Bionic tissue implant and preparation method therefor

Abstract

The present invention provides a bionic tissue implant and a preparation method therefor. The bionic tissue implant has good biocompatibility and mechanical properties and can be used for repairing tissue defects. The bionic tissue implant comprises: a decellularized tissue layer; a first biological scaffold layer arranged on one side of the decellularized tissue layer and comprising a first biocompatible high molecular polymer and a target cell; and a second biological scaffold layer arranged on the other side of the decellularized tissue layer relative to the first biological scaffold layer and comprising a second biocompatible high molecular polymer. The first biocompatible macromolecule is the same as or different from the second biocompatible high molecular polymer.

Description

Bionic tissue implant and preparation method thereof Technical Field

[0001] The present invention relates to tissue engineering, in particular to implants used in tissue engineering, especially implants with biocompatibility; the present invention also relates to the preparation and use of implants used in tissue engineering. Background Art

[0002] Fibrous cartilage, elastic cartilage, and hyaline cartilage are three different forms of cartilage in the human body, each with specific biological and structural properties. Hyaline cartilage, a thin connective tissue that anchors the distal ends of bones and is covered by synovial fluid, is known as articular cartilage and is used to withstand the compression, tension, and shear forces of joints such as the shoulder, elbow, and knee. Articular cartilage provides a low-friction, lubricated, and smooth surface for bearing and transmitting loads during joint movement.

[0003] Adult articular cartilage consists of 95-99% porous and permeable extracellular matrix (ECM) and approximately 1-5% chondrocytes. The ECM is composed of 80% water and 20% collagen, proteoglycans, and other components. Collagen is the most abundant macromolecule in articular cartilage, accounting for 90-95% of the total collagen in the cartilage ECM. Proteoglycans are the second most abundant solid component of the cartilage matrix and consist of a core protein covalently bound to glycosaminoglycans (GAGs). Like collagen, proteoglycans and GAGs play a major role in depth-dependent load-bearing capacity.

[0004] Articular cartilage has regional heterogeneity and is divided into superficial, middle, deep and calcified zones, each with its own specific matrix composition, chondrocyte type, distribution, biomolecular orientation and organization. Chondrocytes are the only resident cells in articular cartilage and are region-dependent. They are responsible for synthesizing, constructing, maintaining and repairing ECM components to stabilize articular cartilage. Because chondrocytes exist in the lacunar space and are separated from the pericellular matrix region, they have poor proliferation ability and low cell density, resulting in a lack of intercellular interaction and limited self-renewal ability. In addition, the lack of blood vessels, nerves and lymphatic vessels in articular cartilage makes it difficult for chondrocytes to regenerate themselves. Therefore, once the cartilage is damaged and conservative treatment is lacking, chondrocytes will develop osteoarthritis (OA) due to the lack of direct nutritional supply. This is a major disability in the elderly and a major burden on global healthcare.

[0005] Osteoarthritis is a common degenerative joint disease and a major disabling condition. The articular cartilage covering the ends of bones gradually breaks down and thins, leading to stiffness, swelling, and pain. In some severe cases, the cartilage can wear out and tear completely, causing severe pain, inflammation, and even disability. There are many risk factors associated with osteoarthritis, such as gender, family history, obesity, trauma, and age. Due to the impact of an aging society, although osteoarthritis is not a fatal disease, it poses a considerable burden worldwide. In addition, the knee joint is the largest synovial joint and is the most susceptible to osteoarthritis due to its high usage rate. According to the latest statistics, there were 654.1 million people with knee osteoarthritis worldwide in 2020.

[0006] Currently, the simplest method for repairing cartilage defects in varying degrees of knee osteoarthritis is the injection of compounds including hyaluronic acid (HA), but clinical trials have yet to demonstrate significant results. Other surgical approaches used clinically include mosaicplasty, microfracture, autologous chondrocyte implantation (ACI), and matrix-induced autologous chondrocyte implantation (MACI). These conventional techniques still have significant limitations, including high costs and limited effectiveness in forming fibrocartilage.

[0007] The long-term performance of remodeled fibrocartilage has been shown to be limited by its biochemical and biomechanical properties compared to hyaline cartilage. Therefore, novel strategies are attractive as they could address the poor outcomes of mainstream clinical management. In the latest area of ​​cartilage research, tissue engineering is one of the alternative therapies gaining momentum to promote cartilage regeneration.

[0008] At present, cartilage tissue engineering can provide a new treatment strategy for the above dilemma. Tissue engineering comprehensively considers cell sources, engineering scaffolds and biochemical factors to restore, maintain, improve or replace tissues. For example, growth factors such as transforming growth factor-β family (TGF-β), fibroblast growth factor (FGF), insulin-like growth factor-I (IGF-1) and platelet-derived growth factor (PDGF) have shown the potential for cartilage repair; the ideal cell source must be able to promote hyaline cartilage regeneration and differentiate into chondrocytes when transplanted to the damaged site, such as autologous chondrocytes, mesenchymal stem cells (MSC) and induced pluripotent stem cells (iPSC).

[0009] Bioscaffolds are crucial in cartilage tissue engineering, providing an environment for adhesion, proliferation, and differentiation. In general tissue engineering, the primary goal is to replicate the structure of the extracellular matrix (ECM). Conventional scaffold materials are generally synthetic and natural polymers. For example, polylactic-co-glycolic acid (PLGA) and polyethylene glycol (PEG) are standard synthetic matrices that provide a porous environment to promote cell attachment and growth. Compared to synthetic scaffolds, scaffolds made from polysaccharides and proteins can reduce inflammation and immune responses. Several commercially available bioscaffolds using collagen, fibrin, and chitosan are available. However, these products have fallen out of favor with clinical users due to their lack of safety and efficacy.

[0010] Accordingly, decellularized materials have received increasing attention in the field of tissue engineering. Decellularized materials retain the original tissue ECM and remove genetic material, reducing its immunogenic response and providing a broad path for xenogeneic transplantation.

[0011] Conventional methods for removing cells from tissues include chemical, physical, and biological methods, which are used to remove cells while minimizing damage to the ECM structure. However, chemical methods cannot avoid inflammatory reactions and collagen fiber damage caused by residual reagents, while physical methods that utilize temperature differences or ultra-high hydrostatic pressures are difficult to completely remove cells from thicker tissues. Biological methods that utilize specific enzymes to degrade nucleic acid components can remove unwanted cells, but they take a long time and still have the disadvantage of inducing immune responses.

[0012] In the existing technology, the most common decellularized material is decellularized cartilage powder, which retains natural cartilage GAGs and collagen and has the advantages of promoting cartilage differentiation and repairing cartilage defects. On the other hand, hydrogels are also widely used in tissue engineering and can provide cells with an appropriate three-dimensional environment. Due to the high hydration properties of hydrogels, they can retain a large amount of water to mimic the high water content environment in the ECM. Hydrogels are also biocompatible and can be injected for minimally invasive clinical regenerative medicine applications. In irregular cartilage defects, hydrogels have encapsulation capabilities and can promote cell aggregation and differentiation.

[0013] However, decellularized cartilage powder lacks sufficient mechanical properties and is not conducive to cell seeding. Similarly, although hydrogels have excellent biocompatibility and can be implemented in clinical bone and joint repair by simple injection, the mechanical properties of hydrogels themselves are poor and they cannot resist the pressure and shear force caused by joint movement in the long term.

[0014] In regenerative medicine, ideal tissue engineering materials must meet the following prerequisites: (1) flexible shape to adapt to irregular defects, (2) meet the mechanical properties of the target tissue, and (3) mimic the extracellular matrix to adapt to irregular defects to create an environment suitable for cell differentiation and growth; therefore, there is an urgent need to create an innovative composite scaffold for cartilage tissue repair in regenerative medicine. Summary of the Invention

[0015] Based on the technical issues encountered in the aforementioned field, the main purpose of the present invention is to provide an innovative composite scaffold that combines decellularized tissue and biocompatible polymers for the repair of tissue defects. Furthermore, multiple layers of decellularized tissue and biocompatible polymers can be stacked together to establish a novel sandwich composite scaffold model, and its biochemical, biophysical, and biocompatibility properties and applicability can be further verified.

[0016] Accordingly, an object of the present invention is to provide a biomimetic tissue implant (100), comprising: a decellularized tissue layer (10); a first bioscaffold layer (1), disposed on one side of the decellularized tissue layer (10), comprising a first biocompatible polymer and target cells; and a second bioscaffold layer (2), disposed on the other side of the decellularized tissue layer (10) relative to the first bioscaffold layer (1), comprising a second biocompatible polymer, wherein the first biocompatible polymer is the same as or different from the second biocompatible polymer.

[0017] The biomimetic tissue implant (100) as described above, wherein the first biocompatible polymer is cross-linked with the acellular tissue layer (10) through multiple covalent bonds, and the second biocompatible polymer is cross-linked with the acellular tissue layer (10) through multiple covalent bonds.

[0018] The biomimetic tissue implant (100) as described above, wherein the first biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof; and the second biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof.

[0019] The bionic tissue implant (100) as described above, wherein the target cells include fibroblasts, glial cells, interstitial cells, embryonic stem cells, umbilical cord mesenchymal stem cells, bone marrow mesenchymal stem cells, adipose stem cells or lymphocytes.

[0020] The bionic tissue implant (100) as described above, wherein the decellularized tissue layer (10) is a decellularized tissue, which includes decellularized liver tissue, decellularized fat, decellularized skin, decellularized cartilage, decellularized myocardial tissue, decellularized pericardium or decellularized blood vessels.

