Innovations in 3D bioprinting and biomaterials for liver tissue engineering: Paving the way for tissue-engineered liver

Abstract

The liver is a pivotal organ that maintains internal homeostasis and actively participates in multiple physiological processes. Liver tissue engineering (LTE), by which in vitro biomimetic liver models are constructed, serves as a platform for disease research, drug screening, and cell replacement therapies. 3D bioprinting is used in tissue engineering to create microenvironments that closely mimic authentic tissues with carefully selected functional biomaterials. Ideal functional biomaterials exhibit characteristics such as high biocompatibility, mechanical strength, flexibility, processability, and tunable degradability. Biomaterials can be categorized into natural and synthetic biomaterials, each with its own advantages and limitations, and their combinations serve as a primary source of 3D bioprinting materials. It is noteworthy that the liver decellularized extracellular matrix (dECM), obtained by removing cellular components from tissues, possesses traits such as bioactivity, biocompatibility, and non-immunogenicity, making it a common choice among functional biomaterials. Furthermore, crosslinking of biomaterials significantly impacts the mechanical strength, physicochemical properties, and cellular behavior of the printed structures. This review covers the current utilization of biomaterials in LTE, focusing on natural and synthetic biomaterials as well as the selection and application of crosslinking methods. The aim is to enhance the fidelity of in vitro liver tissue models by providing a comprehensive coverage of functional biomaterials, thereby establishing a versatile platform for tissue-engineered livers.

1. Introduction

The liver is the largest internal organ in the human body. It plays a pivotal role in numerous physiological processes, such as bile secretion, glycogen storage, and macronutrient metabolism, including proteins, lipids, and carbohydrates. Additionally, it participates in detoxification, hematopoiesis, and coagulation processes, all of which are essential for the maintenance of vital life functions [1,2]. The liver extracellular matrix (ECM) is a complex insoluble composite consisting of collagen (fibrillar and network), adhesion proteins (e.g., laminin and fibronectin), and proteoglycans (PGs) (e.g., heparan sulfate PGs and chondroitin sulfate PGs). It plays a crucial role in various physiological functions and processes, including structural support, cell communication, and growth regulation [3,4]. The liver parenchyma is composed of 60%–70% of liver cells and 30%–40% of non-parenchymal cells (NPCs). Among the NPCs, liver sinusoidal endothelial cells (LSECs) form the inner layer of the hepatic sinusoid, serving as a structural scaffold for hepatocytes. The porous endothelial layer facilitates free exchange of substances between the bloodstream and hepatocytes. Additionally, LSECs secrete biologically active factors and matrix components essential for maintaining hepatocellular function. Hepatic stellate cells (HSCs) store vitamin A and by-products, along with triglycerides, while Kupffer cells (KCs) and natural killer cells (NKs) contribute to immune regulation [5,6].

Tissue engineering employs cells, functional biomaterials, and specific biochemical and physical factors to replicate the structural characteristics of target tissues [7]. One of the pivotal tools in tissue engineering is 3D bioprinting. In contrast to conventional fabrication methods, 3D bioprinting enables the precise delivery and deposition of bioinks, which facilitates cells to reorganize in a 3D environment mimicking the complexities of natural tissue structures [[8], [9], [10]]. In 3D bioprinting, cells are combined with biomaterials to form bioinks, and thus the biomaterial is a critical determinant of printing success. Biomaterials should ensure model stability post-printing. Additionally, biomaterials must create a biocompatible environment for cells, promoting adhesion, self-assembly, and growth within the biological context. Last, biomaterials should contribute to the construction of complex structures, such as lumens formed by sacrificial materials [11].

Functional biomaterials can be categorized into natural and synthetic materials [12,13]. Natural biomaterials are predominantly water soluble and have properties that are advantageous in LTE applications, including enhanced compatibility with cells and ease of manipulation and reshaping. However, they also exhibit various shortcomings, such as diminished mechanical properties, limited availability, and variable degradation rates, which has led to the emergence of synthetic biomaterials characterized by high mechanical strength, flexibility, processability, and tunable degradability. Nevertheless, synthetic biomaterials lack cell recognition and adhesion sites, resulting in comparatively lower biocompatibility compared with natural biomaterials [14]. Therefore, combining the two types of biomaterials represents one of the primary approaches to obtain suitable bioinks for 3D bioprinting. It is worth noting that the ECM of the liver consists of a complex network of macromolecules synthesized and secreted by various tissues and cells in the body. It is primarily composed of protein components, predominantly collagen, and glycoproteins such as fibronectin [15]. Natural biomaterials offer opportunities for the reconstitution of the liver ECM. Another common choice of biomaterials for constructing microenvironments for liver cells is the liver decellularized extracellular matrix (dECM), which is derived directly from animal sources. This biomaterial is favored owing to the cross-species tolerance of decellularized biological matrices, effectively reducing the risk of immune rejection [16,17].

Biomaterial solutions are typically transformed into 3D structures through crosslinking, which significantly affects the mechanical strength, physicochemical properties, and behavior of the printed structure within the biological context. Crosslinking methods can be classified into several categories according to the nature of the biomaterial backbone and functional groups, including chemical, physical, enzymatic reactions, and combinations thereof [18].

Advances in biomaterial research have laid a robust foundation for the development of precise, reproducible, and biologically supportive liver tissue engineering (LTE). Here, we introduce the characteristics and value of natural and synthetic biomaterials commonly used in LTE, review recent innovations in the field of biomaterials, provide examples of strategies used to enhance liver cell viability and functionality, and summarize the crosslinking methods employed in the 3D bioprinting of bioinks. A timely, comprehensive, and well-organized review facilitates tissue engineering scientists gain a holistic understanding of the latest research advancements in biomaterials. This expedites progress in LTE, promoting early translation and application. Our aim is to explore optimal combinations of biomaterials and crosslinking techniques for the construction of liver tissue models, thereby establishing a promising platform for liver disease exploration, drug screening, and cell replacement therapies (Fig. 1).

Fig. 1.

Biomaterials, cells, techniques, and applications of LTE. The figure of 3D printed hepatic lobule model was adapted with permission from Khati et al. (2022) [16], Creative Commons Attribution 4.0 International License; the figure of microfluidic assembly of hepatic lobule model was adapted with permission from Du et al. (2021) [19], Copyright © 2021, Acta Materialia Inc. Published by Elsevier Ltd; the figure of microfluidics was adapted with permission from Ya et al. (2021) [20], Copyright © 2021, American Chemical Society. PLA: polylactic acid; PLACL: poly (L-lactic acid)-co-poly (ε-caprolactone); PEG: polyethylene glycol; PVA: polyvinyl alcohol; LTE: liver tissue engineering.

