs a continuous liquid interface production approach, in which a continuous liquid interface is established between the resin and a transparent window, that enables continuous printing. Lastly, 2PP utilises a focused laser beam and a photosensitive resin to achieve high-resolution printing via a nonlinear optical process . These vat photopolymerization techniques offer versatile and AM capabilities for fabricating complex and detailed structures. 2.2. Material Extrusion Materials extrusion is a widely adopted AM technique due to its rapid manufacturing, cost efficiency, simplicity, user-friendly nature and the potential to produce complex components [ , , ]. This technique involves extruding the material through an orifice and depositing it onto a construction platform [ , , ]. The most common technique is fused filament fabrication (FFF), which is also referred to as fused deposition modelling (FDM) . In this method, the filament is heated to a temperature typically between 150 and 250°C until it melts, and the molten material is extruded through a nozzle . The extruded material is then placed onto a build plate, which can be adjusted in the z direction. This technique usually uses more polymers, although it is also used to print metal and ceramic components. FFF uses a support to create overhang features, which can be removed either mechanically, by detaching it from the printed part, or chemically, by dissolving it in a solvent . In contrast to FFF, which employs thermoplastics, direct ink writing (DIW) is a 3D printing technique that utilises viscoelastic inks. In the DIW process, these inks are extruded through a nozzle mounted on a 3D translation platform, with the flow regulated by the application of pressure . 2.3. Powder Bed Fusion The powder bed fusion (PBF) technique uses a laser or an electron beam to melt and fuse the material into powder . The principle used in this technique is to produce the product layer by layer and melt it. A heat source concentrates its heat on a powdered base material and heats the cross-sectional area . PBD is mainly used due to the low cost of producing the object and because the powder can be recycled to produce another piece . This method encompasses several commonly utilised processes, including direct metal laser sintering (DMLS), selective laser sintering (SLS), electron beam melting (EBM), and selective laser melting (SLM). PBF utilises a laser source or an electron beam (EBM) to selectively melt or sinter layers of material, thereby fabricating a solid component. This technique is applicable to a range of powder-based materials, with metals and polymers being the most commonly processed . 2.4. Material Jetting The material jetting (MJ) deposits the liquid in droplets to bind powder material . In the MJ, the material is injected into the surface/building platform, where it solidifies, and the model is assembled layer by layer. The layers are subsequently cured or toughened with ultraviolet light. In this method, the material must be deposited in droplets, and therefore the materials used are limited. Generally, polymers and waxes are used, considering their viscous nature and ability to produce droplets . This is a fast and proficient method and offers greater freedom when designing and printing complex models . 2.5. Binder Jetting In binder jetting (BJ), a liquid bonding compound is applied selectively to bind powder materials. A relevant feature of this technique today is the possibility of using colour inkjet technology to create coloured objects in the binders . Binders are used to ensure adhesion between the powdered material particles. These binders contribute to obtaining the strength of the part and the desired form of the final product . The materials commonly used in this method are metals (e.g., stainless steel), ceramics (e.g., glass) and polymers (e.g., acrylonitrile butadiene styrene, polyamide and polycarbonate) . 2.6. Directed Energy Deposition Directed energy deposition (DED) is an AM technique that utilises focused thermal energy to melt materials, which are fused as they are deposited. In this method, an energy source—such as an electron beam, a laser or a plasma—is used to melt the materials, which are then deposited . The DED involves a nozzle assembled on a multi-hub arm that deposits the dissolved material onto a surface layer, where it solidifies. A significant advantage of this method is its ability to precisely control the grain structure of the deposited material [ , , ]. This process is analogous to material extrusion; however, the nozzle in DED has the capacity to transverse multiple directions rather than being restricted to a single axis. While the process is applicable to polymers and ceramics, it is predominantly utilised for metals, which are provided in either powder or wire form . 2.7. Sheet Lamination Sheet lamination involves the sequential bonding of thin sheets of material, normally fed through a set of rollers . This technique can employ various materials, including paper, polymer and metal . Although this is a less precise method, it offers advantages in terms of speed and cost-effectiveness . The predominant sheet lamination techniques are laminated object manufacturing (LOM) and ultrasonic additive manufacturing (UAM). LOM employs a layer-by-layer method similar to other additive manufacturing processes, utilizing paper as the primary material and adhesive for bonding, rather than welding. Conversely, UAM involves metals such as aluminium, stainless steel and titanium, and operates at relatively low temperatures, enabling the production of complex internal geometries . To highlight the advantages and drawbacks of each of the previously discussed AM techniques, a summary is provided in . 

This method is characterized by the principle of selectively solidifying a liquid photopolymer resin by exposing it to specific light sources or patterns. In the procedure, a vat contains the liquid resin, which is selectively cured layer by layer in the desired regions, gradually constructing the intended object. The laser, guided by predetermined digital coordinates, provides the necessary information to cure the resin at precise locations within each layer. To facilitate this curing process, a photoinitiator is introduced into the resin material. The photoinitiator absorbs the incident ultraviolet (UV) radiation and generates active species that initiate the photopolymerization reaction . The vat photopolymerization technique includes various commonly utilised processes, such as stereolithography (SLA), continuous light interface production (CLIP), digital light processing (DLP) and two-photon polymerization (2PP) . These methods involve SLA, which uses a laser or digital light projector to cure the resin layer by layer, whereas DLP employs a digital light projector to cure an entire layer at once. CLIP employs a continuous liquid interface production approach, in which a continuous liquid interface is established between the resin and a transparent window, that enables continuous printing. Lastly, 2PP utilises a focused laser beam and a photosensitive resin to achieve high-resolution printing via a nonlinear optical process . These vat photopolymerization techniques offer versatile and AM capabilities for fabricating complex and detailed structures. 

Materials extrusion is a widely adopted AM technique due to its rapid manufacturing, cost efficiency, simplicity, user-friendly nature and the potential to produce complex components [ , , ]. This technique involves extruding the material through an orifice and depositing it onto a construction platform [ , , ]. The most common technique is fused filament fabrication (FFF), which is also referred to as fused deposition modelling (FDM) . In this method, the filament is heated to a temperature typically between 150 and 250°C until it melts, and the molten material is extruded through a nozzle . The extruded material is then placed onto a build plate, which can be adjusted in the z direction. This technique usually uses more polymers, although it is also used to print metal and ceramic components. FFF uses a support to create overhang features, which can be removed either mechanically, by detaching it from the printed part, or chemically, by dissolving it in a solvent . In contrast to FFF, which employs thermoplastics, direct ink writing (DIW) is a 3D printing technique that utilises viscoelastic inks. In the DIW process, these inks are extruded through a nozzle mounted on a 3D translation platform, with the flow regulated by the application of pressure . 

The powder bed fusion (PBF) technique uses a laser or an electron beam to melt and fuse the material into powder . The principle used in this technique is to produce the product layer by layer and melt it. A heat source concentrates its heat on a powdered base material and heats the cross-sectional area . PBD is mainly used due to the low cost of producing the object and because the powder can be recycled to produce another piece . This method encompasses several commonly utilised processes, including direct metal laser sintering (DMLS), selective laser sintering (SLS), electron beam melting (EBM), and selective laser melting (SLM). PBF utilises a laser source or an electron beam (EBM) to selectively melt or sinter layers of material, thereby fabricating a solid component. This technique is applicable to a range of powder-based materials, with metals and polymers being the most commonly processed . 

The material jetting (MJ) deposits the liquid in droplets to bind powder material . In the MJ, the material is injected into the surface/building platform, where it solidifies, and the model is assembled layer by layer. The layers are subsequently cured or toughened with ultraviolet light. In this method, the material must be deposited in droplets, and therefore the materials used are limited. Generally, polymers and waxes are used, considering their viscous nature and ability to produce droplets . This is a fast and proficient method and offers greater freedom when designing and printing complex models . 

In binder jetting (BJ), a liquid bonding compound is applied selectively to bind powder materials. A relevant feature of this technique today is the possibility of using colour inkjet technology to create coloured objects in the binders . Binders are used to ensure adhesion between the powdered material particles. These binders contribute to obtaining the strength of the part and the desired form of the final product . The materials commonly used in this method are metals (e.g., stainless steel), ceramics (e.g., glass) and polymers (e.g., acrylonitrile butadiene styrene, polyamide and polycarbonate) . 

Directed energy deposition (DED) is an AM technique that utilises focused thermal energy to melt materials, which are fused as they are deposited. In this method, an energy source—such as an electron beam, a laser or a plasma—is used to melt the materials, which are then deposited . The DED involves a nozzle assembled on a multi-hub arm that deposits the dissolved material onto a surface layer, where it solidifies. A significant advantage of this method is its ability to precisely control the grain structure of the deposited material [ , , ]. This process is analogous to material extrusion; however, the nozzle in DED has the capacity to transverse multiple directions rather than being restricted to a single axis. While the process is applicable to polymers and ceramics, it is predominantly utilised for metals, which are provided in either powder or wire form . 

Sheet lamination involves the sequential bonding of thin sheets of material, normally fed through a set of rollers . This technique can employ various materials, including paper, polymer and metal . Although this is a less precise method, it offers advantages in terms of speed and cost-effectiveness . The predominant sheet lamination techniques are laminated object manufacturing (LOM) and ultrasonic additive manufacturing (UAM). LOM employs a layer-by-layer method similar to other additive manufacturing processes, utilizing paper as the primary material and adhesive for bonding, rather than welding. Conversely, UAM involves metals such as aluminium, stainless steel and titanium, and operates at relatively low temperatures, enabling the production of complex internal geometries . To highlight the advantages and drawbacks of each of the previously discussed AM techniques, a summary is provided in . 