[0021] Another object of the present invention is to provide a method for preparing a biomimetic tissue implant (100), which comprises: taking a decellularized tissue and planting target cells on one side of the decellularized tissue; taking a first biocompatible coating comprising a first biocompatible precursor polymer and applying the first biocompatible coating to the side of the decellularized tissue on which the target cells are planted to form a first bioscaffold layer (1), wherein the first biocompatible precursor polymer comprises a first biocompatible high molecular polymer; and taking a second biocompatible coating comprising a second biocompatible precursor polymer and applying the second biocompatible coating to the side of the decellularized tissue opposite to which the target cells are planted to form a second bioscaffold layer (2), so as to obtain the biomimetic tissue implant (100), wherein the second biocompatible precursor polymer comprises a second biocompatible high molecular polymer, wherein the first biocompatible high molecular polymer is the same as or different from the second biocompatible high molecular polymer.

[0022] Another object of the present invention is to provide a method for preparing a biomimetic tissue implant (100), which comprises: taking a decellularized tissue as a decellularized tissue layer; mixing a first biocompatible coating and target cells to obtain a first biocompatible cell mixture, and coating the first biocompatible cell mixture on one side of the decellularized tissue layer (10) to form the first bioscaffold layer (1), wherein the first biocompatible precursor polymer comprises a first biocompatible high molecular polymer; and taking a second biocompatible coating comprising a second biocompatible precursor polymer, and coating the second biocompatible coating on the side of the decellularized tissue opposite to the side where the target cells are implanted to form a second bioscaffold layer (2), so as to obtain the biomimetic tissue implant (100), wherein the second biocompatible precursor polymer comprises a second biocompatible high molecular polymer, and wherein the first biocompatible high molecular polymer is the same as or different from the second biocompatible high molecular polymer.

[0023] The preparation method of the biomimetic tissue implant (100) as described above, wherein the first biocompatible coating comprises: a first leading coating, which comprises the first biocompatible precursor polymer and a first cross-linking enzyme; and a second leading coating, which comprises the first biocompatible precursor polymer and a first cross-linking agent; the second biocompatible coating comprises: a third leading coating, which comprises the second biocompatible precursor polymer and a second cross-linking enzyme; and a fourth leading coating, which comprises the second biocompatible precursor polymer and a second cross-linking agent, wherein, before forming the first bioscaffold layer (1) and forming the second bioscaffold layer (2), the preparation method of the biomimetic tissue implant (100) further comprises: simultaneously or separately applying the first leading coating and the second leading coating to the debonding layer. The cell tissue is seeded with the target cells on one side to obtain a first leading scaffold layer; the third leading coating and the fourth leading coating are simultaneously or separately applied to the side of the decellularized tissue opposite to the target cells to obtain a second leading scaffold layer; the first cross-linking enzyme is brought into contact with the first cross-linking reagent for a first time to cross-link the first biocompatible polymer with the decellularized tissue layer (10) to form the first bioscaffold layer (1); the second cross-linking enzyme is brought into contact with the second cross-linking reagent for a second time to cross-link the second biocompatible polymer with the decellularized tissue layer (10) to form the second leading scaffold layer (2), wherein the first time is equal to or less than the second time.

[0024] The preparation method of the bionic tissue implant (100) as described above, wherein the first cross-linking enzyme comprises horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxidoreductase, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase or catechol oxidase; and the second cross-linking enzyme comprises horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxidoreductase, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase or catechol oxidase.

[0025] The preparation method of the biomimetic tissue implant (100) as described above, wherein the preparation method of the first biocompatible precursor polymer comprises: mixing water and the first biocompatible high molecular polymer to obtain a first aqueous solution; preparing a bridging aqueous solution and mixing it with the first aqueous solution to obtain the first biocompatible precursor; the preparation method of the second biocompatible precursor polymer comprises: mixing water and the second biocompatible high molecular polymer to obtain a second aqueous solution; taking the bridging aqueous solution and mixing it with the second aqueous solution to obtain the second biocompatible precursor, wherein the preparation method of the bridging aqueous solution comprises: mixing a biphasic solvent and a bridging agent to obtain a bridging precursor solution, wherein the biphasic solvent comprises water and an organic solvent; and adding a curing agent to the bridging precursor solution to activate the bridging agent to obtain the bridging aqueous solution. BRIEF DESCRIPTION OF THE DRAWINGS

[0026] FIG1A illustrates components and their corresponding relationships in a first embodiment of the present invention;

[0027] FIG1B illustrates components and their corresponding relationships in a second embodiment of the present invention;

[0028] 1C to 1D illustrate components and their corresponding relationships in a third embodiment of the present invention;

[0029] FIG1E illustrates components and their corresponding relationships in a fourth embodiment of the present invention;

[0030] 2A to 2C are flow charts illustrating the steps of the preparation method according to the tenth embodiment of the present invention;

[0031] FIG2D is a flow chart illustrating the steps of the preparation method according to the eleventh embodiment of the present invention;

[0032] 2E to 2F are flow charts illustrating the steps of the preparation method according to the twelfth embodiment of the present invention;

[0033] Figures 3a to 3f are macroscopic images of porcine cartilage discs and decellularized cartilage slices;

[0034] 4a to 4d are histograms showing the biocompatibility test results of acellular cartilage slices;

[0035] Figures 5a to 5d are the UV absorbance, FI-IR transmittance and 1 H-NMR peak shift spectrum;

[0036] FIG6 shows the appearance and microstructure of the hydrogels provided in Examples 1 to 3;

[0037] 7a to 7b are line graphs showing the swelling ratio test results of Examples 1 to 3;

[0038] Figures 7c to 7d are line graphs showing the in vitro degradation test results of Examples 1 to 3;

[0039] 7e to 7f are line graphs showing the mass loss evaluation results of Examples 1 to 3;

[0040] FIG8 is a histogram showing the protein adsorption test results of Examples 1 to 3;

[0041] 9a to 9b are line graphs showing the rheological properties test results of Examples 1 to 3;

[0042] 10a to 10d are histograms showing the biocompatibility tests of Examples 1 to 3 on different cell lines;

[0043] FIG11 shows a plurality of cell images showing the cell morphology after the cells were respectively implanted in Examples 1 to 3;

[0044] FIG12 is a cartoon diagram illustrating the specific process of Example 4;

[0045] FIG13a is a histogram illustrating the adhesion strength test results of Examples 1 to 6;

[0046] FIG13 b is a bar graph illustrating the compression modulus test results of Examples 1 to 3 and 7 to 9;

[0047] FIG14 shows the appearance and morphology of the hydrogels provided in Examples 10 to 12, wherein Example 10 is shown in (a), (d), and (g), Example 11 is shown in (b), (e), and (h), and Example 12 is shown in (c), (f), and (i). The scale bars in (d), (e), and (f) are 1 mm, and the scale bars in (g), (h), and (i) are 500 μm.

[0048] FIG15 is HE tissue staining images showing the cross-section and top view HE tissue staining morphologies of Examples 13 to 15 at 1 and 5 days after cell seeding. The scale bar is 100 μm.

[0049] FIG16 is a diagram of SO tissue staining, showing the cross-section and top view of SO tissue staining morphology of Examples 13 to 15 at 1 and 5 days after cell seeding. The scale bar is 100 μm.

[0050] FIG17 is a diagram of MT tissue staining, showing the cross-section and top view MT tissue staining morphology of Examples 13 to 15 after cell seeding and culture for 1 and 5 days, with a scale bar of 100 μm;

[0051] Figures 18 to 19 show the Live / Dead staining results of cross-sections and top views of Examples 13 to 15 on day 1 after culture, respectively. The scale bar is 100 μm.

[0052] Figures 20 to 21 show the Live / Dead staining results of cross-sections and top views of Examples 13 to 15 on day 5 after culture, respectively. The scale bar is 100 μm.

[0053] Figures 23 to 24 are histograms showing the results of type II collagen quantification in Examples 16 to 18 and Examples 13 to 15, respectively;

[0054] FIG25 is a cartoon diagram illustrating the specific process of Example 20;

[0055] 26A to 26B are macroscopic structure images of Example 19 and Example 20;

[0056] 26C to 26E are SEM images showing the microstructures of Examples 19 and 20, wherein the scale bars in FIG. 26C and 26E are 100 μm, and the scale bar in FIG. 26D is 500 μm;

[0057] FIG27 is a histogram showing the compression modulus test results of Example 19 and Example 20;

[0058] FIG28 is a histogram showing the cell activity test results of Example 19 and Example 20;

[0059] FIG29 is a block flow chart illustrating the experimental design and implementation process of Experimental Example 5;

[0060] FIG30 is an image of a New Zealand white rabbit knee joint sample after surgery in Experimental Example 5;

[0061] FIG31 is a histogram showing the cartilage macroscopic scoring results of the New Zealand white rabbit knee joint specimens after surgery in Experimental Example 5;

[0062] FIG32A is a representative 2D micro-CT image of the New Zealand white rabbit knee joint specimen after surgery in Experimental Example 5;

[0063] FIG31 is a histogram showing the bone tissue recovery results of the New Zealand white rabbit knee joint specimens after surgery in Experimental Example 5;

[0064] 33A and 33B are HE staining images showing the tissue regeneration status of the knee joint specimens of New Zealand white rabbits after surgery in Experimental Example 5;

[0065] FIG34 is a SO tissue staining image showing the matrix glycosaminoglycan content of the New Zealand white rabbit knee joint specimens after surgery in Experimental Example 5;

[0066] FIG35 is a histogram showing the tissue staining scoring results of the New Zealand white rabbit knee joint samples after surgery in Experimental Example 5. DETAILED DESCRIPTION

[0067] In order to more clearly illustrate the spirit of the present invention and its specific implementation, several embodiments are listed and described in detail with reference to the drawings as follows: Please refer to Figure 1A, which illustrates the first embodiment of the present invention, which provides a biomimetic tissue implant (100), which includes a decellularized tissue layer (10), a first bioscaffold layer (1) and a second bioscaffold layer (2), wherein the first bioscaffold layer (1) and the second bioscaffold layer (2) are respectively arranged on both sides of the decellularized tissue layer (10), and together with the decellularized tissue layer (10) form a sandwich structure.

