Current strategies for bioartificial organ engineering

In order to generate functional replacements of native organs, two major endpoints need to be reached: the correct constituent cell types have to be generated, and a physiologic spatial and functional organization has to be established, both of which are highly unique to specific organs [55]. Several technology platforms emerging in recent years provide promising starting points and at least the theoretical translational potential of transplantable bioartificial organs. These platforms include whole-organ decellularization and regeneration based on the utilization of organ’s native extracellular scaffold reflecting its unique geometry [8], three-dimensional organ printing based on computer-aided construction of matrix scaffold and targeted positioning of desired cell types [18, 50], and organoid engineering highlighted by the self-organization of different cell types into specific structures (Fig. 1) [38, 83].

Fig. 1
figure 1

Current strategies for bioartificial organ engineering. a Whole-organ decellularization and regeneration, b three-dimensional organ printing, and c organoid engineering

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Whole-organ decellularization and recellularization

The native extracellular matrix (ECM) is produced and maintained by the resident cells of each tissue or organ, which reflects the unique cellular organization and constitutes the microenvironment for each cell type to survive and function [7]. Whole-organ decellularization by detergent perfusion through native blood vessels removes all the cellular components leaving the organ’s ECM scaffold more or less intact. Importantly, distinct compartments within decellularized organs are well-preserved, which enables the delivery of different cell types to their physiologic locations. Since the initial introduction in heart engineering [60], this approach has been applied to other organs, including lung [59, 63], liver [86], kidney [75], and intestine [85]. Proof-of-principle regeneration of bioartificial organs based on these scaffolds in rodent models has been achieved by seeding vascular endothelial cells and organ-specific cell types (cardiomyocytes for heart, corresponding epithelial cells for lung, liver, and kidney). Some of the resulting bioartificial organs were transplanted into animal models and mimicked functions of their native counterparts, such as gas exchange in lung, metabolic processing in liver, and urine production in kidney [59, 63, 86]. This regenerative approach has recently been upscaled to large animal and human scale [9, 11, 21, 56, 79, 93].

Most organ decellularization protocols are detergent-based, although enzymatic and physical methods have been evaluated [16]. Several detergents have been explored for the purpose of decellularization, including sodium dodecyl sulfate (SDS), sodium deoxycholate (SDC), and 3-[(3-cholamidopropyl)dimethylammonio]-1-propanesulfonate (CHAPS). Many protocols allow the generation of acellular scaffolds; consensus has not been reached regarding the optimal approach/reagent for decellularization to maximize the preservation of ECM and removal of cellular components for each respective organ [21, 62, 88]. Although selective cell delivery into different compartments within decellularized organ scaffold has been well-demonstrated (e.g., vascular vs. airway compartment in lung), the achievement of homogenous cell distribution within each compartment remains challenging, resulting in patchy tissue formation inside whole-organ scaffolds. Furthermore, targeted seeding of different locations within a compartment (e.g., proximal vs. distal airway in lung, proximal vs. distal tubule in kidney) has proven to be challenging. While there is evidence for cell-specific niches preserved within the matrix, as seen in glomerular epithelium in kidney regeneration [75], further strategies to enable site-specific delivery and engraftment have to be developed.

Bioprinting

Many tissue-engineering approaches are based on scaffolds, on which cells adhere, survive, and function to enable the formation of living tissue. Instead of utilizing native organ scaffolds by stripping off all cells, bioprinting generates three-dimensional (3D) scaffolds de novo through layer-by-layer additive robotic fabrication. The design files are generated by identifying key architectural structures of target organs and decomposing these exterior and interior features into series of parallel slices. These slices then serve as templates for additive fabrication [18, 51]. Indirect bioprinting involves generation of patterned scaffolds through 3D printing and seeding cells onto the printed scaffolds. For example, cartilage tissues were regenerated by delivering chondrocytes onto 3D-printed PEG/PBT scaffold, which was proven to be superior to that regenerated based on PEG/PBT scaffold manufactured by conventional leaching method [27]. Direct bioprinting of cells incorporates scaffold materials and cells at the same time and has been explored to engineer small-size cartilage and bone tissues [17, 20]. Cartilage implants printed using poly(ethylene glycol) dimethacrylate together with human chondrocytes integrated into surrounding native cartilage and showed evidence of cell-based extracellular matrix modification [17].

Compared to conventional scaffold manufacturing using molds, 3D scaffold printing through additive fabrication allows precise control of internal architecture with high reproducibility. This facilitates efficient screen and optimization of structural variables including scaffold porosity, pore size, and permeability, all of which are critical factors determining subsequent biological delivery and tissue regeneration [18, 27, 65]. In the case of indirect bioprinting for subsequent cell seeding, a larger range of fabrication technologies can be used, including contact-involving analog printing with a high spatial resolution of less than 1 μm [18]. A further advantage of this technology is the possibility of incorporation of bioactive molecules. Growth factors or cell adhesion molecules can be added during scaffold printing and delivered following predefined patterns or concentration gradients, to control the behavior of subsequently seeded cells [13, 35].

In direct cell printing, the entire manufacturing process needs to be biocompatible and free of cytotoxicity. Contactless inkjet printing is one of the commonly used methods for printing cells, which offers an attainable spatial resolution limit of 100 μm [18]. As we have discussed above, efficient and homogenous cell delivery into preexisting scaffolds remains challenging, especially for engineering solid tissues/organs such as bone and muscle. Direct cell printing potentially overcomes this cell delivery challenge by incorporating cells and scaffold materials at the same time. The grafts generated by 3D bioprinting currently remain small and simple in structure [18]. To print grafts with more complex architecture and increased size closer to those of native organs, improvements have to be made to increase the resolution and scale of direct cell printing. Moreover, a deeper understanding of the cellular and ECM organization within target organs will lead to design files that mimic the hierarchical organization of native organs more closely.

Organoid engineering

Organoid engineering is an emerging technology that enables the engineering of mini-organs, or “organoids,” from differentiating embryonic stem cells (ESCs) or induced pluripotent stem cells (iPSCs) in 3D culture conditions. Through this approach, organ development has been mimicked for systems such as intestine [69, 68], pituitary [78], thyroid [3], retina [19, 54], and most recently cerebrum [38]. These studies highlight the self-assembly capability of differentiating cells in 3D culture. Vascularized liver organoids that mimicked many aspects of native liver’s anatomy and function were generated by cultivating human iPSC-derived hepatocytes with human endothelial and mesenchymal stem cells in 3D culture [83]. This further emphasizes the importance of stromal cells in organ construction and demonstrates the inherent capability of different cell types to coordinate their assembly during organ formation.

Organoid engineering is generally believed to complement conventional 2D culture to further our understanding of organ development and to promote disease modeling for mechanism study and drug screen [30, 38, 68]. From a regenerative medicine point of view, these mini-organs reveal the striking self-assembly capability of dissociated cells forming organized structure when proper 3D culture conditions are provided. In order to generate transplantable bioengineered grafts from such organoids, several challenges need to be addressed. Scaling organoids from currently millimeters to a clinically relevant size has to be accomplished. Although organoids mimic many aspects of native organs, even for complex structures such as the brain [38], they are still relatively simple structures. In order to bring organoids to the complexity and hierarchy of mature organs, supporting scaffold or structural guidance may need to be introduced.

During development, many organs assemble in a modular fashion with interconnected functional units, such as human lung having about 500 million alveoli and human kidney having about 500 thousand glomeruli [33, 58]. Constructing the functional units or a cluster of these units through self-assembly may be accomplished through organoid engineering. These units can then serve as building blocks for subsequent organ engineering approaches. For example, mini-tissue spheroids with different cellular compositions and organizations have been used as building blocks for macrotissue formation during organ printing [51]. Knowledge acquired and protocols developed in organoid engineering will certainly benefit recellularization of decellularized organ scaffold and organ printing. Understanding the unique interaction between different cell types during self-assembly will help to develop strategies and culture conditions promoting organ regeneration.

Providing the right cells for organ engineering

With the tremendous progress that has been made in stem cell biology and especially the development of iPSC technology, generating “personalized” bioartificial organs from patient-derived cells has become a theoretical possibility. While primary adult-derived stem and progenitor cells and embryonic stem cells continue to be of scientific interest, we decided to focus the following discussion on topics specifically related to the use of iPSC-derived cells in organ engineering.

Cellular differentiation

One of the strategies currently used for directed differentiation of human ESCs and iPSCs is to recapitulate developmental programs using sequential administration of growth factor cocktails. This strategy has proven to be successful in generating various differentiated cell types of all the three germ layers. As an emerging trend, compound screens were performed and identified small molecules that can enhance or replace the action of growth factors during directed differentiation, resulting in improved iPSC differentiation towards endoderm [12], pancreatic progenitors [14], hepatocytes [71], endothelial cells [32], specific neuronal lineages [22], and many others. Growth factor-free protocols have been developed for cardiomyocyte differentiation with up to 98 % efficiency [43, 49]. The use of small molecules to replace growth factors during iPSC differentiation offers advantages including improved reproducibility, lowered risk of contamination, and reduced cost, which will eventually facilitate the generation of scalable sources of clinical-grade cells for tissue and organ engineering.

There are several concerns associated with the use of iPSC-derived cells for clinical applications. First, the conventional way of generating iPSCs involves virus-mediated transduction of reprogramming factors [81]. The safety concerns related to viral transduction were alleviated by the development of non-integrating and non-viral methods of expressing reprograming factors using synthetic mRNAs [90] and episomal vectors [94]. Most recently, mouse iPSCs were generated from somatic cells using only chemical compounds [28]. This technique could facilitate clinical application of iPSC technology if applied to generate human iPSCs. Second, the differentiation competency of different iPSC lines can be highly variable [36, 52], which indicates that newly generated iPSC lines need to be carefully characterized regarding their pluripotency and differentiation potential. Lastly, human iPSC lines were found to have a high level of mutational load, the majority of which were enriched in genes mutated or having causative effects in cancers [23]. This raises a concern that some of these mutations could increase disease risk, especially cancer risk, when iPSC-derived cells/tissues are used in clinical applications. Thus, standardized genetic evaluation of iPSC lines is required to ensure safety before clinical use [23].

Functional characterization

Different cell types and their spatial organization determine the identity of an organ and provide the basis for its function. Cells can be characterized both molecularly and functionally. Molecular characterization generally refers to the analysis of a combined expression of genes (mainly transcriptional factors or surface markers) and represents the initial characterization of cell lineages differentiated from iPSCs. Functional characterization on the other hand can be complicated. The functions of cells are determined not only by cell-autonomous programs (gene expression profiles) but also by their niche, which is composed of neighboring cell types, ECM, and physiological stimuli [66]. For example, the acquisition of polarization and ciliation and the establishment of proper barrier function in iPSC-derived airway epithelial cells require ex vivo air-liquid interface culture that mimics post-natal airway epithelial environment [91]. Although many of ESC/iPSC-derived cell types display molecular profiles similar to their in vivo counterparts, they are usually functionally immature [80]. Cardiomyocytes derived from ESCs exhibit fetal instead of adult ventricular characteristics [25, 34, 53]. It was recently shown that iPSCs derived from murine ventricular myocytes exhibited markedly higher propensity to further differentiate into cardiomyocytes with a ventricular phenotype [92]. This indicates that establishment of certain epigenetic signatures is required for further maturation of iPSC-derived cells. These signatures are currently not recapitulated during in vitro differentiation. It remains challenging to develop in vitro culture conditions that enable the crossing of epigenetic barriers to efficiently differentiate desired cell types from generic cell sources that are easy to obtain, such as skin fibroblasts. As the differentiation efficiency of various cell types improves, more attention is required to evaluate and promote the acquisition of mature functions.

Cellular heterogeneity

Cell types such as vascular endothelial cells and fibroblasts are important building blocks for almost all organs. Up to now, generic endothelial cells and fibroblasts have been used for the purpose of tissue engineering. These generic cells include mature cells isolated from primary tissues (e.g., human umbilical vein endothelial cells) [37, 83], adult-derived progenitors (e.g., circulating endothelial progenitors or mesenchymal stem cells) [6, 83], and those differentiated from ESCs or iPSCs [67, 89].

Recent studies indicate that endothelial cells contain highly heterogeneous and specialized subpopulations with morphological and functional differences among different organs (e.g., non-fenestrated endothelium in lung capillary vs. discontinuous fenestrated endothelium in liver capillary) [1, 2, 5]. Microvasculature also displays remarkable differences within organs such as liver and kidney (e.g., glomerular capillary vs. descending vasa recta endothelium) [1]. We need to consider these functional differences when choosing endothelial cells for regenerating higher organ functions such as filtration or gas exchange.

The molecular mechanism directing endothelial heterogeneity just starts to be understood. Microarray expression profiling has been performed comparing a large number of in vitro cultured endothelial cell lines and comparing microvascular endothelial cells freshly isolated from different organs. These studies reveal molecular signatures unique to the microvasculature of each organ [15, 57]. These molecular signatures seem to be acquired through the interaction between endothelial cells and their organ-specific niches, as generic ESC-derived endothelial cells underwent in vivo tissue-specific maturation after engrafted in regenerating tissues [57]. Similarly, blood-brain barrier endothelial cells were generated through sequential neural and endothelial differentiation of human iPSCs [44], which further indicates that the in vivo endothelial maturation can be mimicked by in vitro co-culture of endothelial cells with specific neighboring cell types. These recent advances open up the possibility of generating tissue-specific endothelial cells, which can be used to construct morphologically and functionally specified vascular networks to achieve closer modeling of organ function through bioengineering.

Similar heterogeneity is also present in stromal support cells such as fibroblasts. Fibroblasts with distinct marker expression are associated with specific tissues or neighboring cell types. Taking the lung as an example, fibroblasts expressing NG2 proteoglycan are mainly associated with lung capillaries and are usually referred as pericytes [64]. On the other hand, fibroblasts expressing platelet-derived growth factor receptor alpha (PDGFRA) but not NG2 were able to recapitulate the alveolar type 2 stem cell niche supporting stem cell growth and differentiation [10]. These studies imply that different fibroblast subpopulations may need to be recruited for optimal construction of lung capillaries and airway structures.

In vitro expansion

Regeneration of human-sized organs requires a strikingly large number of cells. For example, human lung is predicted to contain about 220 billion cells [76]. Although human iPSCs theoretically represent an unlimited source of cells, purification is generally required during or after iPSC differentiation to enrich the desired cell types. Purification is usually done by fluorescence-activated cell sorting (FACS) or magnetic-activated cell sorting (MACS). Both techniques are not suited for purification of the large number of mature cells necessary for human organ regeneration. Ideally, the purification step would be performed at an earlier differentiation stage, with the resulting lineage-committed cells to be expanded to the desired number. However, once certain developmental milestones are achieved, expansion of the resulting cell types during long-term in vitro culture without senescence becomes challenging, risking potential loss of identity or functional competency. In recent years, several attempts have been made to improve long-term expansion of ESC/iPSC-derived cell types by introducing assisting molecules, matrices, and feeder cells. Through a compound screen, TGF-β inhibitor was identified as a potent small molecule additive to promote the expansion ESC-derived endothelial cells after FACS purification [32]. ESC/iPSC-derived hepatoblast-like cells can be extensively expanded and functionally maintained by using selected ECM coating materials [82]. Furthermore, the proliferation and self-renewal of isolated pancreatic progenitors can be achieved when co-cultured with matching mesenchymal cells [74]. It is likely that these approaches or their combination will make it possible to expand desired iPSC-derived cell lineages at multiple differentiation stages, which can then be tested for optimal performance in organ engineering. One of the questions we are currently trying to answer is the identification of the ideal seeding stage, when the transition from two-dimensional expansion to three-dimensional organ culture can be made. Matching cell differentiation stage to environment will be key in order to achieve cell engraftment and formation of functional tissues [83, 87].

Organ assembly and maturation

Once scaffolds have been repopulated with regenerative cells, in vitro culture conditions of the resulting 3D multicellular constructs become key drivers of functional tissue formation and maturation. Proper in vitro culture conditions need to (1) support cell metabolism, (2) promote cell-cell and cell-scaffold interaction, and (3) promote the acquisition of desired tissue/organ function.

Vascularization

Proper metabolism of tissue implants with a volume more than 2 to 3 mm3 cannot be supported by passive diffusion and requires a preexistent perfusable vascular network [47, 72]. Therefore, regeneration of functional vasculature within tissue-engineered organs to support efficient oxygen and medium perfusion is a prerequisite for maintaining cell metabolism and survival during long-term in vitro culture. This has been one of the main focuses of organ engineering based on all the platforms we have discussed. Whole-organ decellularization by perfusing detergent through the main vessel supplying the organ preserves the acellular vascular bed with high integrity of underlying basement membranes [60, 86]. Vascular regeneration was therefore attempted by seeding the decellularized scaffold with endothelial cells through the main artery, resulting in re-endothelialization of the previous vascular track, and recapitulating the complexity of organ-specific vascular hierarchy [59, 60, 63, 75]. In 3D bioprinting, patterned vascular networks were generated by seeding endothelial cells around a three-dimensional filament network of carbohydrate glass. After removal of the template carbohydrate, the engineered vascular network enabled blood perfusion under high-pressure pulsatile flow, supporting the survival of other cell types seeded in the perivascular space [48]. In the case of organoid-based approaches, all the constituent cell types, including endothelial cells, are mixed together in ECM gels. In liver organoid engineering, vascular networks formed by endothelial self-assembly integrated well with the perivascular hepatocytes and mesenchymal cells, giving rise to functional tissues enabling blood perfusion and metabolite secretion [83].

Stimulating tissue formation

In an ideal experiment, selective cell delivery positions specific cells in close vicinity to each other and thereby mimics native tissue organization. Subsequent tissue formation involves massive cellular remodeling setting up proper connections between cells and between cells and scaffold. Taking vascular regeneration as an example, endothelial cells seeded into decellularized organ scaffold start as a cord of cells. Vessel formation involves vascular lumen formation and establishment of apicobasal polarity. This process somewhat mimics vasculogenesis in vivo [26, 77]. Several approaches are being developed to facilitate tissue formation in bioengineered grafts, including introducing assisting stromal cells and scaffold functionalization with growth factors.

Mesenchymal cells or fibroblasts are the most commonly used stromal cells for tissue engineering. Interestingly, the presence of fibroblasts or mesenchymal stem cells significantly improved long-term survival of endothelial cells in vascular grafts and promoted perfusable vessel formation [6, 37, 46, 67, 89]. Similarly, co-seeded fibroblasts facilitated skeletal muscle tissue formation from primary myoblasts by maintaining myotube viability, mediating biomaterial degradation, and promoting muscle self-assembly [41]. In liver organoid engineering, mesenchymal stem cells were essential for 3D tissue formation of co-seeded hepatocytes and endothelial cells [83]. Protocols are being established to isolate and expand clinical-grade mesenchymal stem cells from human tissues for the purpose of cell therapy [70]. Mesenchymal stem cells differentiated from iPSCs can also be an alternative source of stromal cells [42].

Growth factors constitute a rich source of extracellular signals directing cell specification and tissue formation during animal development. It is noted that a significant portion of growth factors are bonded to ECM, and this ECM interaction allows the control of growth factor presentation in a temporal and spatial fashion [39]. Therefore, improving organ scaffolds by immobilizing growth factors offers the potential to facilitate tissue formation. Soluble growth factors are usually internalized after binding to their receptors. Compared to soluble factors, scaffold-immobilized growth factors cannot be internalized, which leads to more sustained activation of receptors and downstream signaling pathways [31]. Growth factors can be immobilized through biopolymetric gels, such as hydrogel, in which growth factors conjugated to matrix moieties are incorporated into the gel during matrix polymerization [39, 73]. This approach can be readily incorporated into bioprinting and organoid engineering, because in both systems, biopolymetric gels are conventionally used as cell carriers and supporting scaffolds. Another approach is direct chemical cross-linking, which is used to immobilize growth factors onto preexisting synthetic or natural scaffolds. Various growth factors have been immobilized through direct chemical cross-linking to promote angiogenesis, bone repair, dermal wound healing, and stem cell differentiation [31, 45]. However, challenges remain when applying direct chemical cross-linking to functionalizing decellularized whole-organ scaffolds. The non-selectivity of chemical reactions usually induces massive cross-linking of ECM, leading to significant change in scaffold properties, especially mechanical properties. This will be detrimental to the engineering of organs such as heart and lung, in which the tissue elasticity is essential for organ function.

Biomimetic culture conditions

Once cells are differentiated, positioned, and integrated into the physiologic context of multicellular tissue, functional stimulation promotes organ maturation during the later stages of in vivo development. Therefore, incorporating biomimetic culture conditions mimicking native physiological stimuli should facilitate the maturation of bioartificial organs in vitro. Electrical and mechanical stimulations have been incorporated into the engineering of heart tissues on various scales, including monolayer cell sheet, 3D tissue, and whole heart regenerated from decellularized scaffolds [24, 60, 84]. These conditions improve the maintenance of differentiated cell phenotype and promote the assembly of synchronously contractile cardiac tissues [24, 84]. Organ regeneration based on decellularized whole-organ matrix makes it possible to apply more physiological stimulations during in vitro organ culture, due to the anatomical features with preserved mechanical strength. Various bioreactors have been developed for in vitro culture of bioengineered whole heart, lung, kidney, and liver [60, 59, 75, 86], enabling organ-specific stimulations and functional measurements. The lack of mechanical strength remains a challenge for organ engineering based on bioprinting and organoid, which limits the types and strength of mechanical stimulations that can be applied. Nonetheless, it was recently shown that biomimetic stimulations can be performed even on microfluidic devices that mimic the essential composition of an organ [29].

Moving into clinical application

Preclinical evaluation in animal models

The ultimate goal of organ engineering is to implant bioartificial constructs to augment or replace lost tissue and organ function [8, 25]. From a regulatory perspective, bioartificial organs are biologic devices and need to undergo vigorous characterization and safety examination as governed by the Food and Drug Administration (FDA) [4, 40]. In order to guarantee patient safety, it will be critical to test the in vivo performance of bioartificial organs in animal models of human disease before moving into clinical application.

Small tissue-engineered grafts can be implanted into the cranial window of immunodeficient rodent models, in which graft survival, blood perfusion, and metabolic functions can be studied. This approach has been applied to various tissue-engineered vascular grafts and liver organoids [37, 83]. Vascular grafts without in vitro culture became perfused with recipient’s blood within 2 weeks after implantation [89]. Liver organoids, after in vitro culture for 4 to 6 days, connected with host vasculatures within 48 h after implantation, allowing various metabolic functions to be measured in vivo [83]. Another approach is to implant tissue-engineered graft as a patch attaching to preexisting organs, which has been well-explored in heart tissue engineering. Heart tissue was engineered using neonatal rat heart cells and ECM gel reconstituted in circular mold, which improved systolic and diastolic function in infarcted rat heart after implantation as patch [95].

The preservation of almost all the surgically relevant structures in decellularized organs could allow for orthotopic transplantation after in vitro regeneration. Paired organs such as lung and kidney further make it possible to have one side of the organ to be replaced with the regenerated counterpart, which allows the recipient to survive while the bioartificial organ is to be tested in vivo. Single lungs regenerated from decellularized scaffold have been transplanted orthotopically by connecting pulmonary artery, pulmonary vein, and bronchus. The resulting transplanted bioartificial lungs were perfused with recipient’s circulation, ventilated through recipient’s airway, and participated in gas exchange in vivo for hours [59, 63]. Similarly, tissue-engineered kidneys were transplanted in an orthotopic position made available by experimental left nephrectomy and were able to produce urine [75]. Alternatively, ectopic transplantation can be performed, in which the regenerated organ is not positioned in its normal anatomical position. For example, liver grafts engineered from decellularized scaffold were transplanted by connecting to recipient’s renal vein and artery [86]. These studies highlighted the advantage of physiologic graft size and the anatomic features within the tissue-engineered grafts [75].

These implantation/transplantation models represent the preliminary proof of principle of bioartificial grafts generated in vitro providing measurable function in vivo. Most of these tissue-engineered constructs did not last for prolonged periods of time in vivo [59, 63, 86]. Optimizations on the following aspects are required to improve their long-term in vivo performance. First, graft vascularization needs to reach high endothelial coverage and proper barrier function to avoid blood clotting or hemorrhage. This is especially critical in models where engineered grafts will be connected to the host’s circulation by surgical anastomosis and receive blood perfusion immediately after transplantation [59, 63]. Second, the length and degree of in vitro maturation that is necessary before successful long-term transplantation has to be defined [59]. Lastly, immunosuppression protocols need to be developed to optimize the long-term performance of transplanted grafts. Athymic rodent models with compromised T cells have been used for transplantation of organs engineered using human cells [59]. However, it has been suggested that other components of the immune system, such as macrophages, may also be involved in the response against tissue-engineered organ scaffold after implantation [61].

Regulatory considerations

Bioartificial organs represent complex products with a combination of cells, scaffolds, and other factors. This complexity necessitates the establishment of procedures during product development, manufacturing, and characterization to ensure safety and effectiveness [40]. Recognition and identification of the regulatory pathways and understanding of the requirements at clinical phases will help to structure the early development of bioartificial organs as potential regenerative medicine products [4].

Cells are regulated by FDA using the tissue rules, which emphasize the prevention of infectious disease transmission. To comply with this requirement, procedures need to be established to control cell sources (such as donor eligibility), manufacturing facilities and protocols, and cell characterization (such as biological phenotype and sterility) [40]. Differentiated cell types derived from human iPSCs become prevalent for the purpose of regenerative medicine. As we have discussed in the previous sections, these cells need to be well-characterized regarding their identity, purity, and functional maturity before being incorporated into tissue-engineered products [4]. Any pluripotent cells need to be removed due to their teratogenicity. When multiple cell types are to be combined for organ regeneration, it needs to be a known mixture, which means that all the participating cell types need to be pure and well-characterized before being combined.

Scaffold can be derived from synthetic materials (such as polymeric materials), from extracted biomaterials (such as proteins and peptides), and from decellularization of human or animal tissues/organs. Similar to the aforementioned tissue rules regulating cell products, bioburden within the scaffolds needs to be strictly controlled to minimize infectious disease transmission. Moreover, the biocompatibility of scaffolds needs to be thoroughly investigated [40]. Complete removal of cytotoxic chemicals need to be demonstrated for materials derived from chemical synthesis or organ decellularization, because in both cases, reactive chemicals or detergents are heavily used during scaffold production. Moreover, scaffolds need to be fully characterized regarding their stability, physical, and mechanical properties.

After cells and scaffolds are being well-characterized, novel properties will emerge after their combination. On the one hand, cells may behave differently after seeding onto 3D scaffold, which is different from their regular 2D culture environment. On the other hand, cells are constantly modifying the scaffold they are attaching to through ECM degradation and production. Therefore, the final tissue-engineered products require further characterization in vitro and in vivo to ensure desired quality [40]. The other challenge for cell-scaffold composite products is the control of product consistency [40]. This necessitates the development of procedures that ensure consistent cell delivery, in vitro culture, and functional stimulations.

Conclusions and future directions

The engineering of bioartificial organs aims to generate functional replacement of native organs and offers the hope to alleviate donor organ shortage. Several strategies were successful in providing proof of principle that cells and scaffolds can be combined to form functional tissue. These strategies either mimic the organization of native organs by selective cell delivery (e.g., whole-organ decellularization and bioprinting) or create a microenvironment facilitating tissue formation by self-assembly of participating cells (e.g., organoid engineering). Optimization of culture conditions and application of biomimetic stimulations further allows tissue-engineered constructs to survive and mature in vitro, which can then be tested in vivo. In parallel, advances in stem cell biology continue to generate differentiated cell types with higher similarities to their native counterparts not only molecularly but also functionally.

Bioartificial organ engineering needs to emphasize on the development of grafts capable of providing physiological functions. Here, we propose four levels of functional tests from cells to organs during the engineering process, taking lung regeneration as an example. The first level of functional test occurs at the level of cells and scaffolds before their combination (e.g., surfactant secretion of type 2 alveolar epithelial cells and proper elasticity of lung scaffold suited for ventilation). The second level is the functional test of different tissue units within each organ (e.g., perfusion and barrier function of lung vasculature and ventilation of the airway). The third level is the in vitro test of combined functional output of multiple constituent units within the regenerated organ (e.g., in vitro gas exchange of the regenerated lung). The final level is in vivo functional test of the regenerated organ after transplantation (e.g., blood perfusion and in vivo gas exchange of a regenerated lung after orthotopic transplantation). The last level is similar to regular physiological examination of native organs. However, care should be taken to ensure the functional tests on the initial three levels are physiologically relevant and representative. Identification of this stepwise hierarchy of functional tests will facilitate the definition of clear endpoints directing optimizations on each step during technology development and bring us closer the ultimate goal of generating functional organ replacements for clinical use.