According to Moroni et al., in the context of tissue engineering and regenerative medicine, the term “bioprinting” refers to the use of computer-assisted transfer techniques for the modelling and assembly of living and non-living materials in a predefined 2D or 3D configuration to create bioengineered structures . By incorporating active substances, such as biomaterials, living cells and bioactive molecules, bioprinting imparts biological functions to the manufactured structures . Bioprinting technology allows for the manufacture of complex, functional structures that promote cell growth and tissue formation. The prospect of manufacturing complete tissues or organs using 3D printing is very promising and has the potential to revolutionise regenerative medicine . In recent decades, significant advancements have been made in the field of 3D bioprinting ( ) [ , , , ]. The evolution of this technology began with the invention of SLA by Charles Hull in 1984, which marked the beginning of 3D bioprinting . Subsequently, in 2002, Landers et al. introduced extrusion-based bioprinting technology, which was later commercialized under the name “3D-Bioplotter” . In 2003, Thomas and Boland’s research team adapted a conventional inkjet printer to develop the first inkjet bioprinter capable of printing living cells . Later, the engineering of scaffold-free vascular tissue via bioprinting was achieved by Norotte et al. . The subsequent years witnessed the development and introduction of various bioprinted constructs, including an artificial liver in 2012, full human skin in 2014, a heart valve in 2016, and a lung-mimicking air sac with surrounding blood vessels in 2019, among other innovations [ , , , ]. 3.1. Techniques Among the diverse AM techniques, the most employed in bioprinting are laser-based printing (SLA and SLS), extrusion printing and inkjet printing [ , , ]. SLA is a vat photopolymerization process that consists of photocurable bioinks that are subjected to UV, infrared or visible light to produce 3D pieces using the layer-by-layer procedure . Among the advantages of SLA is its capacity to rapidly cure at physiological temperatures, which facilitates the production of constructs suitable for regenerative medicine applications. This technique has been widely employed to fabricate micro-needles designed for transdermal drug delivery, hearing aids, surgical guides for placing dental implants, temporary crowns and bridges and supports for tissue engineering with/without encapsulated cells . During SLS printing, fine particles of the entire substance are fused by the heat of a high-powered laser to fabricate a 3D structure . In this technique, several categories of powders, including polymers, ceramics and metals, must be processed into powder. SLS is a PBF process that is used in numerous applications in the medical field, namely the fabrication of prototypes for medical devices, physical models used in surgery and scaffolds for tissue engineering. The popularity of SLS printers is attributed to their affordability, high productivity and material versatility . Extrusion printing is considered the most popular bioprinting technology . In this method, material is melted and extruded through a nozzle, orifice or needle using a screw, piston or high-pressure pneumatic force to form successive layers of the part [ , , , , ]. This technique has been used extensively in the medical field, enabling the biofabrication of tissues, organs, implants and personalised drug delivery systems. Extrusion is also highly applicable in the field of disease modelling. Models made by extrusion can provide a baseline for the comprehension of the underlying biological mechanisms behind disease progression, thus contributing to the identification of effective treatments . Inkjet bioprinting is considered the pioneering technology in the field of bioprinting. The printing process utilizing this technique comprises two phases: first, the generation of discrete droplets that are precisely directed to specific locations on a substrate, and second, the subsequent interaction between these droplets and the substrate . It is employed in various medical applications, including the creation of patient-specific or project-specific implants for static load-bearing purposes (such as dental crowns and prosthetic structures), joint applications (such as osteochondral cartilage implants) and for facilitating in vivo blood vessel formation and tissue regeneration . presents the aforementioned techniques, as well as the advantages and drawbacks of each. 3.2. Materials Used in 3D Bioprinting Nowadays, a diverse range of biomaterials are being used in 3D bioprinting . The interaction between biomaterials and cells is fundamental for cell viability, proliferation, and differentiation. Hence, it is fundamental to consider the characteristics of biomaterials, such as non-toxicity, biocompatibility, and the absence of immune reactions and foreign body responses . Bioprinting materials are categorized into two types: natural and synthetic biomaterials. Natural biomaterials are particularly attractive due to their bioactivity, being both similar to the extracellular matrix (ECM) and biocompatible. The following are some examples of natural materials used in this field: collagen, xanthan gum (XG), silk fibroin (SF), gelatin, pectin, gellan gum, albumin, chitosan, sodium alginate, agarose, fibrin, keratin and hyaluronic acid (HA). However, these types of materials generally exhibit poor mechanical properties [ , , , , , , , , , , , , , , , , , , , , , ]. Conversely, synthetic materials provide distinct advantages compared to natural materials, including the ability to be customized with specific physical properties and greater uniformity. The following are some examples of synthetic materials used in this field: poly (ℇ-caprolactone) (PCL), poly(lactic-co-glycolic) acid (PLGA), poly(l-lactic) acid (PLA), poly (glycolic acid) (PGA), polyurethane (PU), polyethylene glycol (PEG), polyether ether ketone (PEEK), polyvinylpyrrolidone (PVP) and pluronic. Nevertheless, synthetic materials for 3D bioprinting have some disadvantages, such as poor biocompatibility, the potential release of toxic degradation byproducts and the absence of bioactive ligands [ , , , , , , , , , , , , , ]. To address these drawbacks, a more comprehensive knowledge of the physiological characteristics of the ECM and a higher ability to replicate the complex 3D structures mimetic of the ECM would represent a significant advance in 3D bioprinting. Advances in composite and hybrid bioprinting materials, as well as multimaterial bioprinting technologies, have emerged in the field of tissue regeneration . Nonetheless, human tissues and organs operate within a highly dynamic biochemical environment, and conventional bioprinted materials often fail to adapt to the evolving spatial and temporal requirements of tissue and organ development. Consequently, there is increasing interest in developing sophisticated stimuli-responsive biomaterials for 3D bioprinting, which would facilitate the adaptation of artificial constructs to intricate and dynamic physiological conditions [ , , ]. 

Among the diverse AM techniques, the most employed in bioprinting are laser-based printing (SLA and SLS), extrusion printing and inkjet printing [ , , ]. SLA is a vat photopolymerization process that consists of photocurable bioinks that are subjected to UV, infrared or visible light to produce 3D pieces using the layer-by-layer procedure . Among the advantages of SLA is its capacity to rapidly cure at physiological temperatures, which facilitates the production of constructs suitable for regenerative medicine applications. This technique has been widely employed to fabricate micro-needles designed for transdermal drug delivery, hearing aids, surgical guides for placing dental implants, temporary crowns and bridges and supports for tissue engineering with/without encapsulated cells . During SLS printing, fine particles of the entire substance are fused by the heat of a high-powered laser to fabricate a 3D structure . In this technique, several categories of powders, including polymers, ceramics and metals, must be processed into powder. SLS is a PBF process that is used in numerous applications in the medical field, namely the fabrication of prototypes for medical devices, physical models used in surgery and scaffolds for tissue engineering. The popularity of SLS printers is attributed to their affordability, high productivity and material versatility . Extrusion printing is considered the most popular bioprinting technology . In this method, material is melted and extruded through a nozzle, orifice or needle using a screw, piston or high-pressure pneumatic force to form successive layers of the part [ , , , , ]. This technique has been used extensively in the medical field, enabling the biofabrication of tissues, organs, implants and personalised drug delivery systems. Extrusion is also highly applicable in the field of disease modelling. Models made by extrusion can provide a baseline for the comprehension of the underlying biological mechanisms behind disease progression, thus contributing to the identification of effective treatments . Inkjet bioprinting is considered the pioneering technology in the field of bioprinting. The printing process utilizing this technique comprises two phases: first, the generation of discrete droplets that are precisely directed to specific locations on a substrate, and second, the subsequent interaction between these droplets and the substrate . It is employed in various medical applications, including the creation of patient-specific or project-specific implants for static load-bearing purposes (such as dental crowns and prosthetic structures), joint applications (such as osteochondral cartilage implants) and for facilitating in vivo blood vessel formation and tissue regeneration . presents the aforementioned techniques, as well as the advantages and drawbacks of each. 

Nowadays, a diverse range of biomaterials are being used in 3D bioprinting . The interaction between biomaterials and cells is fundamental for cell viability, proliferation, and differentiation. Hence, it is fundamental to consider the characteristics of biomaterials, such as non-toxicity, biocompatibility, and the absence of immune reactions and foreign body responses . Bioprinting materials are categorized into two types: natural and synthetic biomaterials. Natural biomaterials are particularly attractive due to their bioactivity, being both similar to the extracellular matrix (ECM) and biocompatible. The following are some examples of natural materials used in this field: collagen, xanthan gum (XG), silk fibroin (SF), gelatin, pectin, gellan gum, albumin, chitosan, sodium alginate, agarose, fibrin, keratin and hyaluronic acid (HA). However, these types of materials generally exhibit poor mechanical properties [ , , , , , , , , , , , , , , , , , , , , , ]. Conversely, synthetic materials provide distinct advantages compared to natural materials, including the ability to be customized with specific physical properties and greater uniformity. The following are some examples of synthetic materials used in this field: poly (ℇ-caprolactone) (PCL), poly(lactic-co-glycolic) acid (PLGA), poly(l-lactic) acid (PLA), poly (glycolic acid) (PGA), polyurethane (PU), polyethylene glycol (PEG), polyether ether ketone (PEEK), polyvinylpyrrolidone (PVP) and pluronic. Nevertheless, synthetic materials for 3D bioprinting have some disadvantages, such as poor biocompatibility, the potential release of toxic degradation byproducts and the absence of bioactive ligands [ , , , , , , , , , , , , , ]. To address these drawbacks, a more comprehensive knowledge of the physiological characteristics of the ECM and a higher ability to replicate the complex 3D structures mimetic of the ECM would represent a significant advance in 3D bioprinting. Advances in composite and hybrid bioprinting materials, as well as multimaterial bioprinting technologies, have emerged in the field of tissue regeneration . Nonetheless, human tissues and organs operate within a highly dynamic biochemical environment, and conventional bioprinted materials often fail to adapt to the evolving spatial and temporal requirements of tissue and organ development. Consequently, there is increasing interest in developing sophisticated stimuli-responsive biomaterials for 3D bioprinting, which would facilitate the adaptation of artificial constructs to intricate and dynamic physiological conditions [ , , ]. 