[0068] In this embodiment, the decellularized tissue layer (10) is a decellularized tissue, which includes decellularized liver tissue, decellularized fat, decellularized skin, decellularized cartilage, decellularized myocardial tissue, decellularized pericardium or decellularized blood vessels; the specific method of obtaining the decellularized tissue is not limited, and includes chemical methods such as sodium dodecyl sulfate (SDS) and TritonX-100 and other chemical reagents to remove cells, or physical methods such as temperature difference or ultra-high hydrostatic pressure to physically destroy the cell membrane and wash away the cells, or enzyme methods using specific enzymes to degrade cell nuclei and membrane components; in multiple embodiments, the decellularized tissue is not limited to chemical methods such as sodium dodecyl sulfate (SDS) and TritonX-100 to remove cells, or physical methods such as temperature difference or ultra-high hydrostatic pressure to physically destroy the cell membrane and wash away the cells, or enzyme methods using specific enzymes to degrade cell nuclei and membrane components. The decellularized tissue can be sheet-shaped, plate-shaped, mesh-shaped, etc., without special restrictions; in some preferred embodiments, the decellularized tissue is sheet-shaped, specifically, decellularized liver tissue slices, decellularized fat slices, decellularized skin slices, decellularized cartilage slices, decellularized myocardial tissue slices; preferably, any of the decellularized tissue layers (10) is a decellularized cartilage slice; on the other hand, the thickness of the decellularized tissue layer (10) can be selected according to the mechanical performance requirements of the bionic tissue implant (100), which can be 5 to 200 μm. In some embodiments, its thickness is 5 to 50 μm; preferably, its thickness is 10 to 20 μm.

[0069] In this embodiment, in order to improve its biocompatibility, promote cell aggregation performance and enhance mechanical properties, as an implant used in tissue engineering, for example, for repairing tissue defects such as cartilage and bone, the first bioscaffold layer (1) includes a first biocompatible polymer, and the second bioscaffold layer (2) includes a second biocompatible polymer, and the first biocompatible polymer is the same as or different from the second biocompatible polymer; in multiple embodiments, the first biocompatible polymer is the same as the second biocompatible polymer; specifically, the first biocompatible polymer can be fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof; the second biocompatible polymer can be fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof; in a preferred embodiment, the first biocompatible polymer and the second biocompatible polymer are gelatin.

[0070] In order to make the first biocompatible polymer and the second biocompatible polymer have the required mechanical properties and thermal stability, in this embodiment, before forming the first bioscaffold layer (1) and the second bioscaffold layer (2), a first biocompatible precursor polymer mixed with the first biocompatible polymer and a second biocompatible precursor polymer mixed with the second biocompatible polymer are mixed, and the intermolecular covalent bonding of the first biocompatible polymer and the second biocompatible polymer is adjusted by enzyme cross-linking, and the degradation kinetics and rheological properties thereof are adjusted at the same time. It is more suitable for the manufacture and use of bionic tissue implants (100); taking gelatin as an example, it can be functionalized with hydroxyphenyl groups, combined with the contact of peroxidase and peroxide, to initiate a cross-linking mechanism in an enzyme-mediated manner; in addition, this enzyme cross-linking mechanism also enables the bionic matrix gel to react with phenol-rich molecules, such as the extracellular matrix ECM in the decellularized tissue layer (10) or the target tissue to produce a cross-linking reaction, thereby enhancing the binding of the bionic tissue implant (100) with the ECM around the tissue defect to enhance its regenerative effect.

[0071] In some specific examples, the cross-linking enzyme can be horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxiredoxin, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase, or catechol oxidase; in some preferred embodiments, the cross-linking enzyme is horseradish peroxidase (HRP).

[0072] In this embodiment, in order to provide the bionic tissue implant (100) with good tissue repair performance, target cells are added to the corresponding specific target tissue; in multiple embodiments, the bionic matrix gel and the target cells together form the regeneration scaffold layer; in these embodiments, the target cells include but are not limited to mesenchymal stem cells, adipose tissue stem cells, endothelial progenitor cells, mesenchymal skeletal stem cells, synovial stem cells, osteoblasts, chondrocytes, and osteocytes; in other embodiments, the target cells can also be fibroblasts, glial cells, mesenchymal cells, embryonic stem cells, umbilical cord mesenchymal stem cells, bone marrow mesenchymal stem cells or lymphocytes; in some preferred embodiments, the target cells are mesenchymal stem cells, which can be derived from bone marrow, adipose tissue, umbilical cord tissue, umbilical cord blood, peripheral blood, dental pulp, amniotic fluid, amniotic membrane, synovium, and periosteum; in one example, the mesenchymal stem cells are derived from the infrapatellar fat pad.

[0073] In this embodiment, the target cells correspond to a cell density on the decellularized tissue layer (10) ranging from 1 to 2×10 6 cells / cm²; preferably, the cell density is between 1.2 and 1.8×10 6 cells / cm²; more preferably, the cell density is between 1.5 and 1.7×10 6 cells / cm².

[0074] Please refer to Figure 1B. In the second embodiment of the present invention, the specific structure and components of the bionic tissue implant (100) are roughly the same as those of the first embodiment and are not described in detail here. However, the decellularized tissue layer (10) includes two or more sub-decellularized tissue layers (10a). These sub-decellularized tissue layers (10a) can be sheet-shaped, plate-shaped, mesh-shaped, etc., without special restrictions. In some preferred embodiments, any of the sub-decellularized tissue layers (10a) is sheet-shaped, specifically decellularized liver tissue slices, decellularized fat slices, decellularized skin slices, decellularized cartilage slices, and decellularized myocardial tissue slices. Preferably, any of the sub-decellularized tissue layers (10a) is a decellularized cartilage slice. On the other hand, any of the sub-decellularized tissue layers (10a) can be 5 to 200 μm thick. In some embodiments, its thickness is 5 to 50 μm. Preferably, its thickness is 10 to 20 μm.

[0075] 1C to 1D , in the third embodiment of the present invention, the specific structure and components of the biomimetic tissue implant (100) are substantially the same as those of the first embodiment, and are not described in detail herein. However, the biomimetic tissue implant (100) further includes another acellular tissue layer (10′) disposed between the acellular tissue layer (10) and the first bioscaffold layer (1), or disposed between the acellular tissue layer (10) and the second bioscaffold layer (2), wherein the another acellular tissue layer (10′) is disposed between the acellular tissue layer (10) and the first bioscaffold layer (1). The tissue sources are different, wherein the other decellularized tissue layer (10') can be sheet-shaped, plate-shaped, mesh-shaped, etc., without special restrictions; in some preferred embodiments, the other decellularized tissue layer (10') is sheet-shaped, specifically, decellularized liver tissue slices, decellularized fat slices, decellularized skin slices, decellularized cartilage slices, decellularized myocardial tissue slices; on the other hand, the other decellularized tissue layer (10') can be 5 to 200 μm, and in some embodiments, its thickness is 5 to 50 μm; preferably, its thickness is 10 to 20 μm.

[0076] Please refer to FIG. 1E . In the fourth embodiment of the present invention, the specific structure and components of the bionic tissue implant (100) are substantially the same as those of the second embodiment and are not described in detail here. However, a third biocompatible layer (3) is provided between each of the sub-acellular tissue layers (10a), which is the same as the first bioscaffold layer (1) or the second bioscaffold layer (2).

[0077] In the fifth embodiment of the present invention, the specific structure and components of the biomimetic tissue implant (100) are substantially the same as those of the fourth embodiment and are not described in detail herein, except that the third biocompatible layer is the same as the first bioscaffold layer (1) but different from the second bioscaffold layer (2).

[0078] In the sixth embodiment of the present invention, the specific structure and components of the biomimetic tissue implant (100) are substantially the same as those of the fourth embodiment and are not described in detail herein, except that the third biocompatible layer is different from the first bioscaffold layer (1) but is the same as the second bioscaffold layer (2).

[0079] In the seventh embodiment of the present invention, the specific structure and components of the biomimetic tissue implant (100) are substantially the same as those of the fourth embodiment and are not described in detail herein, except that the third biocompatible layer is different from the first bioscaffold layer (1) and the second bioscaffold layer (2).

[0080] In the eighth embodiment of the present invention, the specific structure and components of the bionic tissue implant (100) are substantially the same as those of the first embodiment and are not described in detail herein, except that the second bioscaffold layer (2) further includes auxiliary cells, wherein the auxiliary cells are the same as the target cells.