2. Article sources and screening processes

To ensure an unbiased selection of reviewed studies and materials and to include the latest and most relevant research, we conducted a comprehensive search of the Web of Science™ database using a Topic Search approach. The search encompassed fields such as Title, Abstracts, Author keywords, and Keywords plus. The final selection for the review comprised the most recent, representative, and cutting-edge articles. Additionally, we performed a year-wise analysis of the publication and citation counts for these topics to provide an intuitive visualization of the field's emergence as a recent research hotspot (Fig. 2).

Fig. 2.

Article sources and screening processes. The Web of Science™ All Databases maintained by Clarivate Analytics was searched as source data. Bibliometric analysis was performed using the citexs website (https://www.citexs.com/).

3. Natural biomaterials

The liver ECM is primarily composed of major proteins, such as collagen, fibronectin, and laminin, along with minor proteins, including elastin, vitronectin, and lumican [3]. To faithfully recreate the authentic liver tissue microenvironment, the ideal biomaterial for 3D bioprinting should possess biochemical and biophysical properties similar to those of the native liver ECM [21]. The prevailing approach involves modifying or fusing multi-components, which is a significant direction in the design of bioinks for LTE. This section summarizes the latest advancements in strategies using natural biomaterials to construct tissue-engineered liver models.

3.1. Proteins

3.1.1. Collagen

Collagen is a critical component of the ECM and stands out as one of the preferred materials for LTE applications. Using collagen as a bioink offers several prominent advantages, including biocompatibility, biodegradability, and abundant bioactive sites to support cell adhesion, growth, and differentiation. Additionally, collagen can interact with elastin fibers to provide recoil for the ECM and fibronectin [22,23]. However, compared with biomaterials that are stabilized by crosslinking, pure collagen gels show weak mechanical properties with fast degradation, preventing them from directly completing model construction [23,24]. Consequently, in LTE, collagen is often used as an auxiliary factor to enhance biocompatibility with other materials, such as gelatin-alginate-collagen bioinks [25], alginate-collagen bioinks [26], and collagen-xanthan gum (XG) bioinks [27]. Furthermore, optimizing the properties of collagen is an important approach to improving the printability of the bioink. A uniformly soluble Type I concentrated collagen solution named Viscoll exhibits remarkable viscoelasticity and serves as a carrier of growth factors, enabling the high-fidelity 3D bioprinting of rigid scaffolds suitable for cell alignment and activity [28,29]. Furthermore, Guo et al. synthesized norbornene-functionalized neutral soluble collagen (NorCol) through the reaction of acid-soluble collagen (Col) with carbic anhydride in an aqueous environment. NorCol retains the triple helix conformation of collagen and exhibits improved compatibility with other biomaterials, such as gelatin. Importantly, it maintains structural integrity even after gelatin dissolution, indicating enhanced mechanical properties. As a bioink, NorCol exhibits flexible formability and excellent biocompatibility in 3D bioprinting [30,31] (Fig. 3A). Additionally, acid-soluble fibrous collagen FibercollFlexN® (ColN) bioink demonstrates high printability. Garcia-Villen et al. showed that cell viability is maintained without compromising printing performance by neutralizing the bioink with a biocompatible buffer (TRIS-HCl) [32]. Additionally, numerous researchers have combined methacrylated collagen with other biocompatible materials to craft bio-inks endowed with robust ultraviolet (UV) crosslinking potential, high printability, and outstanding biocompatibility [[33], [34], [35]].

Fig. 3.

Natural biomaterials based on proteins. (A) Fluorescence micrographs showing the viability (visible proliferation) of cells in NorCol and native Col hydrogels. Live cells: green, dead cells: red (Adapted with permission from Guo et al. (2021) [30], Copyright © 2021, American Chemical Society). (B) Each 3DP-HCC is a steric grid containing the designed number of cells. The viability of HCC cells. Live cells: green, dead cells: red. Scale bar = 100 μm (Adapted with permission from Xie et al. (2021) [44], Copyright © 2020 Published by Elsevier Ltd). (C) Printability of the hybrid bioink. (a) Schematic illustration of the structure. (b-i, c-i) Top and (b-ii, c-ii) side views of free-form printed structures with heights of 1.8 mm (b) and 3.4 mm (c). (d) Schematic illustration of embedded printing of honeycomb structures with heights of 3.75 mm (e) and 6.8 mm (f). Scale bars: 5 mm. (D) Adjacent NIH/3T3 (in 135ACG) and HepG2 (in GelMA) cell morphology. (a) Schematic illustration of the structure. Fluorescence images of both types of cells near the boundary of 135ACG and GelMA (dashed line) on (b–i) day 1, (c–i) day 7, and (d–i) day 14 (live cells: green, dead cells: red). Scale bar = 100 μm (Fig. 3C, D were adapted with permission from Wu et al. (2020) [43], Creative Commons Attribution 4.0 International License). (E) Scanning electron microscopy (SEM) images of micropatterned scaffolds and fluorescence images of live/dead stained cells (Adapted with permission from Xiao et al. (2011) [59], Copyright © 2011, Acta Materialia Inc. Published by Elsevier Ltd).

3.1.2. Fibrin

Fibrin, a soluble protein produced and secreted by the liver, is widely employed in LTE [36]. Similar to collagen, fibrin shows high biocompatibility and provides an appropriate scaffold for cell adhesion, proliferation, and differentiation. However, it exhibits low mechanical properties and rapid enzymatic degradation [37]. Currently, fibrin is often used as an additive to enhance biocompatibility [38]. Stevens et al. formulated bioinks composed of fibrin along with patterned human primary hepatocytes, endothelial cells, and stromal cells, enabling 3D bioprinting of microscale liver tissues that rapidly grow upon implantation into mice with liver injury [39]. Additionally, fibrin is frequently employed in dual-crosslinking strategies to facilitate the printing of layered cell structures with increased hardness and stability. For instance, dual gelation of a bioink containing alginate and fibrin components in the presence of Ca²⁺ and thrombin has been used to create layered cell-filled 3D tissue engineering models [40].

3.1.3. Gelatin, GelMA

Gelatin, a water-soluble protein derived from the partial hydrolysis of collagen, is extensively used in LTE. It is a biocompatible material capable of forming a reversible molecular network resembling collagen through physical gelation. Compared with collagen, gelatin exhibits low cytotoxicity and minimal immunogenicity, making it a more biologically advantageous choice [41]. However, gelatin suffers from drawbacks such as low mechanical strength, high degradation rate, and poor thermal stability. In addition, gelatin liquefies at cell culture temperatures (approximately 37°C), rendering it challenging for direct 3D bioprinting [42,43]. Consequently, gelatin is often crosslinked with other biomaterials to improve its printability. Examples include gelatin-alginate bioinks [44] (Fig. 3B), gelatin-alginate-matrigel bioinks [45], and alginate-gelatin-ECM bioinks [46,47]. Furthermore, gelatin can be modified into gelatin methacrylate (GelMA) for UV crosslinking, offering a degree of biological functionality and tunability of mechanical properties. GelMA has been applied to construct structurally stable 3D biological scaffolds with customized degradation characteristics [48]. Currently, GelMA is the predominant biomaterial employed for the 3D bioprinting of liver tissues [43,49,50]. For instance, Ma et al. used GelMA-based bioink comprising human induced pluripotent stem cell (hiPSC)-derived hepatic cells and supporting cells to 3D bioprint honeycomb-like liver lobule structures. Compared with traditional 2D monolayer cultures, this 3D cell culture model exhibits enhanced activity and functionality of hiPSC hematopoietic progenitor cells (HPCs) [51]. Furthermore, Yun Wu et al. developed a hybrid bioink composed of GelMA (5%), alginate (1%), and cellulose nanocrystal (3%) (135ACG hybrid ink) exhibiting excellent shear-thinning properties and solid characteristics, ensuring high printability without compromising cell integrity [43] (Fig. 3C, D).

3.1.4. Elastin

Elastin is a primary constituent of elastic fibers and is predominantly found in ligaments and arterial walls. Studies have revealed that the coexistence of elastin and fibrin imparts tissues with elasticity and tensile strength [52]. Moreover, disruption of the synthesis and metabolism of elastin may play a crucial role in liver fibrosis [53]. Lee et al. developed a highly biocompatible bioink composed of GelMA and methacryloyl-substituted recombinant human tropoelastin (MeTro). Importantly, the latter can be crosslinked under non-cytotoxic visible light. This bioink is suitable for 3D bioprinting of complex soft-tissues [52] and holds promise for engineering highly vascularized tissue-engineered livers. Additionally, Dai et al. created a bioink based on elastin-like polypeptide (ELP), which can be chemically modified for photo-crosslinking and also functionalized with peptide Gly-Arg-Gly-Asp-Ser (GRGDS) or combined with collagen to enhance cell adhesion capabilities, showcasing significant potential applications in LTE [54].

3.1.5. Laminin

Laminin, a structural protein, is primarily synthesized by endothelial cells and adipocytes in the liver. Along with collagen, laminin is a component of the basement membrane. Its biological function in the liver involves mediating cell adhesion to the matrix and regulating tissue development, cell growth, and differentiation. Owing to its limited printability, laminin has found relatively few applications in LTE and is primarily employed in the regulation of cell maintenance and differentiation [55,56].

3.1.6. Silk fibroin

Silk fibroin holds immense potential in tissue engineering owing to its biocompatibility and excellent mechanical properties, such as flexibility and tensile strength [57]. Numerous bioinks based on silk fibroin have been developed. For instance, Sharma et al. used silk fibroin-gelatin-dECM to formulate a bioink for LTE, resulting in enhanced expression of liver-specific genes in cultured liver cells [58]. Furthermore, Xiao et al. devised a photo-crosslinkable interpenetrating polymer network (IPN) hydrogel composed of silk fibroin and GelMA, which exhibits low swelling, high compressibility, and slow degradation, making it suitable for various tissue engineering and regenerative medicine applications [59] (Fig. 3E). Kasoju et al. fabricated regenerated silk fibroin-galactosylated chitosan (GalCS) scaffolds using electrospinning. These scaffolds emulate the structural and biochemical characteristics of the native ECM and provide excellent support for liver cells, which makes them suitable for LTE [60].

3.2. Polysaccharides

3.2.1. Chitosan

Chitosan, derived from chitin through N-deacetylation, possesses numerous properties that are suitable for LTE, such as flexibility, non-toxicity, biodegradability, and cell affinity. In addition, the amino groups of chitosan are more reactive than the acetyl amino groups of chitin, which makes chitosan amenable to chemical modifications, by which it gains distinct biological functionalities. However, owing to its structural characteristics, chitosan suffers from the problems of poor stability, poor mechanical properties, and uncontrollable pore size [61]. Nevertheless, chitosan is frequently employed in LTE owing to its processability, often undergoing modification with functional molecules [62] or integration with other biomaterials [63,64], such as alginate-chitosan [65] or chitosan-collagen [66,67]. These approaches enhance the printability and biocompatibility of bioinks. For instance, modifying chitosan with surface lactose ligands and electrospinning chitosan into nanofiber scaffolds not only improve the mechanical stability of the printed structure but also enhance the in vitro viability of primary liver cells [62]. More recently, Zhang et al. developed cell-loaded nanofiber hydrogels by combining 2,2,6,6-tetramethylpiperidine 1-oxyl-oxidized cellulose nanofiber (TOCNF) and chitosan nanofiber (CsNF). The carboxylate groups of TOCNF and the amine groups of CsNF are directly crosslinked through the chemical reaction between 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS), which significantly enhances the printability and biocompatibility of the hydrogel [64] (Fig. 4A).

Fig. 4.

Natural biomaterials based on polysaccharides and dECM. (A) SEM images of as-prepared pre-gels and hydrogels. (a) T₂₀C₁: TOCNF: CsNF = 20:1. (b) HepG2 cells cultured on the surfaces of hydrogels after 3 d. Live cells: green, dead cells: red. Scale bars = 200 μm (Adapted with permission from Zhang et al. (2023) [64], Creative Commons Attribution 4.0 International License). (B) HepG2 encapsulated in the shell compartment consisting of algMC + matrigel with the core remaining cell free (Adapted with permission from Taymour et al. (2021) [72], Creative Commons Attribution 4.0 International License). (C) (a) Grayscale digital masks corresponding to polymerizing lobule structure. (b) Brightfield images and fluorescently labeled hiPSC-HPCs (green) in GelMA and supporting cells (red) in GelMA with GMHA. Scale bars = 500 μm (Adapted with permission from Ma et al. (2016) [51], Creative Commons Attribution 4.0 International License). (D) SEM images of (a) AC (agarose (2 wt%), chitosan (2 wt%), GO (0 wt%)) and ACGO (agarose (2 wt%), chitosan (2 wt%), GO (1.5 wt%)) composite scaffolds. Fluorescence images of (b) FDA-stained cells after 48 h of incubation with AC and ACGO composite scaffolds (Adapted with permission from Sivashankari et al. (2020) [79], Copyright © 2019 Elsevier B.V.). (E) (a) Liver lobular-like structure fabricated with Hep3Gel. (b) Fluorescently labeled HepG2 cells. Live cells: green, dead cells: red. Scale bar = 5.0 mm (Adapted with permission from Guagliano et al. (2023) [83], Creative Commons Attribution 4.0 International License).

3.2.2. Alginate

Alginate is a negatively charged polysaccharide extracted from brown seaweeds, such as kelp or bladderwrack. Alginate is a natural high molecular weight biomaterial that is non-toxic, hydrophilic, and crosslinkable with multivalent cations, such as Ca²⁺ [68]. However, alginate faces several challenges hindering its applications in LTE. At room temperature, it exhibits low mechanical strength, inferior biocompatibility compared with natural protein-based biomaterials, inherent biological inertness with limited binding domains for cell interactions, poor barrier properties, and low adhesive properties [68,69]. These limitations have been addressed by altering its properties and co-applying it with other biomaterials [70,71]. For instance, Zhang et al. created galactosylated sodium alginate (G-SA) by activating the carboxyl groups with 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC)/N-Hydroxysuccinimide (NHS) in morpholinoethanesulfonic acid (MES) buffer, followed by Ca²⁺ crosslinking to improve the galactose grafting rate. These processes enhance the adhesion of liver cells on the alginate substrate to ensure high vitality of liver cells and excellent printability within the gel [70]. Additionally, Taymour et al. designed a highly printable bioink using alginate and methylcellulose (algMC). The addition of matrigel to the bioink enhances the vitality of liver cells [72] (Fig. 4B). It is noteworthy that metal cations play a reversible role in the chemical crosslinking of alginate. Prolonged cultivation of engineered liver tissues in a metal cation-free environment can lead to the gradual collapse of the tissue structure [11].

3.2.3. Hyaluronic acid (HA)

HA, an anionic, non-sulfated glycosaminoglycan, is a major component of the liver sinusoidal space. HA has significant application value in LTE owing to its good biocompatibility, non-immunogenicity, and ability to support endothelial cell proliferation and vascular formation [51]. Methacrylated HA (MA-HA) has been developed to enhance the chemical crosslinking capability and mechanical properties of HA for 3D printing. MA-HA is formed by substituting the hydroxyl groups of HA with methacrylate functional groups. It exhibits excellent biocompatibility and tunable crosslinking properties, that is, the crosslinking degree and performance can be controlled by altering the quantity and distribution of methacrylate functional groups [73]. Furthermore, its ability to construct a matrix microenvironment is significantly enhanced when co-applied with other biomaterials. For instance, Ma et al. combined glycidal methacrylate-HA (GMHA) and GelMA to provide endothelial cells with a more biologically active microenvironment, resulting in the development of a functional biomimetic liver model [51] (Fig. 4C).

3.2.4. Agarose

Agarose, a highly hydrophilic polymer derived from natural red algal polysaccharides, has a wide range of applications in biomedicine and bioengineering owing to its unique properties, such as hydrophilicity, porosity, long-term stability, slow degradation, good shape fidelity, and non-toxicity [74,75]. At high temperatures, the polymer chains of agarose are randomly coiled in the liquid state, and on cooling to physiological temperatures, the polymer chains form double helical aggregates to form a semi-solid-like gel, which is the basis for its numerous applications. In LTE, agarose is often used to prepare high-throughput micro-well scaffolds for the assembly and culturing of microtissues, such as spheroids [76,77]. Agarose is an uncommon material choice for 3D bioprinting owing to the viscosity clogging problem. In addition, the model structure may be affected by brittleness issues even if the model is printed at controlled temperatures. Campos et al. performed an in vivo long-term biosafety and biocompatibility analysis of fibronectin-agarose tissue-like hydrogels, revealing that cellular invasion, ECM component synthesis, and neovascularization continue after the scaffolds are implanted. In addition, implanted biomaterials complete biodegradation and biointegration by week 12, and there are no associated histopathological changes in the implanted region or distal vital organs [78]. Sivashankari et al. pointed out that the main drawback of agarose scaffold materials is the lack of cell recognition motifs, which can be solved by blending with other natural polymers, such as chitosan, collagen, and gelatin. Therefore, they prepared 3D porous scaffolds based on agarose/chitosan/graphene oxide (ACGO) composites, adding graphene oxide to impart the scaffolds with high mechanical properties and stability. The composite scaffolds exhibit excellent hemocompatibility and support cell proliferation [79] (Fig. 4D).

3.3. dECM

The liver dECM is obtained by removing the cellular components from the liver tissue using enzymatic, physical, or chemical methods. Owing to the high preservation of the specific components and geometric structure of the liver tissue, it offers significant advantages over single- or multi-component bioinks that are extracted or synthesized for the purpose of tissue ECM reconstruction [49,80,81]. However, the liver dECM faces several challenges in 3D bioprinting, such as potential immunogenicity, low viscosity, poor shape fidelity, and rapid degradation [11,82]. Nevertheless, multi-component bioinks based on the liver dECM hold promise for extrusion-based 3D bioprinting. To achieve this objective, Kim et al. developed a dECM-gelatin bioink (dECM gBioink), which offers higher bioprinting resolution and more stable stacking layers than traditional single-component dECM bioinks while maintaining excellent compatibility with liver and endothelial cells. The results demonstrate that gelatin blends can significantly enhance the printability and mechanical properties of the liver dECM while retaining cell compatibility [46]. Additionally, Guagliano et al. developed an alginate-dECM hydrogel (Hep3Gel) and finely tuned its crosslinking kinetics, enabling it to self-assemble and adapt to various culture media for the in vitro fabrication of physiologically relevant liver models. Furthermore, they created an internally cross-linkable bioink based on Hep3Gel, with a focus on modulating the oxygenation of the cell culture microenvironment, highlighting its suitability for long-term liver cell cultivation [83,84] (Fig. 4E).

Matrigel is akin to the ECM and is suitable for bioprinting. Comprising type IV collagen, laminin, heparan sulfate proteoglycan, and growth factors [85,86], matrigel has demonstrated significant utility in various applications. For instance, Schmidt et al. employed an alginate-gelatin-matrigel-based hydrogel to construct a model for assessing long-term aflatoxin B1 toxicity in vitro. Matrigel significantly enhances the functionality of hepatocytes, and 3D-bioprinted HepaRG cultures are more resistant to aflatoxin B1 treatment than 2D cultures [87]. This underscores the substantial potential of matrigel in the context of long-term experimentation.

4. Synthetic biomaterial

Synthetic polymer networks consist of repeatable inert units. Polymers with high molecular weights accommodate strong intermolecular interactions between polymer chains. Therefore, they are superior to natural biomaterials in mechanical properties, processability, tunable degradability, and non-immunogenicity. However, the main challenge with synthetic materials is their low biocompatibility. To ameliorate this situation, synthetic biomaterials are mainly produced by functionalizing polysaccharides, proteins or polypeptides [24]. There are also other issues, such as the generation of CO₂ from aliphatic polyester biomaterials after degradation, which alters the cellular microenvironment and limits their applicability in LTE [88]. This section discusses the latest developments of synthetic biomaterials for LTE.

4.1. Aliphatic polyesters

4.1.1. Polylactic acid (PLA)

PLA is an aliphatic polyester polymer primarily derived from lactic acid. It is a novel biodegradable material exhibiting excellent biocompatibility, mechanical strength, transparency, and heat resistance, with the disadvantages of releasing acidic products as well as structural fragility [89]. In LTE, PLA-based nanofiber scaffolds have been employed to provide a conducive microenvironment for in vitro regeneration of primary rat liver cells [90]. Furthermore, Liu et al. modified PLA by incorporating thermoplastic polyurethane (TPU) and lecithin (Lec) into electrospinning solutions. The resulting PLA TPU-Lec fibers demonstrate favorable morphology, hydrophilicity, and biocompatibility, enabling the establishment of scaffolds suitable for liver cell cultivation in bioreactors [91]. Pimentel et al. used PLA and water-soluble polyvinyl alcohol (PVA) to fabricate and perfuse thick, dense tissue-engineered liver with controlled mechanical properties and a 3D vascular network. This approach provides complex 3D microenvironments for simulations of cell adhesion, matrix elasticity, and molecular transport [92]. Additionally, in a recent development, Lee et al. used carbon black-polylactic acid (CB-PLA) to print an electrode with high stability as well as high biocompatibility and encapsulated the electrode with Nafion/multi-walled carbon nanotubes (MWCNT)/glucose oxidase (Gox) to provide good electrical conductivity. Ultimately, the highly sensitive environmental glucose sensing device was obtained without affecting the activity of hepatocytes. Undoubtedly, this is a key application of PLA and an important advancement in LTE [93] (Fig. 5A).

Fig. 5.

Synthetic biomaterials based on aliphatic polyesters. (A) (a) SEM images of bare CB-PLA and Nafion/MWCNT/GOx-coated CB-PLA. (b) Viability test of HepG2 cell exposed to bare CB-PLA and Nafion/MWCNT/GOx-coated CB-PLA. Live cells: green, dead cells: red. (c) Schematic diagram of integration of glucose biosensor and liver-on-a-chip (Adapted with permission from Lee et al. (2023) [93], Copyright © 2023, Wiley-VCH GmbH). (B) SEM (a) and fluorescence (b) images of hepatocytes cultured on TCP, meshRFs, and disAFs. Hepatocyte membrane: yellow arrows, disAFs: red arrows, live cells: green, dead cells: red (Adapted with permission from Feng et al. (2010) [97], Copyright © 2010, Elsevier Ltd). (C) (a) SEM images of fabricated PLGA membranes with moisture contents of 3%, 4%, 5%, 7%, 8%, and 10%. (b) Expression and localization of cadherin in 3D stacked structures. Pan-cadherin: green; nuclei: red. Note that cadherin is localized at cell–cell borders between vertically adjacent cells as well as horizontally adjacent cells (arrowheads). Scale bar = 20 μm (Adapted with permission from Kasuya et al. (2012) [106], Copyright © 2011, Elsevier Ltd). (D) (a) SEM images of scaffold fibers. (b) Images showing nuclei (DAPI) and f-Actin filaments (Phalloidin) of HepG2 cells on scaffolds. DAPI: blue, phalloidin: green (Adapted with permission from Bate et al. (2022) [114], Creative Commons Attribution 4.0 International License). (E) SEM images of electrospun nanofibers of PLACL/collagen blend (a) and hMSC-derived hepatocyte-like cells or hepatospheres on PLACL (b) and PLACL/collagen (c) nanofibrous scaffolds. (d) CMFDA fluorescent dye expression in trans-differentiating hMSCs on PLACL/collagen nanofibrous scaffolds. Scale bar = 200 μm. (Adapted with permission from Bishi et al. (2013) [120], Creative Commons Attribution-NonCommercial 3.0 Unported License). WE: working electrode; RE: reference electrode; CE: counter electrode.

4.1.2. Poly (L-lactic acid) (PLLA)

PLA exists as two optical isomers: poly (D-lactic acid) (PDLA) and PLLA. PDLA is costly and less readily available, and thus PLLA is the more commonly used isomer. PLLA is non-toxic, non-irritating, mechanically strong, biocompatible, biodegradable, and thermally stable. In addition, oxygen and water vapor can permeate through PLLA [94,95]. Numerous studies have found that PLLA scaffolds, including collagen-coated PLLA nanoporous fibers, promote the growth and functionality of liver cells [[96], [97], [98]]. For example, Feng et al. used collagen-coated PLLA discrete aligned nanofibers (disAF) to induce hepatocyte aggregation, which displays excellent cell adhesion, activity, and function compared with monolayer cultures of hepatocytes on tissue culture plastic (TCP) with collagen as well as on meshes of collagen-coated PLLA random nanofibers (meshRFs) [97] (Fig. 5B).

4.1.3. Poly (lactic-co-glycolic acid) (PLGA)

PLGA, a copolymer derived from the random polymerization of lactic acid and glycolic acid, is a biodegradable organic compound known for its excellent biocompatibility, non-toxicity, and ability to form capsules and films [99]. Numerous studies have employed PLGA as porous tissue-engineered scaffolds, which have been shown to provide optimal support for liver cell adhesion and activity [[100], [101], [102], [103]]. Furthermore, researchers have enhanced liver cell functionality and vitality by modifying PLGA scaffolds with collagen [104,105]. Additionally, Kasuya et al. successfully recreated the hepatic microvascular structure with incorporated HSCs in vitro using highly porous PLGA thin membranes. The membrane pore size and porosity can be controlled by varying the moisture content, leveraging PLGA's film-forming properties to lay the foundation for the in vitro construction of functional, thick, and vascularized liver tissues [106,107] (Fig. 5C).

4.1.4. Polyglycolic acid (PGA)

Within the aliphatic polyester family, PGA has the smallest carbon unit count and exhibits complete degradation with the fastest degradation rate. It boasts excellent biodegradability, heat resistance, biocompatibility, gas barrier properties, and mechanical strength, with the disadvantages of releasing acidic products as well as structural fragility [108,109]. There have been multiple studies applying PGA as biliary tissue engineering scaffolds [110,111]. In LTE, its utilization remains relatively limited.

4.1.5. Poly (ε-caprolactone) (PCL)

PCL is a high molecular weight polymer synthesized through the ring-opening polymerization of ε-caprolactone monomers with the aid of metal cation complex catalysts. By controlling the polymerization conditions, PCL with various molecular weights can be achieved. PCL possesses excellent biocompatibility and biodegradability [112]. Research has demonstrated that PCL scaffolds provide robust support for liver cells [113]. Furthermore, several studies have leveraged dECM to modulate the bioactivity of PCL scaffolds, significantly enhancing hepatocyte proliferation while maintaining the expression of critical liver genes [114,115] (Fig. 5D). Ghahremanzadeh et al. introduced modified PCL-galactosylated chitosan nanofibers for LTE. Galactose in the scaffold significantly enhances cell growth and proliferation without exerting toxicity [116]. Zhang et al. developed covalently galactosylated PCL (GPCL) through the coupling of amino-functionalized PCL (NPCL) with lactobionic acid (LA). Cells on the GPCL membrane exhibit improved adhesion and activity [117]. Additionally, Salerno et al. used PCL to develop a biodegradable hollow fiber membrane-based 3D vascularized liver tissue. This membrane segregates human liver cells and endothelial cells between the outer surface and fiber lumen, generating human vascularized liver units by perfusing 3D liver units, which maintain high activity for an extended period [118].

4.1.6. Poly (L-lactic acid)-co-poly (ε-caprolactone) (PLACL)

PLACL, a composite of PLLA and PCL, is a biodegradable and non-toxic biomaterial widely employed for tissue engineering [119,120]. Bishi et al. discovered that human mesenchymal stem cells (hMSCs) cultured on porous PLACL-collagen nanofiber scaffolds exhibit excellent proliferation and hepatic differentiation. Moreover, hMSC-derived hepatocyte-like cells can aggregate to form highly functional hepatospheres [120] (Fig. 5E).

4.2. Polyethers-polyethylene glycol (PEG)

Known for its excellent biocompatibility, hydrophilicity, non-toxicity, and non-immunogenicity, PEG is often incorporated as a hydrophilic component into poly-lactic acid to form amphiphilic block copolymers [121]. Zhang et al. developed PEG and vascular endothelial growth factor (VEGF) crosslinked decellularized single liver lobe scaffolds (DLSs). These modified DLSs not only improve the mechanical properties of natural DLSs but also demonstrate non-toxicity, promoting cell proliferation in vitro and vascularization in vivo for LTE [122] (Fig. 6A). Khati et al. developed a multi-component dECM bioink enhanced with gelatin and PEG to improve the printability and mechanical stability of dECM. This shear-thinning bioink, enzymatically crosslinked with tyrosinase, allowed for the fabrication of 3D grid structures of liver cells at 37°C, enhancing its long-term applicability [123]. Furthermore, Stevens et al. designed a hydrogel system using PEG-diacrylamide (PEGDAAm) containing matrix metalloproteinase-sensitive (MMP-sensitive) peptides. This system effectively supports the activity and functionality of liver cells both in vitro and in vivo [124].

Fig. 6.

Other synthetic biomaterials. (A) (a) HE staining showing the absence of cellular components and nuclear material in decellularized scaffolds. The porosity is high at high concentrations of PEG crosslinked scaffolds. Scale bars = 50 μm. SEM images of the native liver, DLS, and 0.5-PEG-DLS. Scale bars = 200 μm. (b) Immunofluorescence of HUVECs co-cultured with DLS scaffolds. Scale bars = 50 μm (Adapted with permission from Zhang et al. (2023) [122], Copyright © 2022 Elsevier B.V.). (B) PVA dissolves to form a hollow tubular structure with dLM-PEG-T. Scale bar = 1 mm (Adapted with permission from Khati et al. (2022) [16], Creative Commons Attribution 4.0 International License). dLM-PEG-T: dLM-PEG-tyrosinase.

4.3. Polyamides—(polyvinyl alcohol) PVA

PVA, an important water-soluble polymer, is extensively used for LTE [125,126]. However, there is a need to address its limited bioactivity owing to its lack of functional groups. Bidault et al. combined fibroin hydrogels with PVA within an IPN structure. The resulting material is non-toxic to human fibroblasts and also promotes cell development [127]. Meng et al. modified PVA scaffolds with collagen to create a novel scaffold for culturing liver cells, demonstrating outstanding cell activity and functionality [128]. Furthermore, PVA-gelatin hydrogel scaffolds have been developed for liver cell culturing [129,130]. Notably, Khati et al. used PVA as a sacrificial biomaterial to provide high luminal patency for perfusion in the culture medium, supporting the creation of cell-dense enlarged liver models [16] (Fig. 6B).

4.4. Polyethers combined with aliphatic polyesters—polyurethane (PU)

PU is a series of synthetic polymers composed of organic units connected by urethane links. Wang et al. combined PEG and PCL to synthesize a novel PU material that is biocompatible, biodegradable, thermosetting, printable, and biostable. This PU is suitable for the fabrication of complex bioartificial organs. PU can also be combined with natural materials, such as gelatin, collagen, and alginate, to make it suitable for tissue engineering applications [11]. While this material is not extensively used in LTE, it holds significant promise and value.

5. Crosslinking of biomaterials

During the construction of tissue-engineered models, polymer solutions are transformed into 3D structures through crosslinking. Crosslinking is crucial for building structurally ideal and biomechanically stable tissue models. Therefore, it is essential to strike a balance between crosslinking methods, crosslinking scales, and printability. This section discusses the common crosslinking methods used in LTE, providing researchers with insights into selecting appropriate bioinks or crosslinking approaches.

5.1. Physical crosslinking

5.1.1. Ionic interactions

Crosslinking via ionic interactions is one of the most common methods for transforming hydrogels into 3D structures because polyvalent cation-induced gelation can be accomplished at temperatures and pH environments suitable for cellular survival. Notably, the ease with which alginate undergoes ion-mediated crosslinking is intertwined with the widespread use of alginate-based biomaterials. Although employing ionic interactions for crosslinking avoids subjecting cells to harsh chemical or thermal conditions, several challenges persist. For instance, there is a negative correlation between the mechanical strength of ion-crosslinked structures and the release of metal cations, leading to increased difficulties in bioprinting and culturing [18,131,132] (Fig. 7A). Furthermore, studies have shown that the concentration of ion-crosslinking agents and duration of exposure directly impact cell viability within the model [133]. High concentrations of Ca²⁺, for example, can disrupt cell membranes. Therefore, various factors, including the crosslinking agent (such as calcium chloride (CaCl₂), calcium sulfate (CaSO₄), and calcium carbonate (CaCO₃)), crosslinking method (e.g., bath-assisted printing, aerosol spraying, and pre-crosslinking), alginate to crosslinking agent ratio, exposure duration, and adjuvant biomaterial (e.g., gelatin, PCL, and GelMA), should be optimized to obtain the intended 3D bioprinting structure [18,134].

Fig. 7.

Crosslinking of biomaterials. (A) Stability of cell-free plotted Alg/MC scaffolds prepared with research- and clinical-grade alginate. Sr²⁺-crosslinking. Scale bars = 5 mm. (b) Storage modulus of plotted scaffolds on day 1 after bioprinting, showing complete crosslinking and a higher crosslinking density for research-grade alginate (Adapted with permission from Duin et al. (2019) [132], Copyright © 2019, WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim). (B) Schematic diagram of 3D bioprinted alginate–chitosan polyion complex hydrogels, which crosslink through electrostatic interactions. (Adapted with permission from Liu et al. (2018) [138], Creative Commons Attribution 4.0 International License). (C) Schematic illustration of the biofabrication process of 135ACG hybrid ink (1% alginate, 3% CNC, and 5% GelMA), which crosslink upon UV irradiation (Adapted with permission from Wu et al. (2020) [43], Creative Commons Attribution 4.0 International License). (D) Experimental design and demonstration of preset extruded hepatic lobule. (a) Schematic illustration of preset extrusion bioprinting of hepatic lobules. (b) Average width of the printed construct. Material 'C' (alginate) dissolves at 37°C to obtain a lumen structure (Adapted with permission from Kang et al. (2020) [26], Copyright © 2020, WILEY-VCH Verlag GmbH & Co. KGaA). (E) (a) SEM images of the cross-sections of crosslinked formulations of dLM-gel, dLM-PEG, and dLM-PEG-T. (b) Average pore size (n = 3, ∗∗p < 0.01 and ∗∗∗p < 0.001) (Adapted with permission from Khati et al. (2022) [16], Copyright © 2021, Wiley-VCH GmbH).

5.1.2. Electrostatic interactions

Biological materials can undergo self-crosslinking through electrostatic interactions, offering a means to circumvent the above issues associated with metal cations, which can disrupt the cell culture microenvironment. In this crosslinking approach, two biomaterials with opposite charges are applied in combination (e.g., positively charged gelatin or chitosan and negatively charged alginate or κ-carrageenan), and the bioink is gelatinized through a network of electrostatic interactions. However, this approach may result in asynchronous crosslinking times as well as inhomogeneous regions [135,136]. Combinations such as κ-carrageenan with GelMA [137], chitosan with alginate [40,138,139], and silica nanoparticles with silk fibroin-conjugated PEG [140] have been explored (Fig. 7B).

5.2. Chemical crosslinking

5.2.1. Photo-crosslinking

Photo-crosslinking is a significant crosslinking method in 3D bioprinting owing to its simplicity, cost-effectiveness, and remote controllability. Photo-crosslinking is typically carried out by irradiating the sample with light radiation, from UV to visible light, in the presence of a photoinitiator. Although UV light sources are readily available, UV light is ionizing and thus hazardous to biological samples. Therefore, factors such as UV exposure time and intensity should be carefully considered [141]. Similar to the susceptibility of alginate to crosslinking via ionic interactions, GelMA stands out as the most readily photopolymerizable biomaterial in 3D bioprinting. In LTE, GelMA has been crosslinked under UV light for various applications, from standalone GelMA utilization to its combination with alginate [43,51,142] (Fig. 7C). Other photo-crosslinking methods used in tissue engineering include thiolene coupling and visible-light crosslinking. For instance, Stichler et al. used thiolated HA (HA-SH) as a bioink and crosslinked it with allyl-functionalized poly(glycidol)s (P(AGE-co-G)) for the 3D bioprinting of mesenchymal stem cell-derived cartilage. Moreover, various synthetic and natural biomaterials, such as PEG, gelatin, and HA, can form thiolene photo-crosslinked bioinks through conjugation with –SH groups [143]. These photo-crosslinking methods offer valuable biomaterials for LTE.

5.2.2. Temperature-controlled crosslinking

Temperature-controlled crosslinking is one of the simplest crosslinking methods, although its application in LTE is somewhat limited owing to the strict temperature requirement for biological activity, typically at or below 37°C. For instance, the dLM-PEG composite bioink developed by Khati et al. undergoes crosslinking within just 45 min at 37°C in a cell culture incubator, ensuring the viability of liver cells within the structure [16]. In contrast, a reverse approach involving temperature control ensures the gelation of biomaterials and also the liquefaction or dissolution of these materials, enabling the construction of macroscopic hollow scaffold structures. For example, Kang et al. employed sodium alginate gel as a sacrificial material and removed it at 37°C in the culture environment, creating liver lobule-like structures with hollow channels. This approach enhances the vitality and functionality of liver cells [26] (Fig. 7D).

5.3. Enzymatic reaction

Enzymes are considered important agents for crosslinking bioinks owing to their activity in mild reaction conditions and minimal impact on the cellular microenvironment. In LTE, enzyme-based crosslinking is often used as an auxiliary method to further enhance structural rigidity. For example, Wang et al. designed composite hydrogel capsules (CHCs) consisting of a core of fibroin hydrogel and composite shell of alginate and chitosan. Building upon the electrostatic complexation of alginate and chitosan, thrombin further gels the fibrinogen, resulting in CHCs with excellent biocompatibility and stability. Liver cells exhibit robust cell viability within this structure [40]. Similarly, Khati et al. crosslinked dLM-PEG-based liver lobule-like structures at 37°C and subsequently reinforced the mechanical properties of the model by adding tyrosinase, enabling the formation of liver lobule-like structures with excellent active liver cell properties [16] (Fig. 7E).

6. Future prospects

True recapitulation of the liver tissue microenvironment is the ultimate goal of LTE. However, considering the current state of tissue engineering and advancements in biomaterials, achieving this goal remains a considerable distance away. True recapitulation requires the restoration of both the structural and compositional aspects of tissue engineering models [144]. 3D bioprinting is a critical tool for precisely depositing bioinks containing cells and ECM components, enabling the reconstruction of complex liver tissue structures. As a result, it is considered one of the most crucial approaches in LTE. Micro-scale structural recapitulation depends on the properties of the biomaterial, including scaffold porosity, pore size, and biomechanical characteristics, which directly influence cell activity and function. Recapitulation of the composition relies on the ability of the biomaterial to mimic the ECM of liver tissues. Currently, both natural and synthetic biomaterials have found extensive applications in LTE. While each type of biomaterial has its unique potential and advantages, a single material alone often cannot fulfill the requirements of tissue engineering tasks. Undoubtedly, the biomaterial that comes closest to replicating the natural ECM is the liver-dECM. However, its applicability is limited owing to factors such as immunogenicity, low printability, and low mechanical strength. Therefore, researchers have begun to find ways to remove the immunogenicity of the dECM by fusing it with a variety of biomaterials to form multicomponent bioinks [16,145,146]. Multi-component bioinks can circumvent the limitations associated with single-component hydrogel bioinks, such as poor print fidelity, poor shape retention, poor biofunctionality, and poor cell guidance. The future direction lies in advancing multi-component bioinks based on the dECM, incorporating both natural and synthetic materials, to recapitulate the structural and compositional aspects of liver tissue models more accurately.

It is worth noting that the liver is a highly complex organ with abundant tubular structures, which presents challenges for faithfully replicating the organized arrangement of arteries, veins, and bile ducts within liver lobules through 3D bioprinting for accurate in vitro liver reconstruction. At this juncture, the strategic use of sacrificial materials becomes advantageous. As mentioned earlier, water-soluble materials, such as like PVA, hold immense potential in creating intricate tubular structures [16]. Furthermore, it is imperative to consider integrating various technologies for LTE. The synergistic combination of 3D bioprinting and microfluidic techniques, for instance, proves to be complementary, collectively contributing to the construction of in vitro liver models. Zheng et al. leveraged photo-induced electrodeposition (PIED) and microfluidic techniques to pioneer the preparation of cell-laden alginate hydrogel microcapsules, offering unprecedented possibilities for future applications in LTE [147].

Crosslinking is a critical factor that influences not only the printability and chemical properties of the bioink but also the mechanical strength and biological microenvironment of the tissue-engineered structure. There exists a delicate balance between the degree of crosslinking and the desired structure. Low degrees of crosslinking can result in loose print structures that provide insufficient support, while high degrees of crosslinking may affect the printing process and cell viability. Optimizing the crosslinking method and degree according to the target structure is, therefore, of paramount importance. Moreover, considering the effect of crosslinking on biological components, maintaining high cell viability is a fundamental prerequisite in complex 3D bioprinting. Currently employed crosslinking techniques in LTE, such as photo-crosslinking and temperature-controlled crosslinking, can directly influence biological activity, whereas ion-mediated crosslinking can indirectly affect cells by altering the culture microenvironment. Therefore, it is foreseeable that crosslinking methods with minimal biological hazards, such as visible-light crosslinking and enzyme-mediated crosslinking, hold promise as focal points for future research.

In the realm of clinical applications, advancements in biomaterials for LTE are poised to substantively improve diagnostic and therapeutic outcomes for patients. Through the precise construction of intricate liver structures, encompassing cells, ECM, and vascular systems, more realistic and biologically relevant in vitro models are achieved. In terms of diagnostics, these in vitro liver models can provide clinical practitioners with a better understanding of patient conditions by offering accurate anatomical and pathological features. In the realm of therapeutics, analyzing the metabolic activity and drug responsiveness of patient-specific in vitro models enables clinicians to optimize treatment plans, enhancing the efficacy and individualization of therapies. Moreover, the application of biomaterials enhances the feasibility of liver regeneration and repair. Customized scaffolds and carriers can facilitate the transplantation of highly active hepatocytes into the patient's liver, thereby improving liver function. Furthermore, the future prospect of creating entire organ models through bioprinting allows clinicians to perform simulated operations before surgery, predicting potential surgical challenges to enhance procedures and success rates.

Translating LTE advancements into clinical practice necessitates collaboration between researchers and healthcare professionals, with researchers assisting in addressing specific clinical needs articulated by physicians. Numerous hurdles, such as production efficiency, economic viability, ethical considerations, patient acceptance, long-term safety and efficacy, as well as regulatory oversight, must also be surmounted. In essence, integrating these innovations into clinical practice demands proactive collaboration across multiple sectors and patient engagement to fully realize the potential of LTE in guiding disease treatment and improving patient prognosis.

7. Conclusion

Our review is primarily focused on the utilization of natural and synthetic biomaterials, as well as their combined applications, in 3D bioprinting approaches for LTE. Additionally, we delved into the various crosslinking methods applicable to biomaterials. This review is expected to offer new perspectives for overcoming challenges in LTE and provide valuable references for realizing liver tissue models with applications in drug testing, disease exploration, and alternative therapies, ultimately contributing to the advancement of LTE.

Funding

This research was funded by the National Natural Science Foundation of China (Nos 32371477, 82090051, 32000970, and 92168207), the National Key Research and Development Program of China (Nos 2022YFA1103400 and 2022YFC2406704), and the Tsinghua Precision Medicine Foundation (No. 2022TS013).

Author contributions

Conceptualization, Q.W. and J.L.; original draft preparation, Q.W., J.L., Y.F., A.W. and J.Z.; draft review and editing, Y.H., Y.C., Y.L. and J.L.; supervision, J.L.; funding acquisition, J.L.

Acknowledgments

None.

Declaration of competing interest

The authors declare no conflict of interest.

Data available statement

None.

Ethics statement

None.

Informed consent

None.