Cellular therapies, also known as cell therapy, cell transplantation or cytotherapy, involve the injection, grafting or implantation of cells into a patient (autologous or allogeneic) to achieve a medicinal effect . Cell therapies are frequently applied in combination with biomaterial supports, which are designed to support and guide the cells both during and after transplantation . Advances in 3D printing technology allow for the precise spatial arrangement of multiple cells and biomaterials, thereby enhancing the efficiency of cell delivery and integration onto scaffolds. There are two primary methods to deliver cells via 3D printed scaffolds: post-printing cell seeding and embedding cells within bioinks during the 3D bioprinting process [ , , , ]. The approach of seeding the cells after printing provides more flexible requirements for both the conditions and 3D printing process, though it is hindered by low cell adhesion rates on the scaffolds. To address this challenge, one approach involves encapsulating cells in hydrogels prior to combining them with 3D printed scaffolds. Alternatively, the bioprinting method enhances cell loading efficiency and allows for precise control over the spatial distribution of various cell types. However, this method demands stringent conditions and precise parameters during printing, as cells are sensitive to suboptimal environmental conditions. This necessitates detailed knowledge of the parameters of the bioinks and printing conditions beforehand [ , , , , ]. The selection of a specific type of cell therapy is highly important, since it affects the function and design of the tissue engineering model . There are three categories of cell therapies that can be applied to printed scaffolds, namely, stem cell-based, non-stem cell-based and multicellular therapies . 4.1. Stem Cell-Based Therapies Stem cells are among the most frequently utilised cell types due to their unspecialized nature, which allows them to self-renew and differentiate into various cell lineages [ , , , ]. Their ability to differentiate into specific cell types while continuously dividing and self-renewing makes their prospects interesting for regenerative medicine. They have successfully been used to create functional tissues that replicate the properties of natural organs. These stem cells can be sourced from several origins . In the field of bioengineering, the three most frequently employed types are mesenchymal stem/stromal cells (MSCs), embryonic stem cells (ESCs) and induced pluripotent stem cells (iPSCs) . Also known as multipotent cells, MSCs can be sourced from various tissues, including bone marrow, muscle, lung, teeth, adipose tissue, liver and perinatal/extra-embryonic tissues. They exhibit the capacity to proliferate and differentiate into a broad range of cell types, such as osteoblasts, chondrocytes and adipocytes . Additionally, MSCs can differentiate into other mesenchymal and non-mesenchymal cells, including myocytes, tendocytes, neural cells, ligament cells, smooth muscle cells, endothelial cells, cardiomyocytes and hepatocytes . In regenerative medicine, these cells offer numerous advantages, including ease of expansion in culture and the capacity to differentiate into desired cell lines. They also possess specific immunological properties, such as being immunoprivileged and immunomodulatory, and have tropisms for injury sites. Additionally, they can stimulate trophic responses and modulate tissue functions and inflammation through the secretion of essential bioactive molecules . However, MSCs present some possible challenges, such as poor-quality control and inconsistencies regarding their heterogeneity, stability, differentiation, immunocompatibility, and migratory capacity . In turn, ESCs are pluripotent cells originating from the inner cell mass of blastocytes (an early embryo). They possess the capability to differentiate into nearly all cell types originating from the three germ layers, with the exception of trophoblastic cells . However, ESCs raise ethical concerns, exhibit challenges in controlling differentiation, and have the potential to form teratomas and provoke immune responses [ , , , , ]. Finally, iPSCs are pluripotent stem cells generated by reprogramming adult somatic cells through the introduction of specific genes and factors, thereby mimicking the pluripotent state of ESCs . While iPSCs share many characteristics with ESCs, they do not carry the same ethical issues or immunogenicity concerns. Nonetheless, iPSCs, such as ESCs, are also associated with the risk of teratoma formation in vivo . 4.2. Non-Stem Cell-Based Therapies Non-stem cell therapies commonly use somatic cells that are isolated from humans, subsequently cultured and expanded in vitro and then applied to patients for therapeutic, preventive or diagnostic treatment . These can be categorized into immune cells and non-immune cells. Immune cells, including natural killer cells, dendritic cells and macrophages, can be engineered to target specific antigens. This approach is used in therapeutic strategies, including for cancer, infections, autoimmune diseases and allogeneic transplantation. On the other hand, non-immune cells, such as chondrocytes, fibroblasts, hepatocytes, keratinocytes and pancreatic islet cells, normally are involved in the host’s defence response as structural architectures, regulators and effectors of its protective immune reaction [ , , , , , ]. These somatic cell-based therapies serve as in vivo resource of cytokines, enzymes and growth factors. Additionally, they are frequently used in adoptive cell therapy for cancer treatment and as transplanted cells (e.g., hepatocytes or pancreatic islet cells) to address genetic metabolic disorders. They also find applications as cell-based or scaffold-free systems in the treatment of burns, ulcers and cartilage injuries [ , , ]. 4.3. Multicellular Therapies Multicellular therapies comprise at least two types of cultured stem and/or non-stem cells. This emerging approach holds that using a combination of cell types is more effective for promoting long-term tissue repair compared to single-cell therapy. This effectiveness is attributed to cell–cell interactions that extend beyond embryogenesis and play a crucial role in regenerative procedures [ , , ]. Multicellular therapies include scaffold-supported and scaffold-free cell products, bone marrow aspirate-derived therapies, adoptive cell therapy products, stem cell transplantation and stromal vascular fraction [ , , , , ]. 

Stem cells are among the most frequently utilised cell types due to their unspecialized nature, which allows them to self-renew and differentiate into various cell lineages [ , , , ]. Their ability to differentiate into specific cell types while continuously dividing and self-renewing makes their prospects interesting for regenerative medicine. They have successfully been used to create functional tissues that replicate the properties of natural organs. These stem cells can be sourced from several origins . In the field of bioengineering, the three most frequently employed types are mesenchymal stem/stromal cells (MSCs), embryonic stem cells (ESCs) and induced pluripotent stem cells (iPSCs) . Also known as multipotent cells, MSCs can be sourced from various tissues, including bone marrow, muscle, lung, teeth, adipose tissue, liver and perinatal/extra-embryonic tissues. They exhibit the capacity to proliferate and differentiate into a broad range of cell types, such as osteoblasts, chondrocytes and adipocytes . Additionally, MSCs can differentiate into other mesenchymal and non-mesenchymal cells, including myocytes, tendocytes, neural cells, ligament cells, smooth muscle cells, endothelial cells, cardiomyocytes and hepatocytes . In regenerative medicine, these cells offer numerous advantages, including ease of expansion in culture and the capacity to differentiate into desired cell lines. They also possess specific immunological properties, such as being immunoprivileged and immunomodulatory, and have tropisms for injury sites. Additionally, they can stimulate trophic responses and modulate tissue functions and inflammation through the secretion of essential bioactive molecules . However, MSCs present some possible challenges, such as poor-quality control and inconsistencies regarding their heterogeneity, stability, differentiation, immunocompatibility, and migratory capacity . In turn, ESCs are pluripotent cells originating from the inner cell mass of blastocytes (an early embryo). They possess the capability to differentiate into nearly all cell types originating from the three germ layers, with the exception of trophoblastic cells . However, ESCs raise ethical concerns, exhibit challenges in controlling differentiation, and have the potential to form teratomas and provoke immune responses [ , , , , ]. Finally, iPSCs are pluripotent stem cells generated by reprogramming adult somatic cells through the introduction of specific genes and factors, thereby mimicking the pluripotent state of ESCs . While iPSCs share many characteristics with ESCs, they do not carry the same ethical issues or immunogenicity concerns. Nonetheless, iPSCs, such as ESCs, are also associated with the risk of teratoma formation in vivo . 

Non-stem cell therapies commonly use somatic cells that are isolated from humans, subsequently cultured and expanded in vitro and then applied to patients for therapeutic, preventive or diagnostic treatment . These can be categorized into immune cells and non-immune cells. Immune cells, including natural killer cells, dendritic cells and macrophages, can be engineered to target specific antigens. This approach is used in therapeutic strategies, including for cancer, infections, autoimmune diseases and allogeneic transplantation. On the other hand, non-immune cells, such as chondrocytes, fibroblasts, hepatocytes, keratinocytes and pancreatic islet cells, normally are involved in the host’s defence response as structural architectures, regulators and effectors of its protective immune reaction [ , , , , , ]. These somatic cell-based therapies serve as in vivo resource of cytokines, enzymes and growth factors. Additionally, they are frequently used in adoptive cell therapy for cancer treatment and as transplanted cells (e.g., hepatocytes or pancreatic islet cells) to address genetic metabolic disorders. They also find applications as cell-based or scaffold-free systems in the treatment of burns, ulcers and cartilage injuries [ , , ]. 

Multicellular therapies comprise at least two types of cultured stem and/or non-stem cells. This emerging approach holds that using a combination of cell types is more effective for promoting long-term tissue repair compared to single-cell therapy. This effectiveness is attributed to cell–cell interactions that extend beyond embryogenesis and play a crucial role in regenerative procedures [ , , ]. Multicellular therapies include scaffold-supported and scaffold-free cell products, bone marrow aspirate-derived therapies, adoptive cell therapy products, stem cell transplantation and stromal vascular fraction [ , , , , ]. 