[0081] In the ninth embodiment of the present invention, the specific structure and components of the bionic tissue implant (100) are roughly the same as those of the first embodiment and are not described in detail here, except that the second biological scaffold layer (2) further includes auxiliary cells, wherein the auxiliary cells are different from the target cells, and the auxiliary cells include but are not limited to mesenchymal stem cells, adipose tissue stem cells, endothelial progenitor cells, mesenchymal skeletal stem cells, synovial stem cells, osteoblasts, chondrocytes, and osteocytes; in other embodiments, the auxiliary cells can also be fibroblasts, glial cells, mesenchymal cells, embryonic stem cells, umbilical cord mesenchymal stem cells, bone marrow mesenchymal stem cells or lymphocytes; in some preferred embodiments, the auxiliary cells are mesenchymal stem cells, which can be derived from bone marrow, adipose tissue, umbilical cord tissue, umbilical cord blood, peripheral blood, dental pulp, amniotic fluid, amniotic membrane, synovium, and periosteum.

[0082] Please refer to FIG2A , which illustrates the tenth embodiment of the present invention, which provides a method for preparing a bionic tissue implant (100). Specifically, the method is used to prepare the bionic tissue implant (100) provided by the present invention, and includes the following steps:

[0083] Step S1: taking a decellularized tissue as a cell tissue layer (10), and planting target cells on one side of the decellularized tissue;

[0084] Step S2: taking a first biocompatible coating comprising a first biocompatible precursor polymer, and coating the first biocompatible coating on the side of the decellularized tissue where the target cells are implanted, to form a first bioscaffold layer (1), wherein the first biocompatible precursor polymer comprises a first biocompatible high molecular polymer; and

[0085] Step S3: taking a second biocompatible coating comprising a second biocompatible precursor polymer, and applying the second biocompatible coating to the side of the decellularized tissue opposite to the side where the target cells are implanted to form a second bioscaffold layer (2), thereby obtaining a biomimetic tissue implant (100), wherein the second biocompatible precursor polymer comprises a second biocompatible high molecular polymer, and wherein the first biocompatible high molecular polymer is the same as or different from the second biocompatible high molecular polymer.

[0086] In the tenth embodiment, the step S2 and the step S3 may be performed simultaneously, or the step S2 may be performed sequentially before the step S3, or the step S3 may be performed sequentially before the step S2, without any particular limitation.

[0087] In the tenth embodiment, the target cells correspond to a cell density on the decellularized tissue layer (10) ranging from 1 to 2×10 6 cells / cm²; preferably, the cell density is between 1.2 and 1.8×10 6 cells / cm²; more preferably, the cell density is between 1.5 and 1.7×10 6 cells / cm².

[0088] In the tenth embodiment, in order to form the first bio-scaffold layer (1) and the second bio-scaffold layer (2) by in situ gelation of the first biocompatible coating and the second biocompatible coating, the first bio-scaffold layer comprises a first leading coating and a second leading coating, wherein the first leading coating comprises the first biocompatible precursor polymer and a first cross-linking enzyme, and the second leading coating comprises the first biocompatible precursor polymer and a first cross-linking agent; please continue to refer to FIG. 2B , before forming the first bio-scaffold layer (1), the step S2 further comprises:

[0089] Step S2a: simultaneously or separately coating the first leading coating and the second leading coating on the side of the acellular tissue layer (10) seeded with the target cells to obtain a first leading scaffold layer; and

[0090] Step S2b: contacting the first cross-linking enzyme with the first cross-linking reagent for a first time to cross-link the first biocompatible polymer with the acellular tissue layer (10), so that the first leading scaffold layer forms the first biological scaffold layer (1).

[0091] On the other hand, the second biocompatible coating comprises a third leading coating and a fourth leading coating, wherein the third leading coating comprises the second biocompatible precursor polymer and a second cross-linking enzyme, and the fourth leading coating comprises the first biocompatible precursor polymer and a second cross-linking agent; referring to FIG2B , before forming the second bioscaffold layer (2), the preparation method of the biomimetic tissue implant (100) further comprises:

[0092] Step S3a: simultaneously or separately coating the third leading coating and the fourth leading coating on the side of the acellular tissue layer (10) opposite to the side where the target cells are planted, to obtain a second leading scaffold layer;

[0093] Step S3b: contacting the second cross-linking enzyme with the second cross-linking reagent for a second time to cross-link the second biocompatible polymer with the acellular tissue layer (10), so that the second leading scaffold layer forms the second bioscaffold layer (2).

[0094] In the tenth embodiment, the first time is equal to or less than the second time, the first time is between 1 and 10 minutes, and the second time is between 1 and 10 minutes; preferably, the first time is equal to the second time, which is between 3 and 7 minutes.

[0095] In a tenth embodiment, the first precursor coating comprises 1 to 10 units / mL of a first cross-linking enzyme, such as horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxiredoxin, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase, or catechol oxidase; in some preferred embodiments, the first cross-linking enzyme is horseradish peroxidase (HRP).

[0096] In the tenth embodiment, the second precursor coating comprises 0.005 to 0.08 wt% of a first crosslinking agent such as a reactive oxygen species, wherein the reactive oxygen species include but are not limited to urea peroxide (CH6N2O3), peracetic acid (C2H4O3), sodium percarbonate (2Na2CO3·H2O2), di-tert-butyl peroxide (DTBP), benzoyl peroxide (C 14 H 10 O4), tert-butyl hydroperoxide (TBHP), permethrin peroxide (C 21 H 20 Cl2O3), potassium persulfate (KHSO5), benzoyl peroxide (C 14 H 10 O4), sodium perborate (NaBO3), hydrogen peroxide (H2O2); in some embodiments, the first cross-linking reagent is hydrogen peroxide (H2O2).

[0097] In some preferred embodiments, the first leading coating includes 3 to 8 units / mL of the first cross-linking enzyme, and the second leading coating includes 0.01 to 0.05 wt% of the first cross-linking reagent; in some more preferred embodiments, the first leading coating includes 4 to 6 units / mL of the first cross-linking enzyme, and the second leading coating includes 0.01 to 0.03 wt% of the first cross-linking reagent.

[0098] In a tenth embodiment, the third precursor coating comprises 1 to 10 units / mL of a second cross-linking enzyme, such as horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxiredoxin, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase, or catechol oxidase. In some preferred embodiments, the second cross-linking enzyme is horseradish peroxidase (HRP).

[0099] In the tenth embodiment, the fourth precursor coating comprises 0.005 to 0.08 wt% of a second crosslinking agent such as a reactive oxygen species, wherein the reactive oxygen species include but are not limited to urea peroxide (CH6N2O3), peracetic acid (C2H4O3), sodium percarbonate (2Na2CO3·H2O2), di-tert-butyl peroxide (DTBP), benzoyl peroxide (C 14 H 10 O4), tert-butyl hydroperoxide (TBHP), permethrin peroxide (C 21 H 20 Cl2O3), potassium persulfate (KHSO5), benzoyl peroxide (C 14 H 10 O4), sodium perborate (NaBO3), hydrogen peroxide (H2O2); in some embodiments, the second cross-linking reagent is hydrogen peroxide (H2O2).

[0100] In some preferred embodiments, the third leading coating includes 3 to 8 units / mL of the second cross-linking enzyme, and the fourth leading coating includes 0.01 to 0.05 wt% of the second cross-linking reagent; in some more preferred embodiments, the third leading coating includes 4 to 6 units / mL of the first cross-linking enzyme, and the fourth leading coating includes 0.01 to 0.03 wt% of the first cross-linking reagent.

[0101] In the tenth embodiment, referring to FIG. 2C , the method for preparing the first biocompatible precursor polymer includes:

[0102] Step S11: mixing water and the first biocompatible polymer to obtain a first aqueous solution; and

[0103] Step S12: preparing a bridging aqueous solution and mixing it with the first aqueous solution to obtain the first biocompatible precursor.

[0104] Continuing to refer to FIG2C , the preparation method of the second biocompatible precursor polymer includes:

[0105] Step S13: mixing water and the second biocompatible polymer to obtain a second aqueous solution; and

[0106] Step S14: taking the bridging aqueous solution and mixing it with the second aqueous solution to obtain the second biocompatible precursor.

[0107] In these embodiments, the method for preparing the bridging aqueous solution comprises:

[0108] Step S10a: mixing a biphasic solvent and a bridging agent to obtain a bridging precursor solution, wherein the biphasic solvent includes water and an organic solvent; and

[0109] Step S10b: adding a curing agent to the bridging precursor solution to activate the bridging agent to obtain a bridging aqueous solution.

[0110] In these embodiments, the first aqueous solution is prepared by mixing 1 to 10 parts by weight of the first biocompatible polymer and 100 to 200 parts by weight of water. To maintain the fluidity of the first aqueous solution, the first aqueous solution is maintained at 35 to 45° C. before performing step S12. In a preferred embodiment, the first aqueous solution is prepared by mixing 3 to 7 parts by weight of the first biocompatible polymer and 120 to 180 parts by weight of water. More preferably, the first aqueous solution is prepared by mixing 4 to 6 parts by weight of the first biocompatible polymer and 140 to 160 parts by weight of water.

[0111] In these embodiments, the second aqueous solution is prepared by mixing 1 to 10 parts by weight of the second biocompatible polymer and 100 to 200 parts by weight of water. To maintain the fluidity of the second aqueous solution, the second aqueous solution is maintained at 35 to 45° C. before performing step S14. In a preferred embodiment, the second aqueous solution is prepared by mixing 3 to 7 parts by weight of the second biocompatible polymer and 120 to 180 parts by weight of water. More preferably, the second aqueous solution is prepared by mixing 4 to 6 parts by weight of the second biocompatible polymer and 140 to 160 parts by weight of water.

[0112] In these embodiments, water and an organic solvent are mixed in a volume ratio of 3:(1 to 10) to obtain the biphasic solvent, wherein the organic solvent can be exemplified by dimethylacetamide (DMAc), N-methyl-2-pyrrolidone (NMP), diethylformamide (DEF), 1,4-dioxane, propylene carbonate, hexamethylphosphoramide (HMPA), sulfolane, N-ethyl-2-pyrrolidone (NEP), or dimethylformamide (DMF). In preferred embodiments, water and an organic solvent are mixed in a volume ratio of 3:(2 to 5) to obtain the biphasic solvent. In these preferred embodiments, the organic solvent is dimethylformamide (DMF).

[0113] In these embodiments, the bridging water precursor solution contains 10 to 30 mmol of a bridging agent, and the bridging agent has one or more phenol functional groups, such as phenylacetic acid (PAA), 4-hydroxybenzoic acid (p-hydroxybenzoic acid), cinnamic acid, benzoic acid, tyrosine, salicylic acid, vanillic acid, 3,4-dihydroxybenzoic acid or 3-(4-hydroxyphenyl)propionic acid; in some preferred embodiments, the bridging precursor solution contains 15 to 25 mmol of a bridging agent; more preferably, the bridging precursor solution contains 18 to 22 mmol of a bridging agent, wherein the bridging agent is 3-(4-hydroxyphenyl)propionic acid.

[0114] In these embodiments, the bridging precursor solution is added with 10 to 30 mmol of a curing agent, the curing agent comprising a first curing agent and a second curing agent, wherein the first curing agent and the second curing agent are mixed in a weight ratio of 1: (0.5 to 2.0) to obtain the curing agent; the first curing agent can be exemplified by 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC), 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDCI), dicyclohexylcarbodiimide (DCC) or 1-ethyl-3-(3-dimethylaminopropyl)thiocarbodiimide (EDT); the second curing agent can be exemplified by N -hydroxysuccinimide (NHS), sulfo-NHS (N-hydroxysulfosuccinimide), sulfo-NHS-SS-biotin), 1-cyclohexyl-3-(2-morpholinoethyl)carbodiimide (CMC), BOP (benzotriazol-1-yloxytris(dimethylamino)phosphonium hexafluorophosphate); in a preferred embodiment, the first curing agent and the second curing agent are mixed in a weight ratio of 1: (1.0 to 1.5) to obtain the curing agent; in these embodiments, the first curing agent is 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride (EDC), and the second curing agent is N-hydroxysuccinimide (NHS).

[0115] In the tenth embodiment, the first biocompatible polymer, the second biocompatible polymer, the target cells and the acellular tissue layer (10) are specifically implemented as described in the first embodiment and are not described in detail here.

[0116] Please refer to FIG2D , which illustrates the eleventh embodiment of the present invention, wherein the specific steps and raw materials involved in the preparation method of the bionic tissue implant (100) are substantially the same as those of the tenth embodiment and are not described in detail here. However, before applying the step S3, the method includes:

[0117] Step S1': taking the decellularized tissue as the decellularized tissue layer (10); and

[0118] Step S2': mixing the first biocompatible coating and the target cells to obtain a first biocompatible cell mixture, and coating the first biocompatible cell mixture on one side of the decellularized tissue layer (10) to form the first bioscaffold layer (1).

[0119] Please refer to FIG2E , which illustrates the twelfth embodiment of the present invention, wherein the specific steps and raw materials involved in the preparation method of the bionic tissue implant (100) are substantially the same as those of the tenth embodiment and are not described in detail here. However, before step S3, it includes:

[0120] Step S1″: taking a decellularized tissue as a decellularized tissue layer (10);

[0121] Step S2a"-1: mixing the first precursor coating and the target cells to obtain a first precursor cell mixture;

[0122] Step S2a"-2: simultaneously or separately coating the first leading cell mixture and the second leading coating on one side of the decellularized tissue to obtain the first leading scaffold layer; and

[0123] Step S2b": contacting the first cross-linking enzyme with the first cross-linking reagent for the first time to cross-link the first biocompatible polymer with the acellular tissue layer (10), so that the first leading scaffold layer forms the first biological scaffold layer (1).

[0124] Please refer to FIG2F , which illustrates the twelfth embodiment of the present invention, wherein the specific steps and raw materials involved in the preparation method of the bionic tissue implant (100) are substantially the same as those of the eleventh embodiment and are not described in detail here, except that step S3 includes:

[0125] Step S3a"-1: mixing the third precursor coating and auxiliary cells to obtain a third precursor cell mixture;

[0126] Step S3a"-2: simultaneously or separately coating the third leading cell mixture and the fourth leading coating on the side of the decellularized tissue relative to the first bioscaffold layer (1) to obtain the second leading scaffold layer; and

[0127] Step S3b: contacting the second cross-linking enzyme with the second cross-linking reagent for the second time to cross-link the second biocompatible polymer with the acellular tissue layer (10), so that the second leading scaffold layer forms the second bioscaffold layer (2).

[0128] In the twelfth embodiment, the helper cells are the same as the target cells and are not described in detail herein.

[0129] In the thirteenth embodiment of the present invention, the specific steps and raw materials involved in the preparation method of the bionic tissue implant (100) are roughly the same as those in the twelfth embodiment and are not repeated here, except that the auxiliary cells are different from the target cells, wherein the auxiliary cells include but are not limited to mesenchymal stem cells, adipose tissue stem cells, endothelial progenitor cells, mesenchymal skeletal stem cells, synovial stem cells, osteoblasts, chondrocytes, and osteocytes; in other embodiments, the auxiliary cells can also be fibroblasts, glial cells, mesenchymal cells, embryonic stem cells, umbilical cord mesenchymal stem cells, bone marrow mesenchymal stem cells or lymphocytes; in some preferred embodiments, the auxiliary cells are mesenchymal stem cells, which can be derived from bone marrow, adipose tissue, umbilical cord tissue, umbilical cord blood, peripheral blood, dental pulp, amniotic fluid, amniotic membrane, synovium, and periosteum.

[0130] In a fourteenth embodiment of the present invention, a use of a biomimetic tissue implant (100) for preparing a composition for promoting tissue damage repair is provided, comprising taking a biomimetic tissue implant (100) as provided in the first to ninth embodiments, and implanting the biomimetic tissue implant (100) at a biological tissue damage site, wherein, when implanting the biomimetic tissue implant (100), a side of the first bioscaffold layer (1) corresponding to the decellularized tissue layer (10) contacts the biological tissue damage site, or a side of the second bioscaffold layer (2) corresponding to the decellularized tissue layer (10) contacts the biological tissue damage site.

[0131] In the fourteenth embodiment, the damaged biological tissue includes but is not limited to liver tissue, adipose tissue, skin tissue, cartilage tissue, bone tissue, myocardial tissue, pericardium or blood vessel wall.

[0132] The following specifically describes several embodiments and experimental examples to illustrate the technical effects achieved by the bionic tissue implant (100) provided by the present invention.

[0133] In vitro cell culture

[0134] All cell experiments involved in the present invention used human osteosarcoma cell line (MG-63) or rabbit synovial cell line (HIG-82); specifically, MG-63 cell line was cultured in HG-DMEM supplemented with 10% FBS and 1% antibiotic-antimycotic solution (100X Antibiotic-Antimycotic solution, hereinafter referred to as AA) and cultured at 37°C in an environment of 5% CO2; on the other hand, HIG-82 cell line was cultured in Ham's F-12 medium supplemented with 10% FBS and 1% AA in a 37°C incubator.

[0135] Mesenchymal stem cell isolation

[0136] The infrapatellar fat pad of the rabbit was removed, washed with PBS, minced, and immersed in low glucose Dulbecco's Modified Eagle Medium (LG-DMEM) (Gibco) containing 0.2% type I collagenase (Sigma-Aldrich). The cells were cultured with shaking at 37°C and 5% CO2 for 1 day. The cell suspension was centrifuged and the supernatant was removed to obtain primary infrapatellar fat pad-derived mesenchymal stem cells (IFPSCs) and cultured continuously. The culture medium was changed daily for the first 3 days to remove residual oil and impurities. The culture medium used for IFPSCs was LG-DMEM supplemented with 10% fetal bovine serum (FBS) (Gibco) and 1% AA (Simply) was added and cultured in a 37°C, 5% CO2 incubator (SCA-165D, ASTEC); when the cell confluence reached approximately 80%, IFPSCs were detached with 1X diluted trypsin and recorded as the first passage cells (Passage 1).

[0137] Preparation of acellular cartilage slices

[0138] Porcine femoral articular cartilage was obtained and cartilage discs were excised using a 10 mm diameter tissue punch and a scalpel. The discs were washed several times with PBS to remove oil and other debris. Subsequently, the discs were cut into cartilage sheets (CS) approximately 20 μm thick using a cryostat.

[0139] Next, decellularization was performed to prepare acellular cartilage sheets (ASCs). Specifically, the CS was immersed in 0.25% trypsin-EDTA (HIMEDIA) for 24 hours to loosen the dense structure of the cartilage, with the solution changed every 8 hours. For the next 4 hours, the cartilage sheets were transferred to 10 mM Tris-HCl (JTBaker) containing 50 units / mL DNase and 1 unit / mL RNase and then washed with enzyme-free 10 mM Tris-HCl for an additional 20 hours. Finally, the CS was immersed in 10% Triton X-100 (Sigma-Aldrich) for 24 hours, with the Triton X-100 solution replaced every 8 hours. After removing the Triton X-100 solution, the ACS was rinsed with PBS (Biomate) for 2 days to remove residual solution and degradation products. Finally, the ACS was freeze-dried for an additional 2 days and stored at room temperature.

[0140] Please see Figures 3a to 3f, which are macroscopic images of CS and ACS. As shown in Figures 3a and 3b, the cartilage disc cut from the porcine femoral cartilage is 10 mm in diameter, showing a white structure with uniform thickness, and the average thickness of the cartilage disc is 1 mm. As shown in Figure 3c, the CS was further cryosectioned to obtain a 20 μm CS. After decellularization, the gross appearance of the ACS was transparent and still a circle with a diameter of 10 mm. As shown in Figure 3d, decellularization will cause the ACS to only unfold in solution and cannot be successfully unraveled using tweezers. To solve this operational problem, the ACS was further freeze-dried in PBS for 2 days to obtain better operability, as shown in Figures 3e to 3f.

[0141] Biocompatibility testing-acellular cartilage slices

[0142] The following evaluation of the cytotoxicity of decellularized ASCs was conducted. All sample extractions were prepared according to ISO 10993-12 to assess whether the samples released biotoxic substances after 24 and 48 hours of extraction.

[0143] Figures 4a to 4d show the biocompatibility test results of the decellularized cartilage slices. As shown in Figure 4a, after the MG-63 cell line was treated with the 24-hour extract, ACS had a promoting effect on the cell activity of the MG-63 cell line compared with the control group, and the longer the culture days, the better the cell activity. When the culture reached the third and fifth days, there was a significant difference in cell activity between the ACS and control groups. As shown in Figure 4b, the improvement effect of the 48-hour extract on the cell activity of the MG-63 cell line was not as significant as that of the 24-hour extract, but there was no negative effect of the 48-hour extract on cell activity, and its cell activity was still better than that of the control group.

[0144] As shown in Figure 4c, compared with the MG-63 cell line, the HIG-82 cell line showed a relatively high cell activity on the first day after treatment with the 24-hour extract, and still showed a higher average level than the control group on the 3rd and 5th days, and both showed a significant improvement in cell activity; as shown in Figure 4d, the 48-hour extract treatment group had a similar trend to the 24-hour extract treatment group.

[0145] Although the HIG-82 cell line and the MG-63 cell line showed different trends in the cytotoxicity test, their cell activities were above the minimum acceptable standard line (75% of the control group cell activity) at the three time points; the results showed that the ASCs prepared by the above method showed good biocompatibility with the different cell lines implanted thereon.

[0146] Preparation of precursor hydrogel

[0147] 5 g of gelatin (Sigma-Aldrich) was dissolved in 150 mL of double-distilled water (DIW) and stirred at 40°C for 1 hour. 3.32 g of HPA (Thermo Scientific) was placed in a biphasic solvent mixture of DIW and DMF (JTBaker) with a volume ratio of 3:2 and completely dissolved. Subsequently, 10 mmol of EDC (TCI) and 13.9 mmol of NHS (Sigma-Aldrich) were added to the HPA solution for 1 hour to activate the functional groups of HPA. The activated HPA solution was poured into the preheated gelatin solution and stirred at 40°C in the dark for 24 hours to allow cross-linking. After cross-linking, the gelatin-based hydrogel with hydroxyphenyl groups was transferred to a dialysis membrane and placed on DIW for 3 to 4 days. Finally, the product was centrifuged to obtain the supernatant, which was freeze-dried for 2 days to obtain the hydrogel GAHPA.

[0148] In the following examples, HRP (Sigma-Aldrich) and H2O2 (Panreac, Spain) were added to the hydrogel GAHPA to prepare HRP-precursor hydrogel and H2O2-precursor hydrogel, respectively.

[0149] As shown in Figures 5a to 5d, the ultraviolet-visible spectrophotometer (UV-vis), FT-IR and 1 H-NMR was used to examine the synthesized hydrogel to confirm that the hydroxyphenyl groups of HPA were indeed grafted onto the gelatin backbone. As shown in Figure 5a, the characteristic peak of HPA appears at a wavelength of 275 nm, which is significantly different from that of pure gelatin. The absorbance of the hydrogel GAHPA is 9.04943 au, while that of gelatin is 1.23425 au. The characteristic peak of the hydrogel GAHPA is at 275 nm, while the curve of pure gelatin at the same wavelength of 275 nm is flat.

[0150] As shown in Figure 5b, in the FTIR spectrum, the transmittance trend of the hydrogel GAHPA is roughly consistent with that of pure gelatin, which is because gelatin is the backbone of the hydrogel; among them, amide bonds appear at 1650, 1535, and 1190 cm -1 However, at 3400cm -1 The characteristic peak of HPA hydroxyphenyl at 3300 cm -1 There is an overlapping trend with pure gelatin.

[0151] Therefore, further through 1 H-NMR analysis of the structural differences between hydrogel GAHPA and gelatin; as shown in Figures 5c to 5d, 1The highest peak shown in the H-NMR graph is the background solution D2O. Among them, the chemical shift of the hydrogel GAHPA is significantly different from that of pure gelatin. In particular, the hydrogel GAHPA shows significant characteristic peaks at 6.8ppm and 7.1ppm, which can be attributed to the aromatic -OH group. It also shows that the hydroxyphenyl group of HPA is cross-linked with the gelatin main chain, and the hydrogel GAHPA is indeed successfully synthesized.

[0152] Example 1

[0153] The HRP-precursor hydrogel and the H2O2-precursor hydrogel were respectively injected into a Teflon mold with a diameter of 7 mm and a height of 4 mm, and the mixture of the HRP-precursor hydrogel and the H2O2-precursor hydrogel was allowed to stand for 5 minutes to obtain the hydrogel GAHPA; in Example 1, the HRP-precursor hydrogel contained 5 units / mL of HRP, and the H2O2-precursor hydrogel contained 0.01 wt% of H2O2.

[0154] Example 2

[0155] The preparation method of the hydrogel GAHPA was the same as that in Example 1, except that the H2O2-precursor hydrogel contained 0.03 wt% H2O2.

[0156] Example 3

[0157] The preparation method of the hydrogel GAHPA was the same as that in Example 1, except that the H2O2-precursor hydrogel contained 0.03 wt% H2O2.

[0158] Appearance Observation-Hydrogel GAHPA

[0159] FIG6 shows the appearance and microstructure images of the hydrogels provided in Examples 1 to 3, wherein Example 1 is shown in FIG6(a), (d) and (g), Example 2 is shown in FIG6(b), (e) and (h), and Example 3 is shown in FIG6(c), (f) and (i); As shown in FIG6(a) to (c), the biomimetic matrix gels with three different H2O2 concentrations are cylindrical in shape with a diameter of 7 mm and a height of 4 mm, and their overall appearance is similar; As shown in FIG6(e) and (h), Example 2 is larger than Example 1. and 3 have denser and deeper pores, and the distribution of these pores is more regular; as shown in Figures 6(d) and (g), the micromorphology of Example 1 also has a porous structure, but the pores are superficial and the pore walls are rough and thick; as shown in Figures 6(f) and (i), similar to Example 1, the microstructure of Example 3 is porous, but the depth of the pores is insufficient for nutrient transport and cell migration, and the distribution and size of its porous structure are relatively uneven; all three groups of Examples 1 to 3 can be effectively gelled within 5 minutes, among which Example 2 has a pore structure suitable for cell growth.

[0160] Swelling ratio test-hydrogel GAHPA

[0161] Figures 7a and 7b are line graphs showing the swelling ratio test results of Examples 1 to 3. Examples 1 to 3 were immersed in PBS for 0.5, 1, 24, and 48 hours and weighed to evaluate their swelling ratios. As shown in Figures 7a and 7b and Table 1, the swelling ratios of Examples 1 to 3 did not show any different trends. After swelling for 1 hour, the values ​​reached 450% to 550%, indicating that all groups rapidly absorbed PBS within 1 hour. Thereafter, after immersion for 24 and 48 hours, the swelling ratios of all groups remained almost unchanged, indicating that Examples 1 to 3 reached swelling equilibrium within 1 hour. Furthermore, their volumes did not undergo significant deformation and remained almost identical to their original forms.

[0162] Table 1

[0163] In vitro degradation test-hydrogel GAHPA

[0164] Figures 7c to 7d are line graphs showing the in vitro degradation test results of Examples 1 to 3; Examples 1 to 3 were immersed in PBS with or without lysozyme to estimate their mass loss at 1, 3, 5, 7, 14, 21 and 28 days, and the pH value was recorded once a week during the degradation test; as shown in Figure 7c, from the pH value trajectory, Examples 1 to 3 in PBS without lysozyme ranged from 6.7 to 7.0, which presented a weakly acidic environment that did not affect cell growth; as shown in Figure 7d, Examples 1 to 3 also showed a similar stable pH trend in PBS with lysozyme, from 6.7 to 7.0).

[0165] Mass Loss Assessment - Hydrogel GAHPA

[0166] Figures 7e to 7f are line graphs showing the mass loss evaluation results of Examples 1 to 3. As shown in Figure 7e, the mass losses of Examples 1 to 3 were all below 40% during the first two weeks of immersion in PBS without lysozyme. However, the mass loss of Example 1 began to increase significantly from the 21st day and continued until the 28th day. On the other hand, as shown in Figure 7f, the mass losses of Examples 1 to 3 in PBS containing lysozyme showed similar trends within the first week, all below 40%. However, the mass losses of Examples 2 and 3 reached 50% on the 14th day and dropped to around 20% on the 21st day, similar to the test results without lysozyme. The mass loss trend of Example 1 in PBS containing lysozyme began to increase significantly from the 21st day, which was the same as the result in PBS without lysozyme.

[0167] Protein adsorption test-hydrogel GAHPA

[0168] FIG8 is a histogram showing the protein adsorption test results of Examples 1 to 3. To infer the adhesion of proteins to the GAHPA hydrogel, protein adsorption tests were performed using a MicroBCA assay kit, and a series of FBS concentrations were used as a control group. As shown in FIG8 , the protein concentrations for Examples 1, 2, and 3 were 110.5±13.59 μg / mL, 94.47±7.68 μg / mL, and 96.86±11.85 μg / mL, respectively. There were no significant differences between the groups, indicating that high or low concentrations of H2O2 did not affect the protein's adhesion ability.

[0169] Rheological properties test-hydrogel GAHPA

[0170] Figures 9a to 9b are line graphs showing the rheological property test results of Examples 1 to 3. In cartilage tissue engineering, the injectability and gelation ability of hydrogels are very important. A rheometer was used to test samples within a certain shear rate range to verify whether the hydrogels were injectable. As shown in Figure 9a, Examples 1 to 3 all showed a clear peak at a shear rate of 0.01 / s. From the lowest H2O2 concentration to the highest concentration, the precise viscosity values ​​were 122.71, 151.35, and 290.15 Pa·s, respectively. As the shear rate increased, the viscosity of each group gradually decreased, which can be attributed to the shear thinning property, and also showed that Examples 1 to 3 had very good rheological properties, which was conducive to injection.

[0171] On the other hand, the determination of storage modulus (G') and loss modulus (G") can also be used to confirm that the precursor hydrogels with different H2O2 concentrations can properly induce gelation; the rheometer tested the hydrogel GAHPA in oscillation mode; as shown in Figure 9b, the storage modulus (G') of Examples 1 to 3 is much higher than the loss modulus (G") (Figure 3.9b), indicating that the cross-linked structures of Examples 1 to 3 are good and can be used to manufacture biomimetic tissue implants (100).

[0172] Biocompatibility Testing-Hydrogel GAHPA

[0173] In this test, the samples were prepared according to ISO 10993-12 and extracted for 24 and 48 hours respectively.

[0174] As shown in Figure 10a, after the MG-63 cell line was treated with the extract for 24 hours, the cell viability on days 1, 3, and 5 was higher than that of the control group. Example 1 had the best survival rate on day 3 and showed two significant differences from the control group and Example 3. The cell viability of each group decreased slightly on day 5, but was still better than that on day 1. In addition, as shown in Figure 10b, after the MG-63 cell line was treated with the extract for 48 hours, Example 1 also showed two significant differences from the control group and Example 3 on day 1. Among them, Example 3 did not exceed the minimum acceptable cell viability standard on days 3 and 5. In addition, on day 3, the cell viability of Example 2 was almost lower than 75% of the cell viability of the control group, but recovered to above the standard on day 5.

[0175] As shown in Figure 10c, after the HIG-82 cell line was treated with the extract for 24 hours, Examples 1 and 2 were mild to the cells, and the cell viability reached 250% on the 5th day after treatment. On the other hand, the results of Example 3 were relatively unstable, and its biological activity was unexpectedly reduced on the 1st day after treatment. Secondly, the standard deviation showed a significant difference on the 5th day, which means that the cells could not grow stably in this environment. On the other hand, as shown in Figure 10d, after the HIG-82 cell line was treated with the extract for 48 hours, all groups showed good cell viability on the 1st day, but the cell viability of Example 3 decreased significantly on the 3rd and 5th days. The biocompatibility test shows that Examples 1 and 2 have the effect of promoting the cell viability of MG-63 cell lines and HIG-82 cell lines.

[0176] Cell Adhesion Test-GAHPA Hydrogel

[0177] Plant 1×10 4 HIG-82 cells were plated onto the hydrogels of Examples 1, 2, and 3 and cultured for 1, 3, and 5 days to test the cell adhesion ability of the hydrogel GAHPA and observe its effect on cell morphology. Healthy chondrocytes should be round in shape and contain some synthetic ECM, with an average diameter of approximately 5 to 10 μm. In addition, it is preferred that multiple cells aggregate to promote cell attachment and proliferation.

[0178] As shown in Figures 11(a) and 11(d), the cells of Example 1 on the 1st and 3rd days were round, but slightly smaller in size; as shown in Figure 11(g), the cells of Example 1 cultured for 5 days were larger and had a smooth cell surface. The cells were spindle-shaped, and the pseudopodia extended by the interstitial cells were attached to the hydrogel GAHPA; the cells of Example 2 were round on days 1, 3, and 5, and the roundness was more pronounced on day 3.

[0179] On the other hand, the ECM covered on the hydrogel GAHPA was further observed; on the first day after the cells were seeded in Example 2, as shown in Figures 11(b) and 11(e), the cells synthesized some ECM around themselves, but not as much as the cells on the third day; on the other hand, as shown in Figure 11(c), the cells attached on the first day in Example 3 remained round with a diameter of about 10 μm; as shown in Figure 11(i), after culturing for 3 days, the cells aggregated and synthesized an ECM matrix around them; as shown in Figure 11(f), the cells still aggregated on the fifth day, but their morphology changed as the culture time was prolonged, with some cells growing into irregular shapes and some cells aggregated and becoming round, but smaller in size than the cells on the third day.

[0180] Preparation of biomimetic tissue implant (100)

[0181] Example 4

[0182] Please refer to FIG12 . HRP-precursor hydrogel and H2O2-precursor hydrogel are injected into Teflon molds respectively, and ACS is covered on the mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel for conjugation. HRP-precursor hydrogel and H2O2-precursor hydrogel are again spread on ACS to complete the biomimetic tissue implant (100). In Example 1, HRP-precursor hydrogel contains 5 units / mL of HRP, and H2O2-precursor hydrogel contains 0.01 wt% of H2O2. 2- .

[0183] Example 5

[0184] The preparation method of the bionic tissue implant (100) is the same as that of Example 4, except that the H2O2-precursor hydrogel contains 0.03 wt% of H2O2.

[0185] Example 6

[0186] The preparation method of the bionic tissue implant (100) is the same as that of Example 4, except that the H2O2-precursor hydrogel contains 0.05 wt% of H2O2.

[0187] Example 7

[0188] The preparation method of the bionic tissue implant (100) is the same as that of Example 4, except that two sheets of ACS are superimposed on each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0189] Example 8

[0190] The preparation method of the bionic tissue implant (100) is the same as that of Example 5, except that two sheets of ACS are superimposed on each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0191] Example 9

[0192] The preparation method of the bionic tissue implant (100) is the same as that of Example 6, except that two sheets of ACS are superimposed on each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0193] Example 10

[0194] The preparation method of the bionic tissue implant (100) is the same as that of Example 4, except that three sheets of ACS are stacked on top of each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0195] Example 11

[0196] The preparation method of the bionic tissue implant (100) is the same as that of Example 5, except that three ACS sheets are stacked on top of each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0197] Example 12

[0198] The preparation method of the bionic tissue implant (100) is the same as that of Example 6, except that three ACS sheets are stacked on top of each other and then covered with a mixture of HRP-precursor hydrogel and H2O2-precursor hydrogel.

[0199] Experimental Example 1 - Mechanical Properties Test - Adhesion Strength

[0200] The bonding strength of the hydrogel GAHPA in Examples 1 to 6 was tested using a material testing system (MTS).

[0201] As shown in FIG13a , the hydrogel GAHPA containing ACS is superior to the pure hydrogel GAHPA, while the bonding strength of the pure hydrogel GAHPA of Examples 1 to 3 is not much different, and the bonding strength value is about 0.35 kPa; as shown in Table 2, among the sandwich model scaffolds of Examples 4 to 6, the bonding strength of Example 6 reaches 0.62 kPa; comparing the bionic tissue implants (100) containing and not containing ACS, the bonding strength is more significant with the increase of H2O2 concentration; as shown in FIG13a , except for Example 5, Example 6 achieved statistically significant differences among all groups (**p<0.005).

[0202] Table 2

[0203] Experimental Example 2 - Mechanical Properties Test - Compression Modulus

[0204] The compression modulus of Examples 7 to 9 was measured using an ARES G-2 dynamic mechanical analyzer (DMA).

[0205] As shown in Table 3, the compression modulus was calculated within the 15-30% linear range of the stress-strain curve. In the pure hydrogel groups (Examples 1 to 3), the average compression modulus of Example 3 was 18.9 kPa, higher than that of Example 2 (17.61 kPa) and Example 1 (13.66 kPa). Among the sandwich model scaffold groups (Examples 7 to 9), Example 8 had the highest compression modulus (26.07 kPa), while Examples 7 and 9 had values ​​of 19.36 kPa and 20.48 kPa, respectively.

[0206] Table 3

[0207] As shown in FIG13b , it is obvious that the sandwich scaffold model containing ACS can withstand a larger compressive load than the pure hydrogel without ACS. Among them, Example 8 has the best structural properties and is significantly different from Examples 1 and 2 (*p<0.05, ***p<0.001).

[0208] Experimental Example 3-Appearance Observation

[0209] In Examples 10 to 12, the hydrogel GAHPA was cross-linked for 5 minutes before demolding. As shown in (a) to (c) of Figure 14, Examples 10 to 12 were successfully gelled within 5 minutes, with 3 ACS inside. The ACS was located in the middle of the scaffold as shown by the red arrows, and multiple bubbles were formed around it.

[0210] Further observation and evaluation of the microstructure revealed that ACSs were cross-linked with the hydrogel GAHPA in all groups. As shown in Figures 14(d) and (g), the ACS in Example 10 was completely cross-linked and bridged with the hydrogel, and the exact location of the ACS was not easy to identify. The red arrow indicates the presumed location of the ACS, which was abundantly covered by the hydrogel. As shown in Figures 14(e) and (h), the morphology of Example 11 had some obvious gaps between the ACS layers. These gaps were bubbles formed when the two precursor hydrogels were mixed, and the ACS was covered by the porous hydrogel and completely cross-linked and bridged. As shown in Figures 14(f) and (i), the ACS in Example 12 was also tightly cross-linked with the hydrogel. However, due to the higher concentration of H2O2, the gelation rate was faster. The gap size of the implant in Example 11 was significantly larger than that in the other groups, but the porous structure of the hydrogel was not as obvious as that in the implant in Example 11.

[0211] Example 13

[0212] The preparation method of the biomimetic tissue implant (100) is the same as that of Example 4, except that 6x10 5 Synovial cells HIG82 on ACS.

[0213] Example 14

[0214] The preparation method of the biomimetic tissue implant (100) is the same as that of Example 5, except that before laying the HRP-precursor hydrogel and H2O2-precursor hydrogel on the ACS, 6x10 5 Synovial cells HIG82 on ACS.

[0215] Example 15

[0216] The preparation method of the biomimetic tissue implant (100) is the same as that of Example 6, except that before laying the HRP-precursor hydrogel and H2O2-precursor hydrogel on the ACS, 6x10 5 Synovial cells HIG82 on ACS.

[0217] Experimental Example 4-Histological Staining

[0218] In order to evaluate the cell phenotype and growth status in the biomimetic tissue implant (100), cells were encapsulated into HRP precursor hydrogel and a sandwich model scaffold with a single layer of ACS was fabricated; after 1 day and 5 days of culture, HE (Hematoxylin and Eosin), SO (Safranin-O staining), MT (Masson's Trichrome staining) staining and Live / Dead detection were used, respectively, and the internal structure of the sandwich model scaffold was observed from the top and cross section.

[0219] HE tissue staining

[0220] As shown in Figures 15(a) to (i), HE staining showed that the ACS in Examples 13 to 15 were all stained pink or light red, indicating that the ACS in these implants had a rich matrix; chondrocytes were stained dark purple, and the hydrogel was stained purple; after one day of culture, as shown in the cross-sections of Figures 15(a), (e), and (i) and the top views of Figures 15(b), (f), and (j), the cells in all examples were evenly and independently distributed in the hydrogel, and the cells did not settle to the bottom of the ACS or the hydrogel.

[0221] The culture was continued until the 5th day. As shown in Figures 15(c) and (d), only the cells of Example 13 showed uniform distribution. However, as shown in Figures 15(h) and (l), from the top view (Figures 15h and 15l), Examples 14 and 15 showed multiple cell aggregates, and the cell distribution was not as large as on the 1st day.

[0222] In addition, as shown in Figure 15(k), the ACS in the cross-section of Example 15 could not be observed from the image because bubbles formed between the ACS and the hydrogel, eventually causing the ACS to rupture during paraffin sectioning; in addition, as shown in the staining results in Figure 15(i), the rapid gelation made it challenging to stack a flat and smooth ACS in the sandwich model scaffold; as shown in Figure 15(h), it was not easy to place the ACS in the center, and the ACS was exposed at the edge of the implant.

[0223] SO tissue staining

[0224] The ECM synthesized by the chondrocyte-encapsulated implant samples can be observed by SO staining. As shown in FIG16 , the ACS in Examples 13 to 15 all appeared light blue and almost transparent, while the chondrocytes appeared dark red or brown. In addition, the hydrogel was stained light red with some dark red areas. On the first day of culture, as shown in FIG16( a ), Example 13 showed that the cells aggregated near the ACS. As shown in FIG16( b ), the chondrocytes were still evenly distributed from the top view. The cell aggregation in Example 13 was caused by the partial rupture of the bottom of the scaffold, which forced the cells to aggregate in a limited area. The cells in Example 14 ( FIG16 ( e ) and ( f )) and Example 15 ( FIG16 ( i ) and ( j )) were evenly spread.

[0225] On the fifth day of culture, as shown in Figures 16(c) and (d), the cells of Example 13 proliferated evenly; as shown in Figures 16(g) and (h), cell clusters were generated in Example 14 and also diffused regularly in the hydrogel; as shown in Figure 16(k), the cells in Example 15 were evenly distributed, and as shown in Figure 16(l), cell clumps were observed on the top of the implant; as shown in Figure 16(k), similar to HE staining, SO staining still could not observe ACS in the cross-sectional view of Example 15, because when mixing the HRP-precursor hydrogel and the H2O2-precursor hydrogel, bubbles between the ACS and the hydrogel eventually caused the ACS to rupture during the subsequent paraffin sectioning process.

[0226] MT tissue staining

[0227] As shown in Figure 17 , in MT staining, the ACS stained bright blue, the hydrogel stained light pink, and the chondrocytes stained dark red. Starting from the first day of culture, as shown in Figure 17(a), the cells of Example 13 aggregated near the ACS, and from a top view, as shown in Figure 17(b), the cells were uniformly encapsulated in the implant (Figure 17b). As speculated by the SO staining, the cells aggregated due to the rupture of the ACS. On the other hand, the cells of Example 14 (Figures 17(e) and (f)) and Example 15 (Figures 17(i) and (j)) were evenly distributed. On the fifth day of culture, as shown in Figures 17(c) and (d), fewer cells were stained in Example 13, while as shown in Figures 17(k) and (l) and in Example 15, uniform distribution of cells was observed, but no cell clusters were generated. As shown in Figure 17(g), only Example 14 produced multiple cell clusters during culture, with multiple cells dispersed.

[0228] Experimental Example 5-Live / Dead Staining

[0229] To verify the survival of cells in the sandwich model scaffold, Live / Dead assays were performed after 1 and 5 days of culture. Live cells and dead cells were stained with green and red fluorescence using calcein-AM and EthD-1, respectively. The survival status of cells in the samples was also observed using cross-sectional and top-down views.

[0230] As shown in Figures 18 to 19, according to the live cell and dead cell detection results on the first day, multiple cells can be seen as green bright spots, representing the distribution of live cells in the sandwich model scaffold; in the cross-section and top view of calcein-AM (Figures 18(a), (d) and (g) and Figures 19(a), (d) and (g)), the cells show many fluorescent green spots, indicating that after 1 day of culture, the cells can survive and expand evenly in the sandwich model scaffold.

[0231] On the first day of EthD-1 staining, almost no bright red spots were observed (Figures 18(b), (e) and (h) and Figures 19(b), (e) and (h)), which also means that there was almost no cell death. Comparing Examples 13 to 15, living cells were more obvious than dead cells, and only sparse red spots were present in the images of Examples 13 and 15.

[0232] After 5 days of culture, as shown in Figures 20 to 21, the green cells were less abundant and less uniform compared with the results on the first day; nevertheless, in the three experimental groups, both from the cross-section (Figures 20(a), (d) and (g) compared with Figures 20(b), (e) and (h)) and the top view (Figures 21(a), (d) and (h) compared with Figures 21(b), (e) and (h)), the living cells were still clearer than the dead cells; in addition, as shown in Figure 21(d), some cell clusters were visible in the top view of Example 14; as shown in Figures 20(g) and 21(g), some cells were also aggregated in Example 15; in addition, almost no red fluorescence with detectable intensity was found in the cross-section and top views of Figures 20 to 21, which represents the biocompatibility of the sandwich scaffold model provided by Examples 13 to 15.

[0233] To confirm that the absence of EthD-1 red fluorescence in these fluorescence staining images was not due to low fluorescence intensity detected by conjugate microscopy, the distribution of live cells in the brightest image of each group was calculated using ImageJ. As shown in Table 4 and Figure 22, the live cell distribution results showed that the live cell percentage in each group at different time points was above 80%, indicating that the absence of red fluorescence was indeed due to the inability to detect dead cells.

[0234] Table 4

[0235] Example 16

[0236] Mix 6.5×10 5 HIG-82 cells were mixed and injected into Teflon molds with H2O2-precursor hydrogels respectively, and allowed to stand for 5 minutes to allow the hydrogels to crosslink and bridge. In Example 16, the HRP-precursor hydrogel contained 5 units / mL of HRP, and the H2O2-precursor hydrogel contained 0.01 wt% of H2O 2- .

[0237] Example 17

[0238] The preparation method of the bionic tissue implant (100) is the same as that of Example 16, except that the H2O2-precursor hydrogel contains 0.03 wt% of H2O2.

[0239] Example 18

[0240] The preparation method of the bionic tissue implant (100) is the same as that of Example 16, except that the H2O2-precursor hydrogel contains 0.05 wt% of H2O2.

[0241] Experimental Example 6-Quantification of Type II Collagen

[0242] Type II collagen is the main component of the cartilage ECM, accounting for 90% to 95% of the total collagen, and establishes a network to stabilize the cartilage tissue. To estimate the type II collagen synthesized by chondrocytes in the sandwich model scaffold, the type II collagen content of cells encapsulated in pure GAHPA hydrogels (Examples 16 to 18) and single-layer sandwich model scaffolds (Examples 13 to 15) was tested after 1 day and 5 days of culture.

[0243] Figures 23 and 24 respectively present the results of the quantification of type II collagen in pure GAHPA hydrogels (Examples 16 to 18) and single-layer sandwich model scaffolds (Examples 13 to 15); as shown in Figure 23, there was no significant difference in the concentration of type II collagen in Examples 16 to 18; on day 1, the type II collagen concentrations of Examples 16 to 18 decreased to 17.55 to 18.8 ng / mL, