Progress in Organ Bioprinting for Regenerative Medicine

Xiang Wang; Di Zhang; Yogendra Pratap Singh; Miji Yeo; Guotao Deng; Jiaqi Lai; Fei Chen; Ibrahim T. Ozbolat; Yin Yu

doi:10.1016/j.eng.2024.04.023

Engineering ›› 2024, Vol. 42 ›› Issue (11) :129 -152. DOI: 10.1016/j.eng.2024.04.023

Research

Progress in Organ Bioprinting for Regenerative Medicine

Xiang Wang ^a^,^#
, Di Zhang ^a^,^#
, Yogendra Pratap Singh ^b^,^c
, Miji Yeo ^b^,^c
, Guotao Deng ^a
, Jiaqi Lai ^a
, Fei Chen ^a^,^*
, Ibrahim T. Ozbolat ^b^,^c^,^d^,^e^,^f^,^g^,^h^,^*
, Yin Yu ^a^,^*

Author information +

^aCenter for Materials Synthetic Biology, Chinese Academy of Sciences Key Laboratory of Quantitative Engineering Biology, Shenzhen Institute of Synthetic Biology, Shenzhen Institute of Advanced Technology, Chinese Academy of Sciences, Shenzhen 518055, China

^bThe Huck Institutes of the Life Sciences, Pennsylvania State University, University Park, PA 16802, USA

^cEngineering Science and Mechanics Department, Pennsylvania State University, University Park, PA 16802, USA

^dDepartment of Biomedical Engineering, Pennsylvania State University, University Park, PA 16802, USA

^eMaterials Research Institute, Pennsylvania State University, University Park, PA 16802, USA

^fDepartment of Neurosurgery, Pennsylvania State College of Medicine, Hershey 17033, PA, USA

^gPennsylvania State Cancer Institute, Pennsylvania State University, University Park, PA 16802, USA

^hBiotechnology Research and Application Center, Cukurova University, Adana 01130, Türkiye

^bThe Huck Institutes of the Life Sciences, Pennsylvania State University, University Park, PA 16802, USA

^cEngineering Science and Mechanics Department, Pennsylvania State University, University Park, PA 16802, USA

^dDepartment of Biomedical Engineering, Pennsylvania State University, University Park, PA 16802, USA

^eMaterials Research Institute, Pennsylvania State University, University Park, PA 16802, USA

^fDepartment of Neurosurgery, Pennsylvania State College of Medicine, Hershey 17033, PA, USA

^gPennsylvania State Cancer Institute, Pennsylvania State University, University Park, PA 16802, USA

^hBiotechnology Research and Application Center, Cukurova University, Adana 01130, Türkiye

E⁃mail addresses: fei.chen1@siat.ac.cn (F. Chen),

ito1@psu.edu (I.T. Ozbolat),

yin.yu@siat.ac.cn(Y. Yu)

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Abstract

Organ damage or failure arising from injury, disease, and aging poses challenges due to the body’s limited regenerative capabilities. Organ transplantation presents the issues of donor shortages and immune rejection risks, necessitating innovative solutions. The three-dimensional (3D) bioprinting of organs on demand offers promise in tissue engineering and regenerative medicine. In this review, we explore the state-of-the-art bioprinting technologies, with a focus on bioink and cell type selections. We follow with discussions on advances in the bioprinting of solid organs, such as the heart, liver, kidney, and pancreas, highlighting the importance of vascularization and cell integration. Finally, we provide insights into key challenges and future directions in the context of the clinical translation of bioprinted organs and their large-scale production.

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Keywords

Organ printing / Three-dimensional bioprinting / Regenerative medicine / Tissue engineering

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Xiang Wang, Di Zhang, Yogendra Pratap Singh, Miji Yeo, Guotao Deng, Jiaqi Lai, Fei Chen, Ibrahim T. Ozbolat, Yin Yu. Progress in Organ Bioprinting for Regenerative Medicine. Engineering, 2024, 42(11): 129-152 DOI:10.1016/j.eng.2024.04.023

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1. Introduction

Tissue or organ damage and failure commonly occur in humans, resulting from injury, disease, and aging [1], while human bodies possess limited regenerative capabilities in most tissues/organs. Organ transplantation stands as a vital medical intervention for individuals facing organ failure or severe tissue damage, but it presents significant challenges, such as the shortage of donors and the risk of immune rejection [2], [3], underscoring the urgent need for innovative approaches to tackle these challenges.

Biofabrication holds the potential to create intricate three-dimensional (3D) biological structures that can mimic the functional organization of human tissues and organs. This process facilitates physiologically relevant cellular interactions, and maintaining this dynamic reciprocity within 3D microenvironments is crucial for restoring appropriate biological behavior [4]. In the realm of tissue engineering and regenerative medicine, biofabrication is characterized as the creation of biofunctional products with structural organization. This involves utilizing living cells, bioactive molecules, and biomaterials in various techniques such as 3D bioprinting or bioassembly [5]. In particular, 3D bioprinting exhibits unparalleled potential in tackling these challenges, revolutionizing the field of biomanufacturing and spearheading numerous significant advancements in tissue engineering and regenerative medicine [6]. In addition to 3D bioprinting, Table 1 [7], [8], [9], [10] lists alternative solutions to tackle the challenge of organ shortage and the problem of rejection during organ transplantation.

In this review, we summarize the state-of-art technologies utilized in organ bioprinting and outline key considerations in choosing appropriate bioinks and the types of cells involved (Fig. 1). We then discuss the latest advances in the bioprinting of solid organs, such as the heart, liver, kidney, and pancreas, emphasizing the importance of vascularization and the integration of different cell types during organ bioprinting. Finally, we outline key challenges that remain to be addressed in translating the preclinical success of organ bioprinting to the clinic and suggest future directions for this exciting field.

2. Fundamentals of 3D bioprinting

Bioprinting harnesses the capabilities of 3D printing to precisely pattern and assemble both living and nonliving biomaterials, culminating in the fabrication of intricate biological structures. The process of organ bioprinting typically involves two key steps: first, the generation of detailed 3D models based on tissue/organ information obtained through various imaging methods; second, the subsequent fabrication of components that faithfully recapitulate the structure and function of target tissue/organs [11], [12], [13]. 3D bioprinting demonstrates remarkable potential for the creation of solid organs that not only mirror realistic anatomical features but also exhibit functional capabilities. This technology holds significant promise for advancing research across diverse domains, including regenerative medicine, tissue engineering, drug discovery, pharmacokinetics, and fundamental bio-cytology [14], [15].

2.1. Overview of 3D bioprinting technologies

The technologies of 3D bioprinting are classified into inkjet [16], extrusion [17], [18], and vat photopolymerization-based bioprinting [19], [20], depending on the method employed for dispensing bioinks. The following sections describe these bioprinting technologies.

2.1.1. Inkjet bioprinting

Inkjet bioprinting originates from conventional desktop inkjet printing. Based on the droplet generation mechanism, this method can be classified into two categories: ➀ continuous inkjet bioprinting (CIJ); and ➁ drop-on-demand (DOD) inkjet bioprinting. DOD is considered more suitable due to its superior printing accuracy and bioink utilization and its lower risk of bioink contamination compared with CIJ [21], [22]. DOD can create droplets using thermal, piezoelectric, or electrostatic approaches. Of these, the preferred choice is thermal inkjet bioprinting, which involves heating localized bioinks to generate hot bubbles that expel droplets, achieving a droplet resolution of 30-80 μm. Although the heat generated during bioprinting may result in the loss of a small number of cells in the bioink, the final cell viability after bioprinting can still reach up to 90% [23], [24]. In summary, DOD is distinguished by its high resolution, micro-dropletization, high throughput, non-contact nature, and minimal consumption of bioinks. Consequently, it has found widespread applications in the bioprinting of living tissues, such as blood vessels, heart [25], bone [26], cartilage [27], and liver [28], [29]. Notably, the nozzle must be sufficiently small to ensure high precision in DOD bioprinting, restricting the use of low-viscosity bioinks (≤ 30 mPa⋅s). In addition, the issue of clogging associated with small nozzles imposes a limit on cell density in bioinks, typically restricting it to lower than 10⁶ cells per milliliter [2].

2.1.2. Extrusion-based bioprinting

Extrusion-based bioprinting stands as the most extensively employed method among all bioprinting techniques. It addresses the limitations of inkjet bioprinting concerning bioink viscosity and cell densities, enabling the bioprinting of bioinks with high viscosity (∼600 kPa⋅s) and cell densities reaching up to 10⁸ cells per milliliter [12]. In extrusion-based bioprinting, bioinks form seamless cylindrical filaments (150-350 μm) under pneumatic pressure or mechanical means (piston or screw). These filaments are extruded from the nozzle onto a substrate and are subsequently crosslinked to form mechanically robust structures through pH or temperature changes, enzymatic reactions, chemical processes, or exposure to light, heat, or ions [30], [31]. Extrusion-based bioprinting exhibits superior scalability compared with other methods and can accommodate high-viscosity and high-cell-density bioinks. Consequently, it has found extensive applications in the bioprinting of solid organs such as the heart [32], [33], [34], kidneys [35], [36], liver [37], and spleen. However, extrusion-based bioprinting presents several challenges, including low resolution (100 μm), nozzle clogging susceptibility, requirements for shear-thinning materials, and the potential for shear-induced cell death, impacting cell viability (40%-95%) [18], [30], [38], [39].

Coaxial extrusion printing, a prevalent method within extrusion bioprinting, utilizes a specialized nozzle with an internal lumen and external compartment. This design enables the simultaneous extrusion of bioink and crosslinking agent solutions, leading to rapid gelation in a core-shell manner at the dispensing head. This printing approach enables the efficient fabrication of perfusable tubular structures using a dual coaxial nozzle [40]. In addition, a tri-coaxial nozzle design offers the capability to print more biomimetic multilayered heterogeneous vascular channels [41]. However, it is difficult to use coaxial printing strategies to construct branched microvascular channels due to limitations in generating continuous fibers.