Advancements in combining 3D printing/bioprinting and cellular therapies hold considerable promise for advancing regenerative medicine. This synergy has the potential to enhance tissue engineering capabilities and introduce novel therapeutic approaches for a range of diseases and conditions. Such advancements could substantially improve patient outcomes and have the potential to significantly impact survival rates [ , , ]. This approach has been successful in printing the structures of different tissues, including cardiovascular, bone, liver, skin and neural tissues ( ). The following sections analyse the advancements in the integration of 3D printing/bioprinting and cellular therapies in several regenerative medicine applications. Over recent years, research concerning 3D printing and bioprinting technologies in regenerative medicine has grown significantly. As indicated in , the number of publications related to bone regeneration is noticeably higher than the other fields. This prominence can be attributed to the compatibility of bone with 3D printing technologies, which allow precise control of scaffold porosity and the development of patient-specific implants. Furthermore, the high clinical demand for bone regeneration solutions, the limitations of traditional grafting techniques, and advancements in clinical translation, including regulatory approvals of 3D-printed bone supports, have significantly contributed to the extensive research focus in this area [ , , ]. 5.1. Cardiovascular Tissue Engineering Cardiovascular diseases, which include pathologies affecting the myocardium, heart valves and body’s vasculature, are highly prevalent worldwide. These diseases are a leading cause of morbidity and mortality, especially in developed countries [ , , , ]. Current therapeutic approaches include cellular therapies, bypass grafting, the implantation of medical devices, cardiac tissue patches, and organ transplantation [ , , , , ]. Organ transplantation is often not the optimal solution due to the imbalance between the availability of donor organs and the high demand. Additionally, the success of organ transplants is frequently compromised by complications related to immune rejection . Among the various therapeutic approaches, cell therapy has demonstrated success in regenerating cardiovascular tissue. Nevertheless, the absence of ECM limits cell survival following injection, resulting in reduced long-term viability . To address these concerns, 3D printing and bioprinting have emerged as effective approaches for developing scaffolds that incorporate ECM components and enhance cell viability . These structures can more accurately mimic the spatial and mechanical properties of native tissues, which is relevant for their functionality and integration in vivo . Currently, scaffolds for cardiovascular tissue engineering have been fabricated using various 3D printing technologies, such as inkjet printing, SLA and extrusion-based techniques. These scaffolds commonly incorporate biocompatible materials of natural origin, including fibrin, alginate, gelatin, collagen, HA and fibrinogen . However, to achieve complex structures with optimal physical, chemical and mechanical properties, synthetic materials, including poly (glycerol sebacate) (PGS), PCL and poly (ethylene glycol) methacrylate (PEGMA), are also employed. Additionally, some studies utilise a combination of natural and synthetic materials or semi-synthetic derivatives, such as gelatin methacrylate (GelMA) and hyaluronic acid-gelatin methacrylate (HAGM), to optimize scaffold properties while maintaining cell viability . Furthermore, decellularized matrices are extensively utilised due to their provision of a porous, interconnected polymeric network that facilitates cell migration, proliferation and the delivery of essential nutrients for cell survival . Maiullari et al. developed a technique for fabricating 3D cardiac tissue models that include a vascular network ( A). They created multi-cellular constructs using human umbilical vein endothelial cells (hUVECs) and induced pluripotent stem cell-derived cardiomyocytes (iPSC-CMs), encapsulated in alginate and PEG-Fibrinogen (PF). The constructs were fabricated using a custom-designed microfluidic printhead operating at an extrusion speed of 180 mm/min. Their research demonstrated that bioprinted endothelial cells are capable of forming functional vascular structures within the transplanted tissues and interacting with the host’s existing vascular network . To address damaged myocardium, Beijleri et al., produced a 3D-printed cardiac patch consisting of a decellularized cardiac extracellular matrix (cECM) hydrogel combined with GelMA. This patch was designed for the delivery of paediatric human cardiac progenitor cells (hCPC) ( B). The GelMA-cECM bioinks ensure uniform distribution of cECM and hCPCs, with the hCPCs maintaining over 75% viability. Additionally, conditioned media from GelMA-cECM patches demonstrate more than a 2-fold increase in angiogenic potential. The patches also remain adhered to rat hearts and show vascularization over a 14-day period in vivo . Also, in the development of tissue patches, Melhem et al. proposed a hydrogel patch embedded with multiple microchannels to enhance cell retention and factor delivery at the target tissue ( C). They integrated bone marrow-derived MSCs (BMSCs) into a hydrogel by cross-linking a poly (ethylene glycol) dimethacrylate (PEGDMA) solution containing the cells. Microchannels with precise diameters were created within the cell-loaded hydrogel using an SLA unit for in situ cross-linking. This 3D-printed, microchanneled hydrogel was designed as an advanced therapeutic tool for sustained delivery of multiple therapeutics, aiming to improve outcomes in ischemic heart injury . Noor et al. established a technique to 3D print thick, vascularized cardiac patches tailored to a patient’s specific properties. They used patient-derived cells, reprogrammed into cardiomyocytes and endothelial cells, and a personalized hydrogel made from the patient’s extracellular matrix. These components were combined to create bioinks for printing cardiac tissue and blood vessels. The patches, optimized for oxygen transfer, demonstrated proper structure and function in vitro, and the approach was further validated by successfully printing cellularised human hearts presenting a natural architecture ( D) . In turn, to mimic human microvasculature, Cui et al. developed a bioink combining human microvascular endothelial cells (hMVECs) and fibrin, used to fabricate micron-sized fibrin channels via drop-on-demand polymerization ( E). This aqueous-based printing method reduces cellular damage. The hMVECs printed with fibrin aligned within the channels and proliferated to form continuous linings, resulting in a three-dimensional tubular structure. The study concluded that the concurrent printing of cells and scaffolds enhances the proliferation of hMVECs and supports the formation of microvascular networks . Several researchers have employed 3D printing or bioprinting, with or without cell therapies, to advance cardiovascular tissue engineering. provides a summary of these studies. 5.2. Bone Tissue Engineering Severe bone defects resulting from aging, trauma, osteoporosis, degenerative diseases, autoimmune diseases (such as rheumatoid arthritis) or tumour removal are a leading cause of disability globally, affecting an estimated 1.71 billion people . Current therapeutic strategies for addressing these defects include autografting, allografting, xenografting and bone transplantation. Despite their application, these methods present significant risks, including the potential for infectious disease transmission and immune rejection . To address the limitations of traditional bone defect treatments, 3D printing and bioprinting techniques, with or without cellular therapies, have been developed for bone tissue engineering. These methods enable the large-scale fabrication of custom-tailored bone tissues, meeting the growing demand for functionalized bone implants . The 3D-printed scaffolds serve as a support for cell growth and differentiation, forming a hierarchical bone microvascular architecture . For optimal performance, a scaffold should possess several critical attributes: biocompatibility, sterility, osteoconductivity, biodegradability and a porous, interconnected structure that supports cellular infiltration and nutrient transport. Additionally, it should effectively repair bone defects while closely mimicking the characteristics of native bone tissue [ , , ]. The most commonly employed printing techniques in bone tissue engineering are laser powder bed fusion, vat photopolymerization and extrusion-based methods. Various biomaterials are utilised for printing scaffolds: ceramics (e.g., beta-tricalcium phosphate (β-TCP), hydroxyapatite (HAp) and amorphous calcium phosphate (ACP)); natural polymers (e.g., matrigel, alginate, HA and dextran emulsion); synthetic polymers (e.g., GelMA, PLGA, PCL and polyethylene glycol diacrylate (PEGDA); metals (e.g., titanium alloy (Ti6Al4V), tantalum (Ta) and titanium (Ti)), as well as the combinations of these materials. Lei et al. developed Ti6Al4V-based porous tantalum (Ta) scaffolds with high interfacial strength using laser powder bed fusion. In this process, porous Ta was directly deposited onto a solid Ti6Al4V substrate ( A). The manufacturing process was carried out in an argon atmosphere, using a 300 W laser at 250 mm/s for Ta and a 250 W laser at 300 mm/s for Ti6Al4V. In vitro biocompatibility assessments conducted with rat bone marrow mesenchymal stem cells (r-BMSCs) confirmed the scaffolds’ biocompatibility. The findings demonstrated the strong mechanical compatibility and osteointegration properties of the Ti6Al4V-based porous Ta scaffold, underscoring its considerable potential for orthopaedic applications . For mandibular bone defect reconstruction, Yu et al. encapsulated BMSCs in matrigel and infiltrated this mixture into porous Ti6Al4V scaffolds. The study demonstrated that rats with critical full-thickness mandibular defects treated with Matrigel-infused Ti6Al4V scaffolds exhibited significantly greater new bone formation compared to those treated with either local BMSC injections or Matrigel alone ( B). These results indicate that Matrigel enhances the 3D microenvironment for BMSCs, positioning Matrigel-infused scaffolds as a promising method for improving bone regeneration in 3D-printed Ti6Al4V scaffolds . Wu et al. produced a 3D-bioprintable scaffold combining alginate and β-TCP for the treatment of bone defects ( C). The scaffolds were fabricated at a printing speed of 2 mm/s, with an extrusion pressure ranging from 60 to 200 kPa. MG-63 cells were seeded onto these scaffolds. The 3D-printed scaffolds using a 10% alginate/β-TCP bioink exhibited enhanced physical characteristics and significantly improved cell viability and alkaline phosphatase activity. These findings suggest that the scaffolds have considerable potential for application in personalized bone regeneration therapies . To produce a scaffold that replicates bone microstructure, Ressler et al. developed trabecular-like porous scaffolds using ceramic vat photopolymerization with HAp powders doped with magnesium (Mg 2+ ), strontium (Sr 2+ ) and zinc (Zn 2+ ). Scaffolds sintered at 1100–1300 °C exhibited mechanical properties similar to trabecular bone, with optimal performance at 1300 °C. The microstructure resembled cancellous bone and the incorporation of trace elements resulted in a biphasic calcium phosphate system (HAp/β-TCP), potentially enhancing bioactivity . Numerous studies have utilised 3D printing or bioprinting, either with or without cellular therapies, to advance bone tissue engineering. An overview of these studies is provided in . 5.3. Liver Tissue Engineering The liver plays a vital role in blood protein synthesis, glucose metabolism and the detoxification of metabolites . It is also the only organ in the human body capable of efficient regeneration. However, this regenerative capacity can be compromised by excessive drug use or viral infections, which can cause irreversible damage to hepatocytes and lead to liver failure . Chronic liver diseases, including fibrosis, cirrhosis, chronic viral hepatitis and fatty liver disease, significantly contribute to global morbidity and mortality. Unfortunately, advancements in treatment options for these conditions remain limited . The primary medical intervention for liver failure is partial or total liver transplantation. Nevertheless, this approach faces challenges, including limited donor availability, immune rejection and variable graft success rates. Alternative approaches in tissue engineering encompass bioartificial liver systems, which involve the in vitro creation of liver tissue to repair or replace damaged liver segments, as well as hepatocyte transplantation and cellular therapy methods . There has been a continuous search for a reliable and reproducible source of hepatocytes, whether for liver regeneration therapy, seeding liver support devices or in vitro screening applications . Numerous studies have investigated the 3D bioprinting of liver tissue utilizing either stem cells or immortalized hepatic cell lines . Stem cells are particularly promising due to their ability to express hepatocyte-like phenotypes. In contrast, adult hepatocytes are scarce, challenging to isolate, exhibit poor propagation and experience rapid functional deterioration in vitro . In liver tissue engineering, a variety of biomaterials are employed, categorized into natural and synthetic polymers . Natural polymers, such as alginate, HA, collagen, cellulose nanocrystal (CNC) and gelatin, present the advantages of enhanced cell compatibility and ease of manipulation. Nevertheless, they are constrained by relatively weaker mechanical properties, limited availability and variable degradation rates. Consequently, synthetic polymers, such as PCL, have been developed to provide superior mechanical strength, flexibility, processability and adjustable degradability. Despite these advantages, synthetic polymers generally lack cell recognition and adhesion sites, resulting in reduced biocompatibility compared to natural polymers . To mitigate these issues, a strategy of combining natural and synthetic polymers is employed to create effective bioinks for 3D bioprinting. Additionally, the liver decellularized extracellular matrix (dECM), sourced from animals, is frequently used to establish microenvironments for liver cells, offering cross-species tolerance and minimising the risk of immune rejection [ , , ]. Yang et al. developed a liver tissue model through 3D bioprinting using HepaRG cells, a widely utilised hepatic progenitor cell line. Each 3D bioprinted hepatorganoid was produced through forced extrusion at a rate of 150 mm 3 /min, using a layer-by-layer approach. This 3D bioprinted model, referred to as hepatorganoids, exhibited essential liver functions, including albumin production, drug metabolism and glycogen accumulation, after a 7-day differentiation period ( A). In vivo studies showed that the 3D bioprinted hepatorganoids further matured, exhibiting enhanced synthesis of liver-specific proteins and more human-like drug metabolism. Notably, transplantation of 3D bioprinted hepatorganoids significantly increased survival rates in recipient mice. These findings suggest that 3D bioprinted hepatorganoids can effectively undergo hepatic differentiation and ameliorate liver failure in vivo . In the study conducted by Xie et al., a 3D bioprinted model of hepatocellular carcinoma, derived from patient cells, was created using isolated primary hepatocellular carcinoma cells mixed with gelatin and sodium alginate to form a bioink ( B). The resulting models were successfully generated and exhibited substantial growth over prolonged culture durations. They preserved key characteristics of the original hepatocellular carcinoma tumours, including consistent biomarker expression and stable genetic alterations and expression profiles. Thus, 3D bioprinted hepatocellular carcinoma models prove to be reliable in in vitro systems, suitable for long-term culture and capable of predicting patient-specific drug responses for tailored therapeutic approaches . Lewis et al. studied a technique for 3D printing gelatin into precisely defined geometries, which exhibit distinct biological effects on seeded hepatocytes ( C). Their research reveals that the structural configuration of gelatin can markedly impact biological processes. An undifferentiated hepatocyte cell line demonstrated high viability and proliferation on 3D-printed scaffolds with two distinct geometries. Notably, hepatocyte-specific functions—such as albumin secretion, cytochrome P450 activity and bile transport—were enhanced in more interconnected 3D-printed gelatin structures compared to those with less interconnectivity and traditional two-dimensional (2D) cultures. This study underscores the gap between gene expression and protein functionality in simplistic 2D cultures, highlighting the importance of a physiologically relevant 3D environment for optimizing hepatocyte expression and functionality . Jeon et al. employed 3D bioprinting technology to reconstruct liver tissues and organs using human hepatocellular carcinoma (HepG2) cells, a liver cancer-derived cell line. They created multi-layered 3D structures by integrating alginate with HepG2 cells ( D). The study demonstrated that replicating the 3D hepatic architecture using this technology enhances the stability and gene expression profiles of HepG2 cells. Cells cultured on these 3D alginate scaffolds for three weeks were analysed via fluorescence microscopy, histology and immunohistochemistry. Results indicated that HepG2 cells exhibited improved growth and liver-specific gene expression in 3D cultures compared to 2D cultures, highlighting the effectiveness of 3D bioprinting in mimicking liver architecture and enhancing cellular function . Using the same 3D bioprinting technique, Wu et al. produced a novel bioink containing alginate, CNC and GelMA (namely, 135ACG hybrid ink), aimed at fabricating both cell-laden and acellular structures ( E). The structures were printed using a 32-gauge nozzle at a printing speed of 20 mm/s and an air pressure of 20 psi. The bioink presented good shear-thinning behaviour and solid-like characteristics, ensuring high printability and minimal cell damage. Following crosslinking, it formed a rigid ECM conducive to stromal cell growth. The team engineered a GelMA bioink with suitable mechanical properties to mimic human liver tissue, enabling the printing of liver lobule-mimetic constructs with precise cell placement (fibroblasts and hepG2) in different ECMs (135ACG and GelMA). These constructs were used to study the impact of mechanical stimuli and cellular interactions on cell behaviour. These findings demonstrated that fibroblasts proliferated effectively within the rigid 135ACG matrix, whereas HepG2 cells developed into spheroid structures in the more compliant GelMA matrix. Co-cultures of hepG2 and fibroblasts cells showed increased albumin production, highlighting the role of soluble factors in enhancing hepatic function. The study demonstrated that the developed bioinks and printing methods are effective for creating complex, multi-cellular constructs with varied ECMs, advancing both fundamental research and tissue engineering applications . Several studies have employed 3D printing/bioprinting with or without the incorporation of cellular therapies to advance liver tissue engineering. A summary of these studies is presented in . 5.4. Skin Tissue Engineering The skin, regarded as the largest organ in the human body, plays essential roles in various functions, including serving as a protective barrier, regulating body temperature and preventing dehydration . Extensive full-thickness skin wounds, which damage underlying blood vessels, pose serious risks due to induced cellular hypoxia and nutrient deprivation . While autografts are considered the “gold standard” for treating severe skin injuries, their use is limited by issues such as donor site availability and associated morbidity . Furthermore, existing commercial skin substitutes lack the sufficient vascular networks necessary for effective nutrient delivery in full-thickness wounds. Consequently, skin substitutes are seen as promising alternatives, offering the potential for vascularized skin reconstruction with customized cell compositions and controlled geometrical structures . 3D printing/bioprinting has arisen as an innovative technological approach in skin tissue engineering, creating structures by depositing cell-embedded bioinks layer by layer . This approach includes several techniques, including DLP, extrusion bioprinting, inkjet printing and electrospinning. The materials commonly used in this field are natural polymers (e.g., collagen, alginate, HA, gelatin and fibroin) and synthetic polymers (e.g., PCL, PLGA, polyglycolic acid, polyurethanes, polycarbonates and PEGDA). These materials, often referred to as biopolymers, are biocompatible and biodegradable. In applications related to wound healing, bioinks can be integrated with antibiotic agents or antimicrobial peptides, as well as growth factors such as epidermal growth factor (EGF), fibroblast growth factor (FGF) or vascular endothelial growth factor (VEGF), to enhance cell stimulation, growth, proliferation and migration throughout the healing process . The latest developments in wound care and skin regeneration involve embedding various cell types, including fibroblasts, hUVECs, keratinocytes and human umbilical cord mesenchymal stem cells (hUCMSCs), directly into bioinks. Liu et al. fabricated vascularized full-thickness skin substitute by printing an alginate-gelatin hydrogel to simulate the epidermis, and a phosphosilicate calcium bioglass (PSC)-alginate-GelMA hydrogel containing hUVECs and hUCMSCs to replicate the dermis ( A). For the bioprinting process, the printing speed was 10 mm/s with a filament orientation of 60°. They showed a marked enhancement in blood vessel formation and collagen deposition, demonstrating the effectiveness of these skin substitutes in reconstructing full-thickness skin injuries in rat models . Also, with a view to developing a full-thickness skin substitute, Admane et al. employed extrusion-based 3D bioprinting to create a silk fibroin-gelatin construct containing fibroblasts to mimic the dermis, and a silk fibroin-gelatin layer with keratinocytes to replicate the epidermis ( B). The 3D bioprinted full-thickness skin model showed extensive keratinocyte migration and differentiation, mimicking reepithelialization. Analysis revealed similarities to native human skin, involving pathways related to skin development, extracellular matrix organization and keratinization . Also, Jin et al. developed an advanced 3D bioprinted structure designed to mimic natural full-thickness skin, incorporating the epidermis, dermis and a vascular network. For the 3D bioprinting of the full-thickness functional skin model, the printing speed was set at 5 mm/s, with the air compressor pressure maintained at 0.2 MPa. This model utilised GelMA with HaCaTs for the epidermal layer, an acellular dermal matrix (ADM) with fibroblasts for the dermis and a GelMA mesh with hUVECs for the vascular network ( C). They demonstrated that this functional skin model not only enhanced cell viability and proliferation but also supported epidermal reconstruction in vitro. In vivo, the functional skin model maintained cell viability for at least one week and promoted wound healing, re-epithelization, dermal ECM secretion and angiogenesis, thereby improving wound healing quality . Song et al. developed a bilayer skin scaffold incorporating drug delivery for the repair of full-thickness skin defects. The scaffold features an outer layer of amoxicillin (AMX)-loaded PCL nanofibers fabricated through electrospinning which functions as an antibacterial membrane mimicking the epidermis ( D). The inner layer, designed to replicate the dermis, is a hydrogel composed of sodium alginate and gelatin, infused with recombinant human epidermal growth factor (rhEGF) to maintain wound moisture and promote healing. The successful incorporation of AMX and rhEGF into the scaffold was demonstrated, with the scaffold exhibiting excellent physicochemical properties, effective drug release, and antibacterial activity. Both in vitro and in vivo evaluations revealed enhanced cell adhesion, proliferation and accelerated skin wound healing, alongside favourable biocompatibility. These results suggest that the scaffold holds considerable potential for skin regeneration applications . Numerous studies have utilised 3D printing and bioprinting techniques, both with and without the integration of cellular therapies, to advance skin tissue engineering. A comprehensive overview of these studies is provided in . 5.5. Neural Tissue Engineering The nervous system represents one of the most intricate and complex biological systems formed during development . The human nervous system is divided into two primary components: the central nervous system, consisting of the brain and spinal cord, and the peripheral nervous system, which includes cranial and spinal nerves along with associated ganglia . Traumatic injuries, including traumatic brain injury and spinal cord injury, as well as neurodegenerative diseases such as Alzheimer’s, Parkinson’s, Huntington’s and multiple sclerosis, pose major public health challenges, with limited treatment options available that mainly offer symptomatic relief. Autologous nerve graft transplantation is widely regarded as the optimal approach for addressing severe nerve injuries. Nonetheless, its clinical utility is constrained by several substantial issues, such as the restricted availability of suitable donor nerves and the risk of incompatibilities between the donor and recipient nerves [ , , ]. Despite ongoing clinical advancements, fully effective therapies for neural regeneration are still in early stages, driving interest in neural tissue engineering . Neural tissue engineering focuses on developing biological substitutes that integrate biomimetic 3D scaffolds with cells to improve neural tissue functionality. Advances in 3D printing technology have significantly impacted neural autograft engineering, enabling precise fabrication of tissue-engineered neural implants. Various 3D printing methods, such as extrusion-based printing, laser-assisted bioprinting, SLA and 4D printing, are increasingly employed to create accurately structured implants for nerve injury repair and to develop models and devices for in vitro neural tissue engineering. The materials chosen for the printing of neural structures encompass biocompatible polymers (e.g., PCL and PU), composites (e.g., reduced graphene oxide (rGO)), and hydrogels (e.g., GelMa, HA, collagen and fibrin). These materials must meet specific requirements for printability and biocompatibility, as well as possess suitable physicochemical properties and mechanical strength . Researchers have combined 3D printing/bioprinting with cellular therapies for neural tissue engineering. Lee et al. created photocrosslinkable methacrylated silk fibroin-pectin bioinks ( A). The structures were printed with an extrusion pressure ranging from 50 to 100 mm/min and a printing speed between 1 and 30 mm/s, and they exhibited tunable mechanical properties, favourable biocompatibility and an environment highly supportive of neural induction in 3D bioprinted constructs containing neural stem/progenitor cell spheroids . For spinal cord injury treatment, Song et al. engineered scaffolds composed of PCL microfiber-reinforced spinal cord ECM hydrogels incorporating oxymatrine (OMT) using electrospinning techniques ( B). These scaffolds promoted neuronal differentiation of neural stem cells (NSCs) and suppressed astrocyte proliferation in vitro. In vivo, they promoted the recruitment of NSCs, stimulated neuronal growth, diminished glial scar formation and enhanced motor function recovery in rats with spinal cord injuries . In turn, Lin et al. developed a model designed to forecast cell growth and distribution, aimed at reducing the need for empirical adjustments. They established a multiphysics model that integrates oxygen diffusion and substrate consumption dynamics within a rat adrenal medullary pheochromocytoma (pc-12) cell-laden nerve scaffold ( C). This model was used to simulate and forecast oxygen levels and cellular growth patterns. The scaffold was produced using SLA, and the distribution of cells was assessed through fluorescence staining to confirm the model. The findings demonstrated that the model effectively forecast cellular growth patterns . Numerous studies have utilised 3D printing and bioprinting technologies, both with and without the integration of cellular therapies, to develop the area of neural tissue engineering. provides a comprehensive overview of these investigations. 

Cardiovascular diseases, which include pathologies affecting the myocardium, heart valves and body’s vasculature, are highly prevalent worldwide. These diseases are a leading cause of morbidity and mortality, especially in developed countries [ , , , ]. Current therapeutic approaches include cellular therapies, bypass grafting, the implantation of medical devices, cardiac tissue patches, and organ transplantation [ , , , , ]. Organ transplantation is often not the optimal solution due to the imbalance between the availability of donor organs and the high demand. Additionally, the success of organ transplants is frequently compromised by complications related to immune rejection . Among the various therapeutic approaches, cell therapy has demonstrated success in regenerating cardiovascular tissue. Nevertheless, the absence of ECM limits cell survival following injection, resulting in reduced long-term viability . To address these concerns, 3D printing and bioprinting have emerged as effective approaches for developing scaffolds that incorporate ECM components and enhance cell viability . These structures can more accurately mimic the spatial and mechanical properties of native tissues, which is relevant for their functionality and integration in vivo . Currently, scaffolds for cardiovascular tissue engineering have been fabricated using various 3D printing technologies, such as inkjet printing, SLA and extrusion-based techniques. These scaffolds commonly incorporate biocompatible materials of natural origin, including fibrin, alginate, gelatin, collagen, HA and fibrinogen . However, to achieve complex structures with optimal physical, chemical and mechanical properties, synthetic materials, including poly (glycerol sebacate) (PGS), PCL and poly (ethylene glycol) methacrylate (PEGMA), are also employed. Additionally, some studies utilise a combination of natural and synthetic materials or semi-synthetic derivatives, such as gelatin methacrylate (GelMA) and hyaluronic acid-gelatin methacrylate (HAGM), to optimize scaffold properties while maintaining cell viability . Furthermore, decellularized matrices are extensively utilised due to their provision of a porous, interconnected polymeric network that facilitates cell migration, proliferation and the delivery of essential nutrients for cell survival . Maiullari et al. developed a technique for fabricating 3D cardiac tissue models that include a vascular network ( A). They created multi-cellular constructs using human umbilical vein endothelial cells (hUVECs) and induced pluripotent stem cell-derived cardiomyocytes (iPSC-CMs), encapsulated in alginate and PEG-Fibrinogen (PF). The constructs were fabricated using a custom-designed microfluidic printhead operating at an extrusion speed of 180 mm/min. Their research demonstrated that bioprinted endothelial cells are capable of forming functional vascular structures within the transplanted tissues and interacting with the host’s existing vascular network . To address damaged myocardium, Beijleri et al., produced a 3D-printed cardiac patch consisting of a decellularized cardiac extracellular matrix (cECM) hydrogel combined with GelMA. This patch was designed for the delivery of paediatric human cardiac progenitor cells (hCPC) ( B). The GelMA-cECM bioinks ensure uniform distribution of cECM and hCPCs, with the hCPCs maintaining over 75% viability. Additionally, conditioned media from GelMA-cECM patches demonstrate more than a 2-fold increase in angiogenic potential. The patches also remain adhered to rat hearts and show vascularization over a 14-day period in vivo . Also, in the development of tissue patches, Melhem et al. proposed a hydrogel patch embedded with multiple microchannels to enhance cell retention and factor delivery at the target tissue ( C). They integrated bone marrow-derived MSCs (BMSCs) into a hydrogel by cross-linking a poly (ethylene glycol) dimethacrylate (PEGDMA) solution containing the cells. Microchannels with precise diameters were created within the cell-loaded hydrogel using an SLA unit for in situ cross-linking. This 3D-printed, microchanneled hydrogel was designed as an advanced therapeutic tool for sustained delivery of multiple therapeutics, aiming to improve outcomes in ischemic heart injury . Noor et al. established a technique to 3D print thick, vascularized cardiac patches tailored to a patient’s specific properties. They used patient-derived cells, reprogrammed into cardiomyocytes and endothelial cells, and a personalized hydrogel made from the patient’s extracellular matrix. These components were combined to create bioinks for printing cardiac tissue and blood vessels. The patches, optimized for oxygen transfer, demonstrated proper structure and function in vitro, and the approach was further validated by successfully printing cellularised human hearts presenting a natural architecture ( D) . In turn, to mimic human microvasculature, Cui et al. developed a bioink combining human microvascular endothelial cells (hMVECs) and fibrin, used to fabricate micron-sized fibrin channels via drop-on-demand polymerization ( E). This aqueous-based printing method reduces cellular damage. The hMVECs printed with fibrin aligned within the channels and proliferated to form continuous linings, resulting in a three-dimensional tubular structure. The study concluded that the concurrent printing of cells and scaffolds enhances the proliferation of hMVECs and supports the formation of microvascular networks . Several researchers have employed 3D printing or bioprinting, with or without cell therapies, to advance cardiovascular tissue engineering. provides a summary of these studies. 

Severe bone defects resulting from aging, trauma, osteoporosis, degenerative diseases, autoimmune diseases (such as rheumatoid arthritis) or tumour removal are a leading cause of disability globally, affecting an estimated 1.71 billion people . Current therapeutic strategies for addressing these defects include autografting, allografting, xenografting and bone transplantation. Despite their application, these methods present significant risks, including the potential for infectious disease transmission and immune rejection . To address the limitations of traditional bone defect treatments, 3D printing and bioprinting techniques, with or without cellular therapies, have been developed for bone tissue engineering. These methods enable the large-scale fabrication of custom-tailored bone tissues, meeting the growing demand for functionalized bone implants . The 3D-printed scaffolds serve as a support for cell growth and differentiation, forming a hierarchical bone microvascular architecture . For optimal performance, a scaffold should possess several critical attributes: biocompatibility, sterility, osteoconductivity, biodegradability and a porous, interconnected structure that supports cellular infiltration and nutrient transport. Additionally, it should effectively repair bone defects while closely mimicking the characteristics of native bone tissue [ , , ]. The most commonly employed printing techniques in bone tissue engineering are laser powder bed fusion, vat photopolymerization and extrusion-based methods. Various biomaterials are utilised for printing scaffolds: ceramics (e.g., beta-tricalcium phosphate (β-TCP), hydroxyapatite (HAp) and amorphous calcium phosphate (ACP)); natural polymers (e.g., matrigel, alginate, HA and dextran emulsion); synthetic polymers (e.g., GelMA, PLGA, PCL and polyethylene glycol diacrylate (PEGDA); metals (e.g., titanium alloy (Ti6Al4V), tantalum (Ta) and titanium (Ti)), as well as the combinations of these materials. Lei et al. developed Ti6Al4V-based porous tantalum (Ta) scaffolds with high interfacial strength using laser powder bed fusion. In this process, porous Ta was directly deposited onto a solid Ti6Al4V substrate ( A). The manufacturing process was carried out in an argon atmosphere, using a 300 W laser at 250 mm/s for Ta and a 250 W laser at 300 mm/s for Ti6Al4V. In vitro biocompatibility assessments conducted with rat bone marrow mesenchymal stem cells (r-BMSCs) confirmed the scaffolds’ biocompatibility. The findings demonstrated the strong mechanical compatibility and osteointegration properties of the Ti6Al4V-based porous Ta scaffold, underscoring its considerable potential for orthopaedic applications . For mandibular bone defect reconstruction, Yu et al. encapsulated BMSCs in matrigel and infiltrated this mixture into porous Ti6Al4V scaffolds. The study demonstrated that rats with critical full-thickness mandibular defects treated with Matrigel-infused Ti6Al4V scaffolds exhibited significantly greater new bone formation compared to those treated with either local BMSC injections or Matrigel alone ( B). These results indicate that Matrigel enhances the 3D microenvironment for BMSCs, positioning Matrigel-infused scaffolds as a promising method for improving bone regeneration in 3D-printed Ti6Al4V scaffolds . Wu et al. produced a 3D-bioprintable scaffold combining alginate and β-TCP for the treatment of bone defects ( C). The scaffolds were fabricated at a printing speed of 2 mm/s, with an extrusion pressure ranging from 60 to 200 kPa. MG-63 cells were seeded onto these scaffolds. The 3D-printed scaffolds using a 10% alginate/β-TCP bioink exhibited enhanced physical characteristics and significantly improved cell viability and alkaline phosphatase activity. These findings suggest that the scaffolds have considerable potential for application in personalized bone regeneration therapies . To produce a scaffold that replicates bone microstructure, Ressler et al. developed trabecular-like porous scaffolds using ceramic vat photopolymerization with HAp powders doped with magnesium (Mg 2+ ), strontium (Sr 2+ ) and zinc (Zn 2+ ). Scaffolds sintered at 1100–1300 °C exhibited mechanical properties similar to trabecular bone, with optimal performance at 1300 °C. The microstructure resembled cancellous bone and the incorporation of trace elements resulted in a biphasic calcium phosphate system (HAp/β-TCP), potentially enhancing bioactivity . Numerous studies have utilised 3D printing or bioprinting, either with or without cellular therapies, to advance bone tissue engineering. An overview of these studies is provided in . 

The liver plays a vital role in blood protein synthesis, glucose metabolism and the detoxification of metabolites . It is also the only organ in the human body capable of efficient regeneration. However, this regenerative capacity can be compromised by excessive drug use or viral infections, which can cause irreversible damage to hepatocytes and lead to liver failure . Chronic liver diseases, including fibrosis, cirrhosis, chronic viral hepatitis and fatty liver disease, significantly contribute to global morbidity and mortality. Unfortunately, advancements in treatment options for these conditions remain limited . The primary medical intervention for liver failure is partial or total liver transplantation. Nevertheless, this approach faces challenges, including limited donor availability, immune rejection and variable graft success rates. Alternative approaches in tissue engineering encompass bioartificial liver systems, which involve the in vitro creation of liver tissue to repair or replace damaged liver segments, as well as hepatocyte transplantation and cellular therapy methods . There has been a continuous search for a reliable and reproducible source of hepatocytes, whether for liver regeneration therapy, seeding liver support devices or in vitro screening applications . Numerous studies have investigated the 3D bioprinting of liver tissue utilizing either stem cells or immortalized hepatic cell lines . Stem cells are particularly promising due to their ability to express hepatocyte-like phenotypes. In contrast, adult hepatocytes are scarce, challenging to isolate, exhibit poor propagation and experience rapid functional deterioration in vitro . In liver tissue engineering, a variety of biomaterials are employed, categorized into natural and synthetic polymers . Natural polymers, such as alginate, HA, collagen, cellulose nanocrystal (CNC) and gelatin, present the advantages of enhanced cell compatibility and ease of manipulation. Nevertheless, they are constrained by relatively weaker mechanical properties, limited availability and variable degradation rates. Consequently, synthetic polymers, such as PCL, have been developed to provide superior mechanical strength, flexibility, processability and adjustable degradability. Despite these advantages, synthetic polymers generally lack cell recognition and adhesion sites, resulting in reduced biocompatibility compared to natural polymers . To mitigate these issues, a strategy of combining natural and synthetic polymers is employed to create effective bioinks for 3D bioprinting. Additionally, the liver decellularized extracellular matrix (dECM), sourced from animals, is frequently used to establish microenvironments for liver cells, offering cross-species tolerance and minimising the risk of immune rejection [ , , ]. Yang et al. developed a liver tissue model through 3D bioprinting using HepaRG cells, a widely utilised hepatic progenitor cell line. Each 3D bioprinted hepatorganoid was produced through forced extrusion at a rate of 150 mm 3 /min, using a layer-by-layer approach. This 3D bioprinted model, referred to as hepatorganoids, exhibited essential liver functions, including albumin production, drug metabolism and glycogen accumulation, after a 7-day differentiation period ( A). In vivo studies showed that the 3D bioprinted hepatorganoids further matured, exhibiting enhanced synthesis of liver-specific proteins and more human-like drug metabolism. Notably, transplantation of 3D bioprinted hepatorganoids significantly increased survival rates in recipient mice. These findings suggest that 3D bioprinted hepatorganoids can effectively undergo hepatic differentiation and ameliorate liver failure in vivo . In the study conducted by Xie et al., a 3D bioprinted model of hepatocellular carcinoma, derived from patient cells, was created using isolated primary hepatocellular carcinoma cells mixed with gelatin and sodium alginate to form a bioink ( B). The resulting models were successfully generated and exhibited substantial growth over prolonged culture durations. They preserved key characteristics of the original hepatocellular carcinoma tumours, including consistent biomarker expression and stable genetic alterations and expression profiles. Thus, 3D bioprinted hepatocellular carcinoma models prove to be reliable in in vitro systems, suitable for long-term culture and capable of predicting patient-specific drug responses for tailored therapeutic approaches . Lewis et al. studied a technique for 3D printing gelatin into precisely defined geometries, which exhibit distinct biological effects on seeded hepatocytes ( C). Their research reveals that the structural configuration of gelatin can markedly impact biological processes. An undifferentiated hepatocyte cell line demonstrated high viability and proliferation on 3D-printed scaffolds with two distinct geometries. Notably, hepatocyte-specific functions—such as albumin secretion, cytochrome P450 activity and bile transport—were enhanced in more interconnected 3D-printed gelatin structures compared to those with less interconnectivity and traditional two-dimensional (2D) cultures. This study underscores the gap between gene expression and protein functionality in simplistic 2D cultures, highlighting the importance of a physiologically relevant 3D environment for optimizing hepatocyte expression and functionality . Jeon et al. employed 3D bioprinting technology to reconstruct liver tissues and organs using human hepatocellular carcinoma (HepG2) cells, a liver cancer-derived cell line. They created multi-layered 3D structures by integrating alginate with HepG2 cells ( D). The study demonstrated that replicating the 3D hepatic architecture using this technology enhances the stability and gene expression profiles of HepG2 cells. Cells cultured on these 3D alginate scaffolds for three weeks were analysed via fluorescence microscopy, histology and immunohistochemistry. Results indicated that HepG2 cells exhibited improved growth and liver-specific gene expression in 3D cultures compared to 2D cultures, highlighting the effectiveness of 3D bioprinting in mimicking liver architecture and enhancing cellular function . Using the same 3D bioprinting technique, Wu et al. produced a novel bioink containing alginate, CNC and GelMA (namely, 135ACG hybrid ink), aimed at fabricating both cell-laden and acellular structures ( E). The structures were printed using a 32-gauge nozzle at a printing speed of 20 mm/s and an air pressure of 20 psi. The bioink presented good shear-thinning behaviour and solid-like characteristics, ensuring high printability and minimal cell damage. Following crosslinking, it formed a rigid ECM conducive to stromal cell growth. The team engineered a GelMA bioink with suitable mechanical properties to mimic human liver tissue, enabling the printing of liver lobule-mimetic constructs with precise cell placement (fibroblasts and hepG2) in different ECMs (135ACG and GelMA). These constructs were used to study the impact of mechanical stimuli and cellular interactions on cell behaviour. These findings demonstrated that fibroblasts proliferated effectively within the rigid 135ACG matrix, whereas HepG2 cells developed into spheroid structures in the more compliant GelMA matrix. Co-cultures of hepG2 and fibroblasts cells showed increased albumin production, highlighting the role of soluble factors in enhancing hepatic function. The study demonstrated that the developed bioinks and printing methods are effective for creating complex, multi-cellular constructs with varied ECMs, advancing both fundamental research and tissue engineering applications . Several studies have employed 3D printing/bioprinting with or without the incorporation of cellular therapies to advance liver tissue engineering. A summary of these studies is presented in . 

The skin, regarded as the largest organ in the human body, plays essential roles in various functions, including serving as a protective barrier, regulating body temperature and preventing dehydration . Extensive full-thickness skin wounds, which damage underlying blood vessels, pose serious risks due to induced cellular hypoxia and nutrient deprivation . While autografts are considered the “gold standard” for treating severe skin injuries, their use is limited by issues such as donor site availability and associated morbidity . Furthermore, existing commercial skin substitutes lack the sufficient vascular networks necessary for effective nutrient delivery in full-thickness wounds. Consequently, skin substitutes are seen as promising alternatives, offering the potential for vascularized skin reconstruction with customized cell compositions and controlled geometrical structures . 3D printing/bioprinting has arisen as an innovative technological approach in skin tissue engineering, creating structures by depositing cell-embedded bioinks layer by layer . This approach includes several techniques, including DLP, extrusion bioprinting, inkjet printing and electrospinning. The materials commonly used in this field are natural polymers (e.g., collagen, alginate, HA, gelatin and fibroin) and synthetic polymers (e.g., PCL, PLGA, polyglycolic acid, polyurethanes, polycarbonates and PEGDA). These materials, often referred to as biopolymers, are biocompatible and biodegradable. In applications related to wound healing, bioinks can be integrated with antibiotic agents or antimicrobial peptides, as well as growth factors such as epidermal growth factor (EGF), fibroblast growth factor (FGF) or vascular endothelial growth factor (VEGF), to enhance cell stimulation, growth, proliferation and migration throughout the healing process . The latest developments in wound care and skin regeneration involve embedding various cell types, including fibroblasts, hUVECs, keratinocytes and human umbilical cord mesenchymal stem cells (hUCMSCs), directly into bioinks. Liu et al. fabricated vascularized full-thickness skin substitute by printing an alginate-gelatin hydrogel to simulate the epidermis, and a phosphosilicate calcium bioglass (PSC)-alginate-GelMA hydrogel containing hUVECs and hUCMSCs to replicate the dermis ( A). For the bioprinting process, the printing speed was 10 mm/s with a filament orientation of 60°. They showed a marked enhancement in blood vessel formation and collagen deposition, demonstrating the effectiveness of these skin substitutes in reconstructing full-thickness skin injuries in rat models . Also, with a view to developing a full-thickness skin substitute, Admane et al. employed extrusion-based 3D bioprinting to create a silk fibroin-gelatin construct containing fibroblasts to mimic the dermis, and a silk fibroin-gelatin layer with keratinocytes to replicate the epidermis ( B). The 3D bioprinted full-thickness skin model showed extensive keratinocyte migration and differentiation, mimicking reepithelialization. Analysis revealed similarities to native human skin, involving pathways related to skin development, extracellular matrix organization and keratinization . Also, Jin et al. developed an advanced 3D bioprinted structure designed to mimic natural full-thickness skin, incorporating the epidermis, dermis and a vascular network. For the 3D bioprinting of the full-thickness functional skin model, the printing speed was set at 5 mm/s, with the air compressor pressure maintained at 0.2 MPa. This model utilised GelMA with HaCaTs for the epidermal layer, an acellular dermal matrix (ADM) with fibroblasts for the dermis and a GelMA mesh with hUVECs for the vascular network ( C). They demonstrated that this functional skin model not only enhanced cell viability and proliferation but also supported epidermal reconstruction in vitro. In vivo, the functional skin model maintained cell viability for at least one week and promoted wound healing, re-epithelization, dermal ECM secretion and angiogenesis, thereby improving wound healing quality . Song et al. developed a bilayer skin scaffold incorporating drug delivery for the repair of full-thickness skin defects. The scaffold features an outer layer of amoxicillin (AMX)-loaded PCL nanofibers fabricated through electrospinning which functions as an antibacterial membrane mimicking the epidermis ( D). The inner layer, designed to replicate the dermis, is a hydrogel composed of sodium alginate and gelatin, infused with recombinant human epidermal growth factor (rhEGF) to maintain wound moisture and promote healing. The successful incorporation of AMX and rhEGF into the scaffold was demonstrated, with the scaffold exhibiting excellent physicochemical properties, effective drug release, and antibacterial activity. Both in vitro and in vivo evaluations revealed enhanced cell adhesion, proliferation and accelerated skin wound healing, alongside favourable biocompatibility. These results suggest that the scaffold holds considerable potential for skin regeneration applications . Numerous studies have utilised 3D printing and bioprinting techniques, both with and without the integration of cellular therapies, to advance skin tissue engineering. A comprehensive overview of these studies is provided in . 

The nervous system represents one of the most intricate and complex biological systems formed during development . The human nervous system is divided into two primary components: the central nervous system, consisting of the brain and spinal cord, and the peripheral nervous system, which includes cranial and spinal nerves along with associated ganglia . Traumatic injuries, including traumatic brain injury and spinal cord injury, as well as neurodegenerative diseases such as Alzheimer’s, Parkinson’s, Huntington’s and multiple sclerosis, pose major public health challenges, with limited treatment options available that mainly offer symptomatic relief. Autologous nerve graft transplantation is widely regarded as the optimal approach for addressing severe nerve injuries. Nonetheless, its clinical utility is constrained by several substantial issues, such as the restricted availability of suitable donor nerves and the risk of incompatibilities between the donor and recipient nerves [ , , ]. Despite ongoing clinical advancements, fully effective therapies for neural regeneration are still in early stages, driving interest in neural tissue engineering . Neural tissue engineering focuses on developing biological substitutes that integrate biomimetic 3D scaffolds with cells to improve neural tissue functionality. Advances in 3D printing technology have significantly impacted neural autograft engineering, enabling precise fabrication of tissue-engineered neural implants. Various 3D printing methods, such as extrusion-based printing, laser-assisted bioprinting, SLA and 4D printing, are increasingly employed to create accurately structured implants for nerve injury repair and to develop models and devices for in vitro neural tissue engineering. The materials chosen for the printing of neural structures encompass biocompatible polymers (e.g., PCL and PU), composites (e.g., reduced graphene oxide (rGO)), and hydrogels (e.g., GelMa, HA, collagen and fibrin). These materials must meet specific requirements for printability and biocompatibility, as well as possess suitable physicochemical properties and mechanical strength . Researchers have combined 3D printing/bioprinting with cellular therapies for neural tissue engineering. Lee et al. created photocrosslinkable methacrylated silk fibroin-pectin bioinks ( A). The structures were printed with an extrusion pressure ranging from 50 to 100 mm/min and a printing speed between 1 and 30 mm/s, and they exhibited tunable mechanical properties, favourable biocompatibility and an environment highly supportive of neural induction in 3D bioprinted constructs containing neural stem/progenitor cell spheroids . For spinal cord injury treatment, Song et al. engineered scaffolds composed of PCL microfiber-reinforced spinal cord ECM hydrogels incorporating oxymatrine (OMT) using electrospinning techniques ( B). These scaffolds promoted neuronal differentiation of neural stem cells (NSCs) and suppressed astrocyte proliferation in vitro. In vivo, they promoted the recruitment of NSCs, stimulated neuronal growth, diminished glial scar formation and enhanced motor function recovery in rats with spinal cord injuries . In turn, Lin et al. developed a model designed to forecast cell growth and distribution, aimed at reducing the need for empirical adjustments. They established a multiphysics model that integrates oxygen diffusion and substrate consumption dynamics within a rat adrenal medullary pheochromocytoma (pc-12) cell-laden nerve scaffold ( C). This model was used to simulate and forecast oxygen levels and cellular growth patterns. The scaffold was produced using SLA, and the distribution of cells was assessed through fluorescence staining to confirm the model. The findings demonstrated that the model effectively forecast cellular growth patterns . Numerous studies have utilised 3D printing and bioprinting technologies, both with and without the integration of cellular therapies, to develop the area of neural tissue engineering. provides a comprehensive overview of these investigations. 

Despite the benefits of 3D printing/bioprinting and cellular therapies for regenerative medicine, there are several challenges that need to be addressed. Despite extensive research efforts in recent years, the clinical application of these technologies has been constrained. The lack of sufficient animal studies and the absence of viable 3D models in clinical trials underscore the need for further focus and development in these critical areas [ , , ]. One of the current challenges is the advancement of functional vascular networks within bioprinted tissues and organs. Specifically, creating a vascular system that can seamlessly connect with the host’s native blood vessels is complicated by the complex structural architecture and the variability of tissue components [ , , ]. Vascularization is crucial for providing nutrients and oxygen, which are vital for maintaining cell viability and ensuring tissue functionality. A robust, multi-level vascular network is necessary to support the long-term survival and growth of bioprinted organs, incorporating smooth muscle cells and hUVECs into the blood vessels . To overcome challenges related to nutrient and oxygen transport, researchers have utilised proangiogenic factors, such as vascular endothelial growth factor (VEGF) and basic fibroblast growth factor (BFGF), to promote the development of microvessels. Additionally, incorporating endothelial cells into the culture medium has been shown to promote microvessel formation and the development of angiogenic sprouts in engineered constructs . However, endothelial cells and angiogenic factors generally do not produce perfusable constructs in a fast way . Bioreactors offer a solution by continuously supplying media to porous constructs, reducing the reliance on arterial scaffolds and large tissue samples. Despite this, these constructs often lack microvessels and are stored outside of bioreactors, which can compromise cell survival. Microfluidic systems represent a potential alternative for vascular network fabrication, though scaling these systems to larger physiological sizes remains a significant challenge . The selection and sourcing of cells is another challenge in regenerative medicine. Cells used in bioinks must possess some key characteristics: high proliferative capacity, printability, functionality, safety and economic viability . The choice of cells encapsulated in bioinks critically influences their differentiation potential and ability to develop into various lineages. Although live cells, such as stem cells, are very promising, their practical application is limited by issues related to availability and ethical concerns. Researchers have reported the successful integration of stem cells with bioprinting technologies . An additional challenge is the mass production of cells, which needs substantial quantities of cells and increases the demand on in vitro expansion cultures. Addressing the cost-effectiveness of large-scale cell production is a crucial challenge . Extended processing times and mechanical forces experienced during 3D printing can adversely affect cell viability by altering cell geometry and disrupting signalling pathways . To address these issues, it is essential to enhance existing bioprinting techniques to minimise processing duration and to develop specialized buffers that can protect cells throughout the printing process. In 3D bioprinting, choosing the most appropriate biomaterials is critical for the effective fabrication of tissues with clinical relevance . Biomaterials are essential for providing structural support, maintaining cellular viability and ensuring long-term tissue integration. While numerous polymers traditionally utilised in 3D printing and tissue engineering have been explored for bioprinting due to their availability and previous applications, they may not always provide the optimal biological compatibility required for successful bioprinting outcomes . These materials might display excessive biological reactivity, which can result in undesirable cellular interactions and premature or inappropriate differentiation of stem cells. For a bioink to be suitable for clinical applications, it must have specific characteristics, such as structural stability, the ability to support cell proliferation, and a degradation rate that matches the needs of tissue regeneration. Additionally, bioinks must be compatible with bioprinting technologies to facilitate rapid prototyping . A major challenge is ensuring that printed structures are biocompatible and provide an appropriate environment for cell growth. Current research is focused on developing novel biopolymers and hydrogels that more accurately replicate the nanoscale features and responsive characteristics of the ECM and native tissue microenvironment . However, these advanced materials often encounter compatibility issues with traditional bioprinting techniques. Many of these materials may lack the necessary structural integrity, leading to collapse if they are too soft . One potential solution is to combine various materials to harness their individual strengths, such as merging the mechanical properties of more rigid materials with the cell-supportive and biocompatible features of softer ones . Furthermore, the development and implementation of 3D printing and cell therapies in regenerative medicine pose significant financial challenges, which could prevent the widespread adoption of these technologies. The costs associated with research, development and clinical trials are substantial, and securing funding for these endeavours can be difficult. Moreover, the cost-efficiency of these technologies needs careful consideration, particularly in relation to the high expenses of 3D printers, cellular materials and associated computer software . Generally, the costs of maintaining and scaling bioprinting technologies limit the rapid integration of 3D printing capabilities into clinical settings . Another challenge is the size of bioprinted tissues. Currently, bioprinted constructs are typically small and consist of a limited number of cell types, which restricts their functionality and scalability [ , , , ]. Moreover, 3D printers are often constrained by their build volume, which limits the maximum size of bioprinted tissues and complicates the creation of entire 3D-printed organs . Ethical considerations are increasingly influencing the development of 3D bioprinting technologies. The transition of 3D printing/bioprinting research from the laboratory to clinical applications raises both established ethical concerns, as discussed in the bioethics literature, and emerging issues specific to the unique features of these new treatments [ , , ]. Several ethical concerns arise with the advancement of 3D printing/bioprinting technologies, including issues related to the sourcing of cells and tissues, as previously mentioned. Furthermore, the development of these technologies requires careful ethical consideration, particularly in research involving animal welfare and the design of human clinical trials. It is crucial to find a balance between the necessity of animal testing and the ethical treatment of research subjects, while thoroughly assessing the risks and benefits for human trials . As bioprinting technologies evolve, ensuring equitable access becomes an ethical obligation. Addressing potential healthcare disparities and balancing the advantages of personalized medicine with fair healthcare delivery are significant challenges that must be addressed . The rapid rate of the development of bioprinting has outpaced existing regulatory frameworks, resulting in gaps that need to be closed. Despite their complexity and time demands, valuable guidelines for these processes can be obtained from regulatory agencies such as the Food and Drug Administration (FDA) in the United States and the European Medicines Agency (EMA) in the European Union, especially regarding 3D-printed medical devices. For tissue bioprinting to achieve clinical translation, it is essential to establish a clear and defined regulatory pathway . Additionally, challenges and concerns related to biosafety and liability arise when fabricating internal tissues and organs. The clinical translation of bioprinting techniques will depend on regulatory bodies’ thorough the evaluation of safety, efficacy and risk. Globally, regulatory authorities face challenges in addressing the potential and uncertain risks associated with 3D bioprinting, such as immune responses to bioinks or materials . In the absence of specific regulations, the FDA is currently relying on the guidelines of the Center for Biologics Evaluation and Research (CBER) for 3D bioprinting products. These products require FDA approval, and adherence to regulatory guidelines is mandatory from the initial stages of product development. As the field advances, more 3D bioprinting products are likely to emerge, highlighting the need for more specific regulatory guidelines. Currently, only South Korea’s Ministry of Food and Drug Safety (MFDS) and Japan’s Pharmaceuticals and Medical Devices Agency (PMDA) have developed specific guidelines for 3D bioprinting. Thus, the global development of comprehensive regulations for 3D bioprinting techniques, bioinks and printers is of increasing importance . These limitations and challenges highlight the complexity of integrating 3D printing with cellular therapies in regenerative medicine, emphasizing the need for continuous research and development in this field. Overcoming these obstacles requires transdisciplinary collaboration among cell biologists, engineers, physiologists and pharmaceutical industry partners to advance and expand the potential of this technology . Despite the existing challenges, advancements persist, and it is anticipated that these technologies will increasingly impact the treatment of various diseases and conditions in the future. Ongoing research is progressively making the goal of creating safe and fully functional bioprinted tissues more achievable . 

The current state of clinical trials combining 3D printing/bioprinting and cellular therapies is promising, though they are still in their early stages. Numerous research groups are actively developing and testing new treatments utilizing these technologies, and several clinical trials have been initiated in recent years to assess their safety and efficacy. Tissue engineering currently has broad applications, including the development of various tissues, cardiac, vascular, bone, skin, neural, cartilage and retinal tissues, among others. Conventionally, this method entails seeding cells onto a porous scaffold to promote their growth and, subsequently, promote tissue development . The approach presents several advantages, including providing strong structural support with suitable degradation timing, regulating the cellular environment and allowing for effective nutrient and waste exchange between the scaffold and cells . Bioprinting has opened up a whole new era in tissue regeneration, allowing the production of patient-specific autologous organs and tissues [ , , , , ]. This rapid prototyping technique allows for the fabrication of intricate tissue and organ structures by precisely depositing living cells and biomaterials layer by layer, based on a CAD model. Using this approach, 3D constructs can be produced with high accuracy in terms of positioning and architecture, including shape, pore geometry and interconnectivity. This allows for the development of tissue and organ models that closely resemble the human body, with high reproducibility [ , , ]. By simultaneously printing multiple cell types and biomaterials, bioprinting can replicate the structural and biochemical complexity of living tissues, creating a heterogeneous microenvironment at specified locations [ , , ]. Although in vitro models have advanced significantly for developing new therapies, their application in surgical settings is still not fully realized. However, significant progress is being made in hydrogel design and the development of advanced technological tools. These advancements are bringing us closer to meeting the fidelity and safety standards required for bioprinted constructs to be used consistently and in a personalized way in patients in the near future . Additionally, artificial intelligence (AI) is leveraging significant advances in AM for regenerative medicine, particularly in the production of personalized medical devices, implants and tissue engineering. AI algorithms optimize 3D printing parameters such as speed, temperature and layer thickness, increasing the precision and quality of printed structures .This optimization is crucial for the creation of patient-specific implants and prostheses, as well as for the development of 3D-printed tissues and organs, advancing the field of regenerative medicine . In addition, AI-assisted 3D printing enables the manufacture of detailed models for surgical planning and testing, improving surgical results and precision . Moving forward, researchers predict that AI will continue to revolutionize AM, enabling a more seamless use of the technology for complex designs in artificial organs, flexible electronics and wearable biosensors . The synergy between AI and AM promises to unlock new possibilities for design, production and innovation in regenerative medicine . The future of tissue biomaterials is expected to replicate not only the structural design and characteristics of organs and tissues but also their dynamic, functional behaviours . The concept of time as the fourth dimension (4D) has gained prominence, in the context of bioprinting, introducing two key aspects: materials capable of deformation and structures that mature after printing . This new approach to bioprinting addresses the complexity of the system, which is very important for fully understanding the behaviour of functional living materials during the post-processing stage . 

The combination of 3D printing/bioprinting technologies with cellular therapies represents a significant advancement in regenerative medicine. Advances in this area are leading to the development of new functional tissues that closely resemble native tissues. This review has detailed the fundamental principles and applications of each technology individually and has elucidated how their convergence is pushing the boundaries of tissue engineering This study underlines that the latest developments have markedly enhanced the capacity to produce complex 3D cellular structures presenting high precision. Innovations in bioprinting techniques, coupled with improvements in biomaterials and cellular engineering, have enabled more sophisticated control over tissue architecture and cellular organization. These advancements facilitate the creation of more accurate and functional tissue models, which hold promise for personalized regenerative therapies. Nevertheless, there are still several critical barriers that need to be addressed, such as the development of functional vascular networks, regulatory and ethical issues and the advancement of suitable biomaterials. The potential impact of these integrated technologies in regenerative medicine is expected to be high in the future. As the field progresses, solving these issues will be essential to realizing the full potential of 3D printing/bioprinting and cell therapies. Ongoing research and innovation in these areas is therefore expected to produce transformative advances in personalized medicine and tissue regeneration, ultimately improving outcomes for patients and advancing therapeutic options. 