2.1.3. Vat photopolymerization-based bioprinting

Vat photopolymerization-based bioprinting is a technique that employs ultraviolet light to selectively induce gelation in bioinks. It includes stereolithography (SLA), digital light processing (DLP), and two-photon polymerization (2PP) bioprinting. Among these, the most common method is SLA, which uses a light beam to crosslink photocurable bioinks in a point-by-point manner. After each layer is cured, the platform descends a specific height to cover the cured bioink layer with sufficient bioink for the crosslinking and curing of subsequent layers. This method allows for faster and more precise printing of both internal and external tissue structures compared with inkjet and extrusion-based bioprinting. However, due to the point-by-point curing with a laser beam, the printing speed is slower, and it is more challenging to print scalable structures [42]. Moreover, the use of photoinitiators in bioinks introduces some cytotoxicity [43]. 2PP achieves high-precision crosslinking through the selective irradiation of voxels with two-photon irradiation. It typically produces structures with dimensions ranging from a few hundred micrometers to a few millimeters, making it suitable for printing complex structures at the photonic level and the nanoscale. However, the technology’s high operating cost poses challenges for widespread adoption [44].

2.1.4. Others

In addition to the traditional 3D bioprinting methods mentioned earlier, researchers have pioneered various innovative techniques by seamlessly integrating multiple technologies tailored to specific applications. One of these advancements is acoustic bioprinting, in which an open bioinks reservoir employs an acoustic brake to generate circular sound waves, focusing on the bioink at the nozzle’s tip. This unique process facilitates the formation of cell-encapsulated picoliter droplets, which are subsequently sprayed onto the substrate. Acoustic bioprinting provides precise control over cell placement, and its nozzle-free design mitigates clogging issues, providing enhanced protection for cells [45].

Another groundbreaking technology is magnetic bioprinting, which necessitates the prior magnetic labeling of cells and controlled patterning of cells or cell spheroids through the application of a magnetic field. This technique enables meticulous spatial control, allowing for the rapid and precise bioprinting of various tissue structures. Successful creation of tissues such as vascularized lungs and aortic valves has been demonstrated through this method [46], [47].

Bioprinting in a suspension bath is based on traditional extrusion-based bioprinting but incorporates suspended media. The printhead prints within this bath, balancing the bioactivity and printability of the bioinks. This approach is not constrained by printing location, enabling the high-resolution reconstruction of intricate 3D structures. Recent applications have showcased the 3D bioprinting of full-size human hearts [48] and the direct production of small blood vessels contributing to vascularization within functional tissues [46].

2.2. Importance of bioinks in organ bioprinting

Bioinks play a pivotal role in organ bioprinting by providing a conducive microenvironment and structural support for crucial cellular processes such as adhesion, migration, proliferation, and differentiation. Polymers are a major component of bioinks, and the ideal polymers must possess specific properties aligned with various bioprinting technologies. These properties include printability, tissue specificity [49], biocompatibility [50], shear-thinning behavior [51], [52], appropriate crosslinking mechanisms [51], [52], degradation rates [53], suitable viscosity [19], and tissue-mimicking mechanical properties [54], [55]. Polymers, which are classified into natural and synthetic categories based on their sources, form the foundation of these bioinks [56].

2.2.1. Natural polymers in bioprinting

Natural polymers are usually protein or polysaccharide-based; natural polymers used in bioprinting include collagen, fibrinogen, silk fibroin, gelatin, sodium alginate, hyaluronic acid, and decellularized extracellular matrix (dECM) [57]. Natural polymers are favored for 3D bioprinting due to their tissue-matched physicochemical properties, high biocompatibility, low cytotoxicity, porous structure, and the presence of essential bioactive molecules [58]. For example, collagen, the most abundant protein in humans, has been utilized for bioprinting bladder smooth muscle cells [59]. Gelatin, another natural bioink, is advantageous for organ bioprinting due to its low immunogenicity, cell adhesion structure, and cost-effectiveness. Gaetani et al. [60] encapsulated human cardiac progenitor cells in a hyaluronic acid/gelatin composite bioink to create a cardiac patch. Implantation of the patch improved cardiac function, supporting the long-term survival and proliferation of human cardiomyocytes (CMs) with subsequent differentiation into cardiac and vascular structures.

dECM derived from tissues is treated using chemical, physical, and enzymatic methods to remove cells without compromising the extracellular matrix (ECM) and is then utilized to reduce the antigenicity of bioinks [61], [62]. The composition and structure of dECM closely resemble those of its tissue of origin, so it excels in replicating complex, cell-compatible microenvironments, in comparison with many traditional bioinks. This proficiency in emulating the native ECM facilitates improved cellular function and the reconstitution of the intricate ECM complexities inherent in natural tissues [63]. Biological ECM scaffolds rapidly degrade and are replaced in vivo, producing downstream bioactive molecules with bio-conducive properties [64]. 3D bioprinting based on dECM bioinks has made great strides in the regeneration of the heart, muscle, cartilage, skin, and kidney [65]. Consequently, decellularized organs offer ideal transplantable scaffolds that contain the microstructural and extracellular cues necessary for cell attachment, differentiation, blood vessel formation, and function [66]. However, the reproducibility of dECM production may be affected by large batch-to-batch variations in natural tissue sources [67]. Moreover, scaffolds implanted with naturally derived bioinks often exhibit weak mechanical properties, making it challenging to match the 3D structure of implanted tissues or organs.

2.2.2. Synthetic polymers in bioprinting

Synthetic polymers are human-made polymers created through chemical reactions of monomers. In bioprinting, their networks comprise inert units that provide superior mechanical properties and immunogenicity [68]. The molecular weight and physical properties of synthetic polymers can be flexibly adjusted to meet specific bioprinting requirements. However, some synthetic polymers exhibit poor biocompatibility and generate toxic degradation products because of the absence of functional groups and structural complexity, limiting their application in organ bioprinting.

Notable examples of synthetic polymers used in bioprinting include polyethylene glycol (PEG), polyurethane (PU), and polycaprolactone (PCL). PEG offers hydrophilicity, biocompatibility, and non-immunogenicity [69] and has been commonly employed as a photo-crosslinkable bioink for bioprinting. Its crosslinking principle involves grafting its two hydroxyl groups to other functional groups (acrylate groups, thiols, etc.) through physical radiation, free-radical polymerization, or enzymatic or covalent interactions. This flexibility allows the modulation of its mechanical properties during bioprinting [70]. However, the low viscosity of PEG makes it unsuitable for extrusion-based 3D bioprinting, and its biological inertness results in poor cell adhesion. Researchers have sought to enhance its cellular activity by doping it with bioactive molecules [71], and it is primarily used in the bioprinting of blood vessels [72], [73], [74]. PU has a distinctive segmented structure and a diverse range of physical and mechanical properties. Depending on the synthesis method, it can be a biodegradable or non-biodegradable biomaterial. Biodegradable PU can be utilized alone or in combination with other biomaterials for bioprinting [75], [76]. PCL is a semi-crystalline polymer with excellent rheological properties and viscoelasticity when heated. It is frequently employed as a raw material for preparing organ-chip devices [77]. Researchers [78] have also utilized PCL as a support structure material for the bioprinting of soft hydrogels for cartilage and muscle tissue regeneration.

2.2.3. Choice of appropriate cell types

The inclusion of appropriate cell types is paramount for recapitulating the architecture and function of biological tissues/organs, which consist of diverse cell populations collaborating synergistically. To emulate cellular heterogeneity, organ bioprinting typically involves multiple cell types. For example, CMs are the dominant cells in the human heart responsible for myocardial contraction. Non-myocytes, such as fibroblasts (FBs), endothelial cells (ECs), and immune cells [79], also play distinct and important roles in heart development and function [80]. To mimic the heterocellular composition and interactions, recent cardiac tissue bioprinting efforts have employed a mixture of non-myocytes and CMs [81], [82], [83].

To select cells for bioprinting, one straightforward approach is to explore the use of primary cells found in native tissues, which typically include highly specialized, terminally differentiated cells and progenitors of specific lineages. However, the isolation and in vitro expansion of certain primary cell types supporting bioprinting remains challenging [49]. Concurrently, human multipotent or pluripotent stem cells, such as adult stem cells (e.g., mesenchymal stem cells (MSCs), adipose-derived stem cells (ADSCs), human embryonic stem cells (ESCs), and induced pluripotent stem cells (iPSCs)), have garnered significant attention due to their self-renewal capability and potency. These cells, whether pre-differentiated or undifferentiated, have been applied in bioprinting across a wide range of tissues and organs, including the heart, liver, pancreas, neurons, vasculature, and various cancer models [57]. Furthermore, patient-derived autologous stem cells hold great potential in bioprinting personalized organs for transplantation purposes, mitigating immune rejections, or establishing patient-specific disease models. iPSC technology has enabled the creation of patient-specific stem cell lines from different mature cell types obtained with minimal invasion [84]. Recent studies have differentiated iPSCs into cells of various lineages such as CMs, neurons, hepatic cells, and ECs [85], [86], The field of bioprinting has witnessed a trend of utilizing iPSC-derived cells, such as for the fabrication of cardiac [87], hepatic [43], and pancreatic [88] tissues. Despite the merits of this approach, unaddressed problems in iPSC technology include the risks of tumorigenesis and mechanical sensitivity of undifferentiated stem cells, as well as cell maturity derived from stem cells [89], [90].

While autologous cells offer advantages in terms of safety and providing patient-specific genetic backgrounds, their application is limited by the patient’s health conditions and involves complex manufacturing procedures and long lead times [91]. Allogeneic cells from universal donors have emerged as promising alternatives with great potential for the scalable production of standardized cell lines. Along with systemic immunosuppression, strategies have been proposed to address immune rejection issues [91], [92], [93], [94]. More specifically, universal iPSCs have been developed through gene manipulations [95], [96], [97]. These universal allogeneic cells could serve as a reliable source for the off-the-shelf production of human grafts. Moreover, animals are promising sources of organs, tissues, and cells to address current unmet needs in clinical settings. Recent studies in xenotransplantation have made progress in overcoming xenogeneic immune barriers [98] and the application of immuno-privileged xenogeneic MSCs [99], [100]. Xenogeneic cells might also be used in specialized applications requiring immune alerting, such as cancer wound repair [101], [102].

An underexplored frontier is the integration of solid organ bioprinting with the latest advancements in cell engineering, which would unlock the potential to create artificial organs that surpass mere recapitulation by incorporating human-designed features. For example, stimuli-responsive functions can be established to help researchers gain control of cellular behaviors for disease studies using bioprinted models [103]. Optogenetic engineering has achieved high-resolution control of cell differentiation, enhancing recapitulation [104], [105]. Moreover, functionally enhanced tissue substitutes can be designed to ameliorate a detrimental microenvironment, thereby improving engraftment.

3. Recent advancements in organ bioprinting

3.1. Heart bioprinting

Cardiovascular disease stands as the predominant contributor to global mortality, being responsible for one-third of annual deaths across the world. This category encompasses a diverse spectrum of disorders affecting the heart and blood vessels, such as coronary heart disease, cerebrovascular disease, rheumatic heart disease, and congenital heart disease [106]. Particularly noteworthy is myocardial infarction, which is marked by hypoxia-induced CM loss, tissue degeneration, fibrosis, and eventual heart failure, constituting up to 13% of global deaths [107]. Therefore, engineering cardiac tissues that accurately mimic the heart’s complex nature is crucial for improving our understanding of cardiovascular disease and developing effective treatments.

A multitude of cardiac structures have been created through the application of bioprinting, aimed at reconstructing different components of the human heart, including the myocardium [108], [109], vasculature [110], [111], and heart valves [112]. These constructs have been developed for diverse applications, such as heart-on-a-chip platforms for drug screening [108], [113], [114], in vitro disease models [87], cardiac patches for myocardial infarction tissue restoration [60], [115], [116], [117], and organ-scale heart pumps for future replacement therapies [48], [82]. This section provides a review of recent advancements in the 3D bioprinting of cardiac tissues, highlighting the increasing complexity of these constructs and the associated challenges.

The human heart—a critical organ for blood circulation—consists of four chambers, with two atria receiving blood and two ventricles pumping it out. Its wall is structured in three layers: the pericardium, myocardium, and endocardium. While CMs, which account for 30%-40% of heart cells, are essential for heart contractions [118], other cells such as FBs and ECs also contribute to heart function [119]. The limited self-renewal of adult CMs poses challenges for bioprinting [120]. Stem cell therapies, which provide new CMs for heart failure, depend heavily on the stem cells’ paracrine effects rather than direct maturation into functioning CMs [120]. Bioprinted cardiac tissue often uses pre-differentiated cells from human ESCs or iPSCs [109], [113], [121], but replicating the mature characteristics of adult CMs in vitro is difficult [122].

In bioprinting, bioinks are sought to incorporate pre-differentiated cells within polymer matrices, mimicking myocardial microenvironments (e.g., biochemical compositions, mechanical properties) to promote CM maturation pre- and post-implantation. Gelatin and methacrylate gelatin (GelMA) have been explored for their suitability in this regard [123], [124]. Cardiac-derived dECM bioinks, which foster maturation and improve heart function, have also been developed [32], [125]. Mechanical properties, including matrix stiffness, influence CM maturity, affecting myofibril protein expression and sarcomere alignment [122]. Efforts have been made to develop mechanically tunable dECM bioinks that resemble cardiac biomechanical properties and are optimized for bioprinting [125].

To mimic the developmental environment of the heart, recent strategies involve bioprinting CMs with supportive cells, such as FBs and ECs. This approach, which includes using gelatin/alginate composite bioinks, promotes cell survival, proliferation, and functional maturation [81]. Furthermore, cardiac spheroids, which offer a representation of native tissue structures and interactions, have been integrated into bioprinting to build complex 3D cardiac constructs, enabling the creation of scaffold-free myocardium with high cellular density [83], [126], [127], [128]. These spheroids have been used in cardiac patches, displaying spontaneous beating and vascular formation, with in vivo studies indicating enhanced regeneration and angiogenesis [127].

Spheroid-based bioprinting not only achieves cardiac-like high cell densities but also facilitates the engineering of spatial heterogeneity in cardiac tissues. For example, Daly et al. [83] developed a reductionist model of focal cardiac fibrosis using CM/FB spheroids. By manipulating the ratio of iPSC-derived CMs to FBs in the spheroids (4:1 for “healthy” and 1:4 for “scarred”), the researchers created microtissue rings exhibiting spatially controlled CM and FB ratios with distinct properties in healthy and scarred areas, including variations in contraction and electrophysiological characteristics. These microtissue rings can serve as a model for cardiac fibrosis, enabling the study of therapeutic interventions for cardiac repair.

The native myocardium features anisotropically aligned muscle fibers that contribute to left ventricular torsion and are essential for normal ejection function [129]. Several approaches based on bioprinting have been reported to induce oriented cell alignment in the myocardium. Zhang et al. [108] introduced a scaffold-based method in which the bioprinting of EC-embedded micro-fibrous scaffolds guided the orientation of subsequently seeded CMs, enabling the fabrication of endothelialized myocardium. In addition to surface seeding, anisotropic myofibers have been created using microscale continuous optical printing, wherein cells were encapsulated inside patterned GelMA scaffolds [130]. The aligned fibers, with ventricular CMs encapsulated in a 3D environment, generated nearly twice the force compared with two-dimensional (2D)-seeded controls. Patterned scaffolds with different alignments exhibited varied contractile forces, all surpassing the controls on flat surfaces. Tsukamoto et al. [131] developed orientation-controlled cardiac tissue using a layer-by-layer method combined with bioprinting. In this approach, cell direction was manipulated by linearly shaping the cardiac tissue with a 3D-printed gel frame. A novel approach for programming cellular alignment was utilized to engineer anisotropic organ building blocks (aOBBs) made of iPSC-CMs, which were incorporated into a compacted bioink for extrusion-based bioprinting [109]. The aOBBs consisted of elongated microtissues composed of aligned CMs. During extrusion, the shear and extensional forces resulted in the alignment of aOBBs along the print path, forming macro-filaments containing oriented CMs (Fig. 2(a)).

Establishing functional vascular systems is imperative for the bioprinting of thick myocardial tissues. The native heart, being highly vascularized, features hierarchical vessels that transport nutrients and oxygen and support multicellular crosstalk, tissue maturation, and functional maintenance [90], [133]. Considering the tissue’s oxygen diffusion limit of approximately 200 μm, the need for vascularization strategies becomes evident when bioprinted myocardial patches surpass this limit [120]. Various strategies have been proposed for engineering vasculatures of different scales, ranging from microscale capillaries to large-scale aorta [110], [111], [134], [135], [136], [137], [138]. For microscale vessel formation, numerous studies have leveraged the self-assembly nature of ECs or endothelial progenitor cells to engineer vascularized cardiac spheroids or patches [81], [117], [121], [126], [136], [138]. ECs can be bioprinted alongside CMs and other supportive cells to create vascularized cardiac constructs. For example, Zhu et al. [132] utilized microscale continuous optical bioprinting to fabricate prevascularized tissues with intricate 3D microarchitectures (Fig. 2(b)). The tissues successfully encapsulated diverse cell types and mimicked the native vascular compositions, within hydrogels exhibiting precise spatial control, thereby obviating the necessity for sacrificial materials or perfusion techniques. Additional factors contributing to microvessel formation and maturation, such as biochemical cues (e.g., vascular endothelial growth factor) and mechanical stimuli [133], can be combined with a cell co-culture approach to engineer vessels. For example, Lu et al. [136] introduced an electrical stimulation method to enhance microvasculature formation within cardiac constructs through the co-culture of iPSC-CMs and human umbilical vein endothelial cells (HUVECs). The electrical stimulation significantly promoted HUVEC elongation, migration, and interconnection; increased the expression of related genes; and enhanced interaction with CMs. Mesoscale vasculatures can be created using multiplexing “top-down” approaches, including layer-by-layer assembly, sacrificial templating, indirect bioprinting, embedded bioprinting [110], SLA, diffusion-induced gelation, and coaxial bioprinting [139], as extensively described by Ref. [133].

Heart valves, which encompass tricuspid, bicuspid, pulmonary, and aortic valves, play a crucial role in guiding unidirectional blood flow through the four cardiac chambers. In cases of severe conditions such as congenital heart valve defects, surgical replacement of defective valves is often required. A study has [140] undertaken the bioprinting of aortic valves, employing distinct bioinks to replicate the anatomy and heterogeneity essential for valvular functions. The work of Butcher’s group [34], [38], [141] has yielded significant advancements in this realm. These researchers have successfully bioprinted heterogeneous aortic valves using PEG-diacrylate (DA), maintaining near 100% viability of seeded interstitial cells for up to three weeks. They also developed living alginate/gelatin hydrogel valve conduits that encapsulated spatially patterned dual cells (smooth muscle cells and aortic valve leaflet interstitial cells) with high viability. Another achievement was tri-leaflet methacrylated hyaluronic acid/methacrylated gelatin valve conduits that supported the survival and ECM deposition of embedded interstitial cells. The use of Type I collagen hydrogel as a bioink for heart valve bioprinting has also been explored [142]. Constructs bioprinted with MSCs demonstrated remodeling transitions after rat subcutaneous implantation. In a recent development, Immohr et al. [112] created 3D models for investigating the pathogenic mechanisms of calcified aortic valve disease. These constructs facilitated a 21-day co-culture of valvular interstitial cells and valvular ECs in an alginate/gelatin hydrogel, mimicking the heterogeneous nature of valves (Fig. 2(c)).

Recent strides in bioprinting have ushered in the ability to construct scalable soft tissues, marking progress toward the engineering of full-size human hearts. A notable methodology presented by Lee et al. [143] involves the 3D bioprinting of unmodified collagen into intricate heart components using the freeform reversible embedding of suspended hydrogels (FRESH). The FRESH method allows for the fabrication of filaments at high resolutions (∼20 μm) and has been employed to engineer multiscale vasculature and neonatal-scale human hearts with high fidelity. Building upon this, a full-size model of an adult human heart was created based on patient-specific magnetic resonance imaging (MRI) data and bioprinted with alginate using FRESH (Fig. 2(d)) [48].

More recently, FRESH was successfully employed to develop heart tubes capable of pumping fluid, emulating the ventricles’ natural function [82]. These heart tubes were cellularized with ESC-derivedCMs (ESC-CM) and FBs, forming perfusable constructs that exhibited synchronous contractions for over four weeks. Observations revealed dense layers of interconnected CMs on the tubes’ surface expressing markers of z-lines in the myofibrils, specifically sarcomeric α-actinin. The tubes demonstrated spontaneous and anisotropic conduction, with a typical ESC-CM conduction velocity of approximately 5-6 cm·s⁻¹, effectively displacing fluid within the lumen through each contraction. Despite lacking anatomical complexity and achieving a lower pumping velocity compared with native hearts, this groundbreaking work established the feasibility of and marked an initial stride toward the engineering of functional human heart pumps. Addressing the challenge of replicating similar electrical characteristics is crucial for constructing functional cardiac patches, as mismatched electrical properties between engineered and host tissues may lead to re-entry arrhythmias or conduction blocks [120]. A prospective strategy involves the coupling of bioprinting with conductive materials to mimic the electrophysiology of the human heart [106].

Current investigations have unveiled numerous challenges in the pursuit of bioprinting ideal cardiac tissues tailored for specific applications. It is notable that the criteria for defining “ideal” tissues and the associated challenges differ across applications. While certain objectives, such as tissue maturity and heterocellular interactions, are commonly shared, other objectives exhibit variations or even contradictions. For example, the focus on drug screening underscores the need for reproducible, high-throughput construct generation. This preference leans toward miniaturized cardiac tissues with simplified geometry and composition while preserving a minimal set of essential structural and functional features [90]. Conversely, for tissue replacement, complexity is paramount in faithfully replicating the inherent architecture and functions of the heart. Challenges such as engineering anisotropy, vascularizing thick tissue, establishing conductive properties, and scaling up construct volume have emerged as critical hurdles. Therefore, it is imperative to establish “ideal” standards in order to guide the judicious fabrication of application-specific constructs.

3.2. Liver bioprinting

The liver’s intricate cellular architecture supports diverse metabolic and synthetic functions. Its main cells are hepatocytes, which are organized into hepatic lobules, featuring a central vein from which cell cords extend to the periphery, where portal triads reside [144]. Fenestrated ECs form sinusoids between these cords, allowing efficient substance exchange and creating a zonation of hepatocyte functions. Kupffer cells eliminate pathogens and debris in the sinusoids, stellate cells regulate fibrosis, and cholangiocytes manage bile. These cells collectively regulate glucose, lipid, protein, cholesterol, and vitamin metabolism, synthesize serum proteins, detoxify substances, and more [144], [145].

Hepatocytes’ regenerative capacity allows the liver to rapidly recover from damage [146]. Tissue engineering has aimed to replicate this ability, but traditional methods have failed in terms of volume and architecture fidelity [147], [148]. 3D cell culture methods better emulate liver tissue structure and biomechanics, advancing the in vitro bioengineering of liver tissues [149], [150]. Early research by Kang et al. [151] extended the culture life of primary hepatocytes using alginate scaffolds, while Wang et al. [152] employed rapid prototyping to fabricate hepatocyte-gelatin structures. Arai et al. [28] developed liver constructs using artificial scaffolds and an inkjet bioprinter, whereas Ma et al. [43] created a 3D hydrogel-based triculture model for drug screening and disease modeling (Fig. 3(a)). These advances overcome the limitations of traditional in vitro cultures, which is crucial for drug screening and disease mechanism investigation [155].

The development of a widely accessible liver model is particularly important, to provide a platform that can faithfully replicate the preclinical in vivo human environment. Such a model has now become an indispensable tool for drug screening and gaining insights into disease mechanisms. Recent developments in 3D cultivation have shown promise. Chang et al. [156] developed a micro-organ device that allowed for dynamic drug screening, Bhise et al. [157] designed a liver-on-a-chip platform that supported long-term hepatocyte culture, Norona et al. [158] reported a model for investigating hepatic fibrosis and the roles of Kupffer cells, and a 3D-printed model by Sun et al. [159] revealed differences in tumor-related gene expression for antitumor drug screening. This research highlights the potential of 3D-bioprinted models to enhance drug-screening accuracy and elucidate drug-resistance mechanisms.

In the realm of liver bioprinting, the challenge lies in constructing liver lobules and intricate vascular systems. Although some research has shown successful hepatic structures, few studies have developed in vitro liver sinusoid models with integrated vascular networks. Several studies have addressed this challenge by incorporating vascular networks within the 3D architecture of engineered livers using techniques such as sacrificial bioprinting, extrusion-based bioprinting, and coaxial bioprinting. Kang et al. [151] utilized a preset extrusion bioprinting technique to successfully fabricate hepatic lobules within a highly vascularized structure. Hassan et al. [160] utilized multi-material bioprinting for complex tissues that can be implanted post-printing while retaining their 3D architecture in vivo. In a recent study, Maji et al. [161] presented a single-step bioprinting technique to concurrently print a chip platform and generate a perfusable vascularized liver sinusoid in vitro model (LSOC-P). The LSOC-P exhibited significantly enhanced hepatocyte viability, proliferation, and liver-specific gene and protein expression compared with other 3D and 2D models. These studies indicate significant progress in addressing liver tissue complexity, showing promise for future therapeutic applications.

3.3. Kidney bioprinting

Bioprinting technology has emerged as a promising avenue for advancing kidney regeneration by engineering renal constructs. Researchers have explored innovative approaches to create 3D structures that mimic the intricate architecture of the kidney, aiming to address the critical shortage of donor organs for transplantation. One notable study by Yu et al. [139] utilized a patient derived cells in a bioink to bioprint functional vascularized renal tissue. This groundbreaking work demonstrated the feasibility of constructing complex renal structures with reduced immune rejection and ehnahced biocompatability, paving the way for potential applications in kidney regeneration. Similarly, Jia et al. [162] employed bioprinting to fabricate renal constructs by incorporating renal cells. Their study emphasized the importance of preserving the viability and functionality of renal cells within bioprinted constructs, laying the foundation for the development of bioengineered kidneys for transplantation and regenerative medicine. These advancements in bioprinting renal constructs underscore the potential to revolutionize the field of kidney regeneration, offering new avenues for therapeutic interventions and addressing the challenges associated with organ transplantation.

Bioprinting glomerular and tubular structures for kidney regeneration involves careful consideration of various factors to achieve functional and anatomically accurate renal tissue. The complexity of the glomerulus, which is responsible for blood filtration, requires the selection of suitable bioinks and bioprinting techniques. Studies by Homan et al. [35] and Zhang et al. [108] emphasize the importance of mimicking the intricate architecture of the glomerular filtration barrier, which consists of podocytes, ECs, and mesangial cells (Fig. 3(b)). In addition, it is crucial to address the biomechanical properties of the glomerular structure for successful filtration function. When bioprinting tubular structures, which play a vital role in reabsorption and secretion, the necessary considerations include choosing bioink materials that are compatible with the unique functions of proximal and distal tubules.

Notable progress has been made in recent years on bioprinting functional nephrons, the structural and functional units of the kidney, offering potential solutions for kidney regeneration and transplantation. Studies such as those by Homan et al. [35] and Jia et al. [162] showcase advancements in creating intricate renal structures, including glomerular and tubular components, with the aim of replicating the complex filtration and reabsorption functions of nephrons. The integration of vascularization, which is essential for proper perfusion and functionality, has been a key focus in overcoming the hurdles associated with limited nutrient diffusion. Despite these advancements, several challenges persist. Achieving the precise cellular composition and architecture of nephrons, especially the intricate glomerular filtration barrier, remains a significant hurdle. Furthermore, ensuring long-term functionality, cellular maturation, and immuno-compatibility in vivo poses ongoing challenges. Continued collaborative efforts between researchers in materials science, biology, and medicine are essential to address these hurdles and propel the field of bioprinting toward the creation of fully functional nephrons for effective kidney regeneration.

3.4. Lung bioprinting

The lungs, which are optimized for gas exchange, consist of the conducting zone with the trachea, bronchi, and bronchioles, lined by the pseudostratified ciliated epithelium that facilitates air conduction and mucociliary clearance [163]. Transitioning to the respiratory zone, the lining changes to simple cuboidal and squamous cells in the respiratory bronchioles and alveoli [164], [165]. The alveoli, where gas exchange occurs, are lined with Type I pneumocytes for diffusion and Type II pneumocytes that secrete surfactant to reduce surface tension [166]. Collectively, the highly specialized structures of the lungs work synergistically to accomplish the vital physiological process of respiration.

For bioprinting, the lungs are best approached by dividing them into distinct units with unique cavities, similar to other hollow organs [167]. The goal of lung bioprinting is to create personalized, physiologically accurate pulmonary tissue tailored to individual patients’ needs, which necessitates a deep understanding of lung physiology and defined functional parameters. The notable advantage of bioprinting in lung tissue engineering lies in its ability to construct personalized functional compartments, replicate physiological activities, and accurately reproduce the 3D architecture of the organ.

Although some success has been achieved in the bioprinting of human lung alveolar models, replicating the complex lung structure remains a formidable challenge, limiting the scope of regenerative lung tissue research [168]. Kim et al. [169] presented an innovative methodology for tracheal reconstruction, offering a promising approach for the long-term functional restoration of segmental tracheal defects. Kang et al. [170], [171] utilized inkjet bioprinting to construct a 3D alveolar barrier model composed of four human alveolar cell lines. da Rosa et al. [172] obtained Wharton’s jelly MSCs isolated from human umbilical cords, differentiated them into various lung cell types and subsequently bioprinted them using an alginate/gelatin bioink. Ng et al. [173] employed DOD bioprinting to construct human triple-layered alveolar lung models consisting of human lung epithelial cells, human ECs, and human lung FBs. These bioprinted alveolar lung models maintained cell viability and exhibited proliferation profiles similar to those of non-bioprinted cells, providing an attractive and scalable tool for realistic and reliable in vitro disease models of the human lung.

Reconstructing the vascular network is essential for the survival and growth of transplanted tissue cells, while ensuring effective gas exchange is a key aspect in the development of 3D-bioprinted lung tissues. Grigoryan et al. [153] engineered 3D transport systems with photopolymerizable hydrogels, demonstrating the potential for creating complex vascular networks and functional valves (Fig. 3(c)). Horváth et al. [174] developed a bioprinting technique for a 3D alveolar model that closely mimics the human air-blood barrier; the technique offers advantages over manual construction for high-throughput screening despite a brief three-day survival period.

Overall, refining the parameters of each functional unit is crucial for optimizing bioprinting technologies for the lungs. This optimization is essential to achieve improved reconstruction of the intricate functional units and microenvironments within the lung tissue.

3.5. Pancreas bioprinting

The pancreas is a large and intricate gland that plays a dual role. Its exocrine part, which makes up the majority of the organ, produces digestive enzymes, while the endocrine pancreas houses the islets of Langerhans, which consist of over five cell types crucial for maintaining glucose homeostasis in the body [175]. However, the bioprinting of the pancreas is nascent, with most research focusing on islet encapsulation and pancreatic cancer models.

Type 1 diabetes (T1D) is a major concern, and cell replacement via islet transplantation is a promising treatment, albeit one that presents challenges such as donor shortages, immune rejection, post-transplant islet loss, hypoxia, and scaling of islet-laden devices [176]. Current studies address one or more of these challenges by integrating islet transplantation with bioprinting technologies. For example, hypoxia poses a significant risk of graft failure post-transplantation, as metabolically active islets require a high oxygen supply to survive [177], [178]. To combat this issue, 3D bioprinting is being applied to create oxygenation-enhancing structures. For example, Farina et al. [179] developed a vascularized, porous encapsulation device using bioprinting that protected islets from acute hypoxia post-transplantation.

Another strategy to address donor scarcity is the generation of insulin-secreting β cells from human iPSCs. For example, Song et al. [180] bioprinted a macroporous device housing iPSC-derived β cells and found that the microenvironment is key for cell specification. Kim et al. [181], [182] proposed a bioink formulation based on pancreatic dECM (pdECM) to recreate the inherent microenvironments that enhance insulin secretion and the maturation of iPSC-derived cells. The pdECM-based bioink was further applied in the fabrication of iPSC-derived pancreatic islet-like aggregates. Combined with a porous PCL outer shell, it formed a hybrid macro-encapsulation system for subcutaneous transplantation [88]. Klak et al. [183] bioprinted bionic pancreatic tissues using the pdECM-based bioink; the tissues showed biocompatibility and functional islet formation in vivo.

One limitation of bioprinting in this context is that shear forces may be experienced during bioprinting, which can compromise islet viability [184]. Thus, researchers are exploring alternative bioprinting methods. For example, Wang et al. [154] used SLA for higher-resolution pancreatic tissue construction (Fig. 3(d)).

In pancreatic cancer research, 3D bioprinted constructs are invaluable for pathology studies and drug testing [185]. Hakobyan et al. [186] utilized laser-assisted bioprinting to develop exocrine spheroids for studying pancreatic ductal adenocarcinoma. These bioprinted spheroids, which were composed of both acinar and ductal cells, could function as 3D array models for studying the initiation of pancreatic ductal adenocarcinoma. Similarly, Huang et al. [187] created models to replicate tumor-stroma dynamics, demonstrating their drug testing potential.

The 3D bioprinting of pancreatic tissues that closely resemble their native counterparts—considering the highly complex composition and anatomy of the pancreas and the functions of both its endocrine and exocrine components—is still a significant challenge. While substantial efforts have been directed toward islet replacement strategies, driven by the high demand for addressing T1D, other aspects of the bioprinting of pancreatic tissues have undergone limited exploration. There is a pressing need for pancreatic tissues, whether as in vitro models or as a potential source for pancreas donors in critical situations where the replacement of the entire pancreas, rather than just islets, is required. Thus, further research and development are essential to achieve comprehensive advancements in the bioprinting of pancreatic tissues.

3.6. Integrating diverse cell types and navigating the complexity of organ structures

The vascularization across various organs shares the fundamental purpose of ensuring the critical supply of nutrients and oxygen that is vital for cell survival and functionality [188], [189]. Common bioprinting strategies leverage the innate ability of ECs and their progenitors to self-assemble into capillary networks [174]. This is often complemented by incorporating supportive cell types such as pericytes and smooth muscle cells, which provide structural support and stability to the forming vessels. Techniques such as co-culturing cells within bioinks, electrical stimulation, and creating microvascular channels using sacrificial materials enhance the formation of vascular networks [190], [191].

The integration of multiple cell types is crucial for reconstructing the complex heterogeneity and hierarchical organization of native tissues [81], [173]. For example, the reconstruction of osteochondral defects through the utilization of articular cartilage progenitor cells and bone MSCs, which has been done by employing a two-channel extrusion bioprinting technique, not only demonstrates the simultaneous regeneration of both cartilage and subchondral bone but also exhibits favorable crosstalk between the cartilage, bone, and vascular components [192]. In the realm of cardiac bioprinting, CMs, FBs, and ECs are employed to both establish vascularization and recreate the myocardium’s conductive and contractile properties [81]. Similarly, hepatic and renal bioprinting combines parenchymal cells with vascular cells to mimic the organ zonation and filtration capacities, respectively [145], [193].

Bioprinting techniques must consider the organ-specific ECM, which provides cues for cell adhesion, migration, differentiation, and function. Bioinks that are formulated to resemble the native ECM are often derived from decellularized organs and support cells [194], [195]. These bioinks require optimization for mechanical strength, biodegradability, and bioactivity to facilitate tissue maturation post-bioprinting.

The complexity of organ structures is further addressed through the recreation of microarchitecture using advanced bioprinting methods, such as SLA for high-resolution patterning, inkjet bioprinting for precise cell deposition, and extrusion-based bioprinting for fabricating larger tissue constructs [151]. Innovations such as spheroid-based bioprinting enable cell-dense constructs to be built that closely mimic in vivo tissue conditions [83].

The reconstruction of organs through bioprinting is a symphony of cellular orchestration and biomaterial engineering. It requires a nuanced understanding of tissue biology, biomaterial science, and the interplay between different cell types. Vascularization commonalities suggest the benefits of a unified approach in facilitating organ tissue perfusion and viability, while the integration of various cell types in organ-specific bioinks underscores the versatility and adaptability needed to recreate the complexity of each organ. These approaches hold promise for advances in regenerative medicine, disease modeling, and drug testing.

4. Challenges and obstacles

4.1. Recapitulation of functionality

Clinically relevant bioprinted organs should not only recapitulate the microenvironment but also function for transplantation. In this pursuit, attempts to reconstruct fully functional organs should be made in vitro along with valid evaluation methods [196]. Meanwhile, in vivo, studies can be conducted to assess the maturity of bioprinted organs and the immune responses of the host. Preclinical trials can be gradually established on a scale ranging from treating small defects to eventually replacing the whole organ [197]. Pre-vascularization and reinnervation will continue to be prerequisites for functional organs [198], [199], so 3D bioprinting techniques should provide high resolution and precision in order to attain perfusable, innervated microstructures.

Recapitulating functionality in bioprinted organs presents intricate technical challenges that demand innovative solutions. To achieve proper functionality, the difficult tasks of optimizing cell density and incorporating multiple cell types in bioprinted organs must be achieved. A critical obstacle is the achievement of precise spatial control over cell placement and the organization of various cell types within the printed constructs in order to accurately emulate the native tissue microarchitecture [83]. This requires the development of advanced bioprinting techniques capable of high-resolution, multi-material deposition to recreate the complex cellular composition and ECM characteristics of tissues. Furthermore, ensuring proper cellular viability, proliferation, and differentiation within the printed structures requires the thorough optimization of bioink formulations, including the selection of biomaterials and growth factors that mimic the native tissue microenvironment [194]. Bioprinted tissues must also replicate the gradual development seen in native organs, which involves the dynamic processes of cellular differentiation, matrix remodeling, and tissue maturation [200]. To achieve this, it is necessary to establish bioactive environments, optimize culture conditions, and incorporate growth factors to guide functional integration. Importantly, robust evaluation methods must be established to assess the functionality of bioprinted organs in terms of parameters such as tissue biomechanics, metabolic activity, and physiological responses to stimuli. Addressing these challenges requires interdisciplinary collaboration in bioprinting technology, biomaterials science, tissue engineering, and biomedical imaging to drive the field toward the development of clinically relevant bioprinted organs.

4.2. Reproducibility and standardization

Although significant advancement in 3D bioprinting has been achieved, there are challenges ahead. The success of bioprinting solid organs relies on the development of tissue- or patient-specific biomaterials that can be developed into bioinks possessing various properties such as high biocompatibility, adequate printability (i.e., viscosity), biodegradability, bioactive properties to promote cell growth and vascularization, and capability to integrate with host tissues [201]. Furthermore, bioinks should preserve cell viability against various bioprinting modalities. To address these demands, various bioink preparation strategies have been introduced, including the use of composite bioinks [202], [203], sacrificial materials [143], [204], and modified dispensing systems (e.g., coaxial nozzles and microfluidic channels) [205], [206] that can bioprint multiple materials at once. However, meeting all these requirements remains a challenge.

Moreover, while the development of technology is accelerating, standardization is not being vigorously established for biomaterials and relevant bioprinting technologies. To address this issue, standardization can be developed in two phases. In the first, internal phase, recommendations can be made to promote the wide-scale use of available standards in research methodologies and study reporting [207]. In the second, external phase, protocols and criteria can be translated and applied within different global/international rules, regions, and nations [207]. Although there is an available bioprinting process standard (ISO/DIS 17296-1) [208], standardization should be expanded more into biomaterials and various bioprinting techniques and updated according to the circumstances as a first step in the establishment of quality-control systems and improvements in clinical applicability.

4.3. Scalability and cost-effectiveness

The scalability and cost-effectiveness of bioprinting techniques are closely related and should be considered for realistic clinical applications. For volumetric organ reconstruction, a substantial number of cells should be prepared, which can be up to 100 million cells [209]. This process can be expensive in terms of time and labor and can require the extensive utilization of reagents [210]. The biomaterials with various growth factors (i.e., ECM and collagen) that are required for bioink preparation require labor-intensive processes for extraction or can be costly if commercially available. Currently available bioprinters are reported to be priced at between 500 000 and 200 000 USD, depending on their capabilities [211]. On the other hand, the billed charges per solid organ transplant could reach over a million USD in 2020 (Table 2) [212], and there has always been a shortage of donor organs [213]. Moreover, there is continuous expense in the years following an organ transplant for postoperative treatments such as replantation or dialysis. These costs are difficult for most individuals to afford, and bioprinting solid organs can be an alternative to alleviate both patients’ financial burden and the organ shortage. In brief, despite the investment costs, it is necessary to systemize and streamline the bioprinting process for cost-effectiveness and improvement in organ transplantation.

4.4. Immune system integration and rejection concerns

The success of organ transplantation can be determined by the integration of the bioprinted organ with the host’s immune system. Given that bioprinted organ substitutes are composed of biomaterials, cells, and other biologics (e.g., growth factors), regulating these components is crucial for suppressing immune responses. As a biomaterial, dECM provides a more native tissue-like microenvironment in comparison with hydrogels. Specifically, dECM is a framework that contains extracellular molecules (i.e., elastin, collagen, laminin, fibronectin, and matricellular proteins) without cells [214], [215], [216], [217], [218]. Additionally, dECM preserves immunomodulatory cytokines, such as basic fibroblast growth factor, transforming growth factor-b, and bone morphogenic proteins [219], [220], [221], [222]. When macrophages respond to a foreign body, they release degradation mediators and secrete pro-inflammatory cytokines such as interleukin (IL)-6 or IL-1b [219]. However, these inflammatory responses can be modulated by the cytokines in dECM, which serve as regulatory factors. Nonetheless, the source of dECM is native organs or tissues, which must be extracted from a human body or livestock. Furthermore, the safety of xenografts or residual cells in dECM remains a risk. In this regard, the interaction between the biomaterials composing the organ substitutes and the host tissues should be thoroughly investigated. 3D bioprinting makes it possible to deposit autologous primary or stem cells in a layer-by-layer manner, with careful modulation of the cell density and arrangement of multiple cell types to minimize immune responses after implantation [121], [201]. Also, utilizing cells to regulate cytokines may be a possible approach to evaluate immune responses in vitro and in vivo [223], [224].

4.5. Ethical and societal considerations

To address bioethical issues, the widespread clinical adoption of bioprinting necessitates the implementation of rigorous regulations for scientific research and medical practice [225], including informed-consent protocols for donation, material manipulation, storage, and applications (commercial and research-oriented). Safety, quality control, and procedural efficiency standards for bioprinting and its products are crucial and required according to human rights considerations. Committees must be established to formulate and oversee guidelines covering the technical, ethical, and legal aspects of the development and application of bioprinting. Robust legal safeguards are essential for all patients, including minors and those unable to provide consent. In addition, regulations governing the commercialization, costs, and penalties of the illicit trafficking of bioprinted human organs and tissues must be put in place.

Still, in contrast to alternative technological solutions such as animal sourcing, the regeneration of human organs seems to raise fewer ethical concerns, thus initially reducing concerns [226]. Moreover, it is clear that bioprinting technology has the potential to transform the restoration and replacement of human tissues and organs. However, ethical oversight is vital across the entire process, from production to end results, alongside considerations of accessibility, fairness, biological compatibility, and engineering obligations to replicate native organs. Fig. 4 illustrates these ethical considerations, including those for cell sourcing, human biology modification, patient autonomy, and the crucial need for equitable access to bioprinted tissues and organs.

5. Future directions

5.1. Promising research directions to overcome existing challenges

The escalating demand for the biofabrication of human-scale volumetric organs, coupled with rapid advancements in bioprinting, emphasizes the imperative need for fabricating functional, transplantable, and intricate tissues and organs. Advancements in bioprinters and bioprinting processes, such as multicellular/multi-material systems [227], showcase the potential for rapid, scalable tissue and organ fabrication. Hybrid bioprinting strategies are envisioned for the future to address the complexity of native tissues and organs. Combinations of various bioprinting techniques can be used to improve the viability and efficiency of the manufacturing process [228]. Furthermore, the developments in multichannel multi-material freeform bioprinting utilizing existing bioinks hold potential for forming large-scale structures and addressing the heterogeneity of native tissues. However, challenges such as prolonged bioprinting times and poor resolution for complex structures require attention. Volumetric bioprinting—an approach for fast printing speed—has great potential for fabricating scalable structures within seconds [229], but it is still limited to low compositional complexity. Thus, combining volumetric bioprinting with other approaches, in “volumetric bioprinting plus” models, offers a promising biofabrication strategy to overcome limitations in compositional complexity. For example, Größbacher et al. [230] recently reported an approach that combines melt electro-writing and volumetric bioprinting to fabricate geometrically complicated objects. This approach has been proven to be successful in fabricating multi-material and multicellular structures. The integration of smart bioprinting processes with artificial intelligence or machine-learning approaches also enhances quality control and process controllability.

The development of novel materials with specific biological functions, high precision, and better compatibility with bioprinters and bioprinting software is another research direction. Deepening the understanding of the natural ECM is a crucial path, as simulating complex ECM components and gradient distributions can lead to the creation of bioactive materials with specific functions [231]. Moreover, biomimicry-based research aims to develop adaptive materials that can spontaneously reprogram their shapes, properties, or functions in response to external stimuli, creating favorable conditions for cell growth and tissue function maintenance [232]. Another research direction is the challenge of vascularization in bioprinted constructs. Synergistic coupling between bioprinting and vascularization strategies is required in order to overcome hypoxia-induced cell death in the central regions of tissues thicker than 100-200 μm [233].

Using patient-derived cells through iPSCs for bioprinting in regenerative medicine is promising, but challenges exist in isolating and differentiating cells, particularly for endothelial progenitors [234]. The time-consuming process and rate-limiting steps, along with the high cell quantities required for bioprinting, emphasize the need for robust differentiation protocols and potentially automated methods to ensure uniform cell populations. Concurrently, there is significant interest in the 3D bioprinting of “organoids-on-a-chip,” where organoids are integrated into microfluidic devices, allowing precise control of the microenvironment for the generation of physiologically relevant structures [235]. This technology offers the precise deposition of cells and materials, potentially enabling customized tissues and vasculature based on individual patient needs. While it is currently limited to specific cell types, future developments should aim to incorporate patient-specific cells to prevent immune rejection.

Another prospective area of research is four-dimensional bioprinting, which permits the fabrication of structures that dynamically alter in response to external stimuli. Here, the fourth dimension of time is combined with 3D bioprinting, allowing printed objects to change their forms or functions in response to outside stimuli, cell fusion, or post-bioprinting self-assembly [236]. This cutting-edge technology greatly enhances the development of complex dynamical structures with high resolution that would not be possible with various 3D bioprinting procedures [237]. In addition, the continued development of advanced biomaterials, higher-resolution bioprinting methods, vascularization strategies, patient-derived cells, and integrated bioprinters with monitoring systems offers the possibility of intraoperative (IOP) bioprinting, in which tissues are constructed directly on the living body, as discussed in the next section.

5.2. Technological innovations and emerging trends in organ bioprinting

Advancements in 3D bioprinting have driven tissue fabrication from patch-level constructs to complex functional organs. Initially, bioprinting focused on single-material structures, but now it integrates multiple biomaterials, cell types, and bioinks to create complex, multicellular architectures [14]. The development of bioinks with tunable mechanical properties and bioactive molecules enables precise control over cellular behavior and tissue formation. Furthermore, advancements in printing technologies, such as extrusion-based, inkjet, and laser-assisted bioprinting, offer varying levels of resolution and scalability to accommodate different tissue types and applications [238]. Other techniques such as volumetric bioprinting and embedded bioprinting have taken biofabrication to the next level, where complex constructs a few centimeters in size can be printed in mere seconds in synergy with functional blood vessels [239]. This transition to larger organs requires bioprinted constructs to include vascularization and innervation, which are essential for sustaining larger tissue volumes. Techniques such as sacrificial printing and bioprinting around pre-existing vasculature facilitate the integration of functional blood vessels within engineered tissues [197]. Similarly, strategies for inducing nerve ingrowth and promoting neuronal connectivity enhance the physiological relevance of bioprinted constructs [240]. Advancements in scaffold design, cell patterning, and organ-on-a-chip technologies have further helped this evolution. Scaffold-free bioprinting techniques, such as spheroid-based assembly and cell sheet engineering, enable the creation of 3D organoids with organotypic structure and function [241]. Moreover, organ-on-a-chip platforms integrate bioprinted tissues with microfluidic systems to mimic organ-level physiology and drug responses accurately [242]. These technical advancements in 3D bioprinting offer unique opportunities for personalized medicine, disease modeling, and drug discovery.

IOP bioprinting is emerging as an advanced approach for integrating the principles of 3D bioprinting with operative processes. In brief, 3D bioprinting has developed from non-biocompatible 3D surgical models, biocompatible prostheses, or organ substitutes to IOP bioprinting, which has been proposed for skin [243], [244], bone [244], [245], cartilage [246], [247], muscle [246], [248], and endoscopic [249], [250] treatments. Since IOP bioprinting is performed in surgical settings, certain concerns arise. First, bioprinting-related equipment and materials, including bioprinters, bioinks, and cell sources, should meet clinical standards in terms of sterility and clinical grade. Second, close cooperation among practitioners and engineers is warranted. Third, IOP bioprinting is still in its developing stage, although its application to musculoskeletal tissue regeneration has been demonstrated using small animal models. Hence, more investigations using large animal models and validations on solid organs are necessary to translate IOP bioprinting from the laboratory bench to the patient’s bedside.

6. Growing effort from bench to bedside

6.1. Preclinical studies and translational research

Prior to preclinical applications and translational research, it is essential to evaluate bioprinted organ constructs in order to estimate their similarity to native tissues in terms of anatomical structure, differentiation, and functionality. To observe microstructures, bright-field and electron microscopy (i.e., scanning electron microscopy (SEM) and transmission electron microscopy (TEM)) are widely used. In bright-field microscopy, the simplest and most common method, white light is reflected from the sample to capture images that are visible to the naked eye. It can also capture images of stained or unstained samples with adjustable magnification using object lenses. However, the maximum magnification is often limited to 100 times, which is insufficient to observe fine structures or cells at the microscale or nanoscale [251]. Electron microscopy has a higher resolution than light microscopy owing to the shorter wavelengths involved [252] and can reveal ultrastructural features less than 100 nm in size [253]. SEM images are acquired by transmitting high-energy electrons to the sample; with the interactions between the electron stream and the sample converted into a detailed raster image [254]. SEM resolution can potentially reach a magnification of two million, so SEM is often used for visualizing the surface of bioprinted organ constructs, such as heart tissue [255], liver [61], kidney [256], lung [257], and pancreas [258]. However, despite its advantages of high resolution and 3D imaging, SEM has the limitations of producing black-and-white images that do not reflect stained samples and of requiring sample-disruptive fixing steps [259]. In comparison, TEM uses a broad beam of electrons that transmit through samples to collect information such as crystal structure, composition, and morphology [260]. Its resolution can be up to five million times magnification, but TEM imaging is limited to ultrathin 2D samples [261].

To comprehend the maturity of bioprinted organ constructs, staining techniques are considered to be the gold standard for understanding the degree of tissue regeneration and differentiation qualitatively and quantitatively. In this regard, immunofluorescence, immunohistochemistry, and immunocytochemistry are widely used for analyzing the structure of in vitro and in vivo bioprinted organ constructs. More specifically, immunofluorescence is a concept that utilizes antibodies to localize proteins of interest; it can include immunohistochemistry and immunocytochemistry [262]. Within the scope of the concept, immunohistochemistry generally refers to the staining of samples that are derived from tissues and processed into thin sections [263]. In comparison, immunocytochemistry focuses more on single-cell staining, so ECM or other stromal components can be removed [263]. These histological analyses are useful for identifying the distribution and abundance of organ-specific markers, although they are limited to planer information. Therefore, to detect 3D structures, 3D imaging techniques are required.

Computed tomography (CT) is the most common 3D imaging technique, followed by MRI. CT images are generated by projecting and reconstructing multiple 2D X-ray images taken at slightly different angles. For reconstruction into a 3D image, the 2D pixels obtained from the samples are represented as small volume elements called voxels [264]. Among its various benefits, CT is non-destructive; thus, it is often chosen to observe both internal and external structures [265]. The limitations of CT include exposure to radiation and low capability for visualizing soft tissues [266]. MRI scanning is another noninvasive technique that can conduct 3D imaging of both hard and soft tissues [264]. With these advantages, MRI is used for imaging musculoskeletal, vascular, and neurologic structures of human anatomy and engineered biological substitutes [267]. MRI images obtained from human organs can provide an anatomic atlas, but the MRI procedure is relatively lengthy and requires subjects to remain motionless during scanning.

Studies on mimicking an organ’s entire structure and functionality have been attempted since the advent of bioprinting techniques. Along with the development of bioprinted organs, evaluation methods on functionality have been developed. For example, cardiac muscles are known to be involuntary muscles, with a contraction ranging between 60 and 120 beats per minute (bpm) for functioning [268], [269]. Hence, the contractility of bioprinted cardiac tissues is assessed in terms of amplitude and area of contraction, beating frequency, and synchronous contractile function with host tissues [270]. The hepatic functionality of bioprinted liver models can be evaluated according to albumin secretion or CYP3A4 activity [157]. In addition, the expression of liver-specific genes (e.g., ZO-1, MRP2, and CK18) or the liver-specific secretion of proteins (e.g., A1AT, ceruloplasmin, and transferrin) can be tracked [271]. Similarly, kidney functionality can be analyzed via proximal tubule functionality assays and tubule patterning- and function-related gene expressions (e.g., HNF4A, CUBN, LRP2, and SLC12A1) [272]. The lung is a dynamic organ that expands and contracts spontaneously in respiratory cycles. Hence, a study proposed a bioprinted lung model that applied both cyclic ventilation in the airway channel and perfusion in the vascular channel [273]. The perfused red blood cells then exhibited an oxygen saturation level (> 90%) close to the normal oxygen saturation level (95%-100%). In addition, immunofluorescence (e.g., aquaporin-5, caveolin-1, and prosurfactant protein C) or immunohistochemistry (e.g., Masson-Goldner trichrome) can be carried out to observe the layer formation of bioprinted lung models via histological cross-sections [173], [174]. For bioprinted pancreas models, the islet composition can be analyzed in terms of cell types and ratios. To elaborate, there are α (20%), β (70%), δ (< 10%), γ (< 5%), and ε (< 1%) cells in the pancreas, where this specific ratio is an optimal condition in the engineering and regulation of endocrine hormone-related α, γ, δ, and ε cells [274]. The microstructure of bioprinted pancreas models, which is as important as the hormonal function, can be confirmed via immunofluorescence (e.g., lectin) and immunohistochemistry (e.g., hematoxylin and eosin) staining [275].

Using these analysis methods, preclinical models for predicting clinical outcomes can be assessed. To begin with, various bioprinted heart models have been proposed with improvements in patient-specific bioinks, vascularization, and in vitro or in vivo applications in mice [108], [116], [121]. External interventions such as cellular or extracellular stimuli (e.g., electrical field stimulation and 3D contractile chambers) have also been applied to bioprinted heart models to mimic physiological conditions and accelerate tissue maturation [276], [277], [278]. Thus far, some successful results have been demonstrated, including anatomically accurate cardiac patches, vasculature-like native tissue, and in vivo functional tissues. However, the achievement of fully functional bioprinted hearts or the integration of bioprinted cardiac tissue with the host remains a challenge.

In comparison, the liver-on-a-chip is widely used to estimate tissue responses against various conditions such as virus infections (i.e., hepatotropic hepatitis B virus) and gradients of oxygen and hormones (i.e., insulin and glucagon) [279], [280], [281]. Studies[157], [282] using this platform have provided useful information on long-term primary liver cell culture, the mechanism of metabolism, and drug conjugation metabolic zonation. However, there are few reports of liver-on-a-chip platforms fabricated via bioprinting, and more in vivo studies should be performed to test the compatibility, degradation rate, and toxicity of bioprinted liver constructs for clinical trials [283], [284]. Likewise, the kidney-on-a-chip was developed to study drug screening or renal diseases; it includes diverse renal tissue chips such as the glomerulus- or tubule-on-a-chip and the perfusable chip [285], [286], [287], [288]. In vitro, results have shown that these chips can recapitulate the human renal filtration barrier and display a well-established vasculature with improved tubular and glomerular maturity. However, the kidney functionality and post-printing maturation of the kidney-on-a-chip still require improvement in order to reach a similar level to native kidney.

Various preclinical studies on bioprinted lung models have been conducted, especially after serial outbreaks of respiratory diseases such as severe acute respiratory syndromes (SARSs) or coronavirus disease 2019 (COVID-19) [289]. Although bioprinted lung models have not yet been infected with these viruses, if this concept is carried out, it will have various preclinical applications for the investigation of pathogenesis, post-COVID-19 sequelae, and drug development [289]. From a technical standpoint, since lung structures are composed of intricate alveoli (diameter: 100-200 mm) and alveolar sacs (diameter: 400 mm), bioprinting resolution should be improved to recapitulate the lung’s anatomical structure and functionality [290]. Bioprinted pancreas models are considered to hold great therapeutic potential for the treatment of T1D. In particular, it is highly feasible to use 3D bioprinting techniques to bioprint multiple cell types in desired spatial orientations and complex shapes. In the near future, a pancreas-on-a-chip will be developed and utilized for purposes ranging from reconstructing the pancreatic mass to investigating pancreatic dysfunction and treatments for T1D [291], [292]. Table 3 [35], [43], [108], [116], [121], [159], [161], [173], [174], [186], [187] summarizes several examples of 3D bioprinted tissues and organs for preclinical investigation.

6.2. Implications for the future of organ transplantation and personalized medicine

The overarching objective of bioprinting is to generate viable and functional tissue and organ constructs with the aim of restoring or substituting damaged or non-functional tissues and organs. As a result, the clinical translation of tissue-engineered medical products [293] and bioprinted tissues and organs is critical and challenging [294]. The regulatory requirements for the clinical translation of bioprinted organs differ across different countries. According to the requirements of the US Food and Drug Administration (FDA), the primary consideration revolves around the question of whether bioprinted tissues and organs are categorized as medical devices or as minimally modified tissues (e.g., decellularized tissue grafts). Given the substantial processing involved in creating tissues through the deposition of bioinks, it is unlikely that bioprinted tissues and organs would meet the criteria for being minimally manipulated [197]. Nevertheless, certain tissue-engineered products have gained approval through the 510(k) process, indicating that there is a precedent for some bioprinted tissues to follow suit [295], [296]. Based on a risk-based classification, the regulatory pathway for FDA approval differs: a 510(k) directive is essential for Class I and II devices, while a premarket approval application is required for Class III [297]. A 510(k) premarket submission to the FDA demonstrates that a new device is “substantially equivalent” to an existing one [136]. Within five years (2010-2015), approximately 80 medical devices produced through additive manufacturing received FDA 510(k) clearance [298]. These devices lacked biological components and were used for implants, with 53% being orthopedic, 7% cranial/neuro, and 6% dental, while 34% served as surgical guides. The approval process for living tissues remains complex, hindering their clearance so far. When biologics are involved, a mere 510(k) clearance often proves to be inadequate, necessitating premarket approval or a biologics license application.

Acellular bioprinted scaffolds are likely to fall under the classification of class II medical devices, necessitating adherence to performance standards and special checks including in vivo performance, nonclinical performance assessments, and assessments for sterility, shelf life, and biocompatibility. Cellularized bioprinted tissues will mostly be classified as Class III medical devices, demanding comprehensive premarket approval application. This designation is more likely if these tissues are deemed to be novel and distinct from those already available on the market [197]. For more intricate bioprinted solid tissues and organs, there is currently a lack of comparable existing products or approved alternatives. For solid tissues such as the heart, for example, the initial and pivotal step in this direction is the engineering of small organs that replicate physiological functions and act as models for evaluating drug toxicity and establishing patient-specific heart disease models [143], [299]. Given the rapid pace of research, it is reasonable to anticipate achieving this milestone within the next five years. Subsequently, the focus would shift to fabricating functional organs for testing in preclinical animal models, transitioning from small-scale proof-of-concept studies in animals to larger models that closely mimic human physiology. Ultimately, human clinical trials would be conducted. A crucial point in this journey is the phase three clinical trial for a bioprinted solid organ, which is a prerequisite for regulatory approval. However, due to the challenges involved, it is difficult to predict a precise timeline for this phase, which might be projected to span a couple of decades [197]. Fig. 4 outlines the path from the discovery and development of functional tissues through preclinical testing and clinical trials to patient-specific medical applications of bioprinted tissues and organs.

In 2007, Organovo Holdings, Inc., which is based in California, USA, pioneered an entry into the realm of bioprinting by introducing functional bioprinted blood vessels. In 2014, the NovoGen MMX Bioprinter was used to successfully bioprint liver tissue, which is commercially available as ExVive3D Human Liver Tissue for preclinical drug discovery testing. The printed liver tissue is functional for at least 42 days [300]. Since the launch of Organovo, many other bioprinting companies have entered the market [301], including several companies, start-ups, spinoffs, and 3D bioprinting businesses, aiding in the commercialization of this innovative technology and generating a predicted market value based on the early success and novel applications of bioprinted products [302]. As announced in early 2023, the FDA will no longer require animal testing prior to human trials in the drug development process. This demonstrates the enormous potential for in vitro drug development via bioprinting. The potential standardization of bioprinting for advanced preclinical drug testing hinges on further exploration of reproducibility, quality control, and automation. Industry reports indicate that the worldwide bioprinting market reached a value of 2.0 billion USD in 2022, and it is projected to exhibit a compound annual growth rate of 12.5% from 2023 to 2030, leading to an anticipated revenue of 5.3 billion USD [141].

Currently (as per the National Library of Medicine (NLM)^†), there is only one record in the clinical trial database using the keyword “3D bioprinted.” This observational study (the national clinical trial number (NCT): NCT04755907) validates the potential of 3D bioprinted tumor models for predicting the response to chemotherapy in colorectal cancer [303]. Table 4 provides a list of clinical studies relevant to 3D bioprinted tissues and organs. Concurrently, several companies are already making strides in advancing these technologies. As a major milestone, the first 3D bioprinted human tissue was recently (began in 2020) implanted in patients by 3DBio Therapeutics (USA), a clinical-stage biotechnology company. They implanted a patient-specific ear (AuriNovo^TM implant), 3D bioprinted from chondrocytes, as part of a phase 1/2A clinical trial (NCT04399239). Aimed at surgical reconstruction of the external ear in people born with microtia, it is designed to match the size and shape of the patient’s existing ear, making 3DBio (USA) the first company to administer such a treatment to a human subject [304]. Although the findings of this study are yet to be formally published, and the long-term shape retention and stability of the bioprinted ear remain to be seen, this accomplishment signifies a substantial leap forward in scientific progress. In another collaborative work, the biotech companies United Therapeutics (USA) and 3D Systems (USA) successfully created an acellular intricate 3D-printed human lung scaffold designed for gas exchange in animal models. This scaffold consists of 44 trillion components (voxels), encompassing 4000 km of pulmonary capillaries and 200 million alveoli. Moving ahead, the partners intend to infuse the scaffold with a patient’s personalized stem cells to establish viable and transplant-ready human lungs [305].

Furthermore, in the shift toward personalized healthcare, bioprinting offers the opportunity to develop therapeutics on a large scale and manage the disease through personalized treatments [306], [307]. It is a versatile fabrication technique with no one-size-fits-all approach, but it offers an on-demand approach that could potentially tackle the increasing scarcity of organs. Employed for in vitro organ modeling, its ability to replicate human physiology opens up new avenues, including customized implants and drug innovation, thereby enhancing the field of personalized medicine. Through the effective coupling of diagnosis and intervention, the conversion of patient-specific images into custom implants and prostheses, and the advancement of gene-based therapies and regenerative medicine, bioprinting is projected to enhance and revolutionize customized healthcare.

Overall, although fully functional bioprinted solid organs (e.g., heart, kidney, liver, lung, and pancreas) are still some years away, quick progress is being made. Given the current research trajectory, unless regulatory or bioprinting advancements alter this course, it is projected that bioprinted solid organs could attain clinical approval as a viable transplant option by 2042 [197]. This development holds promise for alleviating the substantial shortage of donor organs currently being experienced. While the field of bioprinting continues to evolve, considerable challenges remain. These include integrating different cell types, creating completely functional multiscale vasculature, and achieving tissue functionality like that of a native organ. Additional progress is needed to enable clinical translation, including advancements in large-scale cell production and differentiation, the establishment of bioreactor platforms to facilitate post-print tissue maturation, and the implementation of a well-defined regulatory framework tailored for organ-scale constructs. In addition, there is a need for collaborative effort from scientists, clinicians, and regulatory bodies to translate this advanced biofabrication technology from bench to bedside.

7. Conclusions

This review provided an in-depth look into the advancements in the field of bioprinting toward the fabrication of solid organs. This technology represents a significant step in the biomedical sciences, as it brings us closer to the possibility of generating replacement organs tailored to individual patients—including not only the external anatomy of organs but also the internal blood vessels and complex networks responsible for their function. This review also acknowledged the challenges that must be overcome before bioprinting can be fully accepted in clinical settings. Maintaining the safety and reliability of bioprinted organs is of utmost importance, demanding rigorous testing and regulatory validation to ensure their effectiveness and safety. Furthermore, the review delved into the ethical aspects of bioprinting, with a particular focus on cell sourcing and the impact of human biology modification. Balancing the transformative potential of bioprinting with ethical concerns presents a multifaceted challenge that requires thorough scrutiny. Overall, the review highlighted the potential of bioprinting to transform organ transplantation and personalized medicine, thereby mitigating organ shortages and providing tailored treatments. Despite the challenges, clinical translation holds promising prospects of expanding the horizons of healthcare to meet patients’ needs more effectively. A future of unique healthcare opportunities is a possibility through the exciting technology of bioprinting.

Acknowledgments

This review article was supported by the National Natural Science Foundation of China (82372403), the Shenzhen Science and Technology Program (ZDSYS20220606100606013), the Shenzhen Institute of Synthetic Biology Scientific Research Program (DWKF20190010 and JCHZ20200005), the Shenzhen Science and Technology Major Project (KJZD20230923114302006), the National Institute of Dental and Craniofacial Research Award (R01DE028614), the National Institute of Biomedical Imaging and Bioengineering Award (R01EB034566), the National Institute of Allergy and the Infectious Diseases Award (U19AI142733), and the 2236 CoCirculation2 of TUBITAK award (121C359).

Compliance with ethics guidelines

ITO has an equity stake in Biolife4D and is a member of the scientific advisory board for Biolife4D, and Healshape. Xiang Wang, Di Zhang, Yogendra Pratap Singh, Miji Yeo, Guotao Deng, Jiaqi Lai, Fei Chen, Ibrahim T. Ozbolat, and Yin Yu declare that they have no conflict of interest or financial conflicts to disclose.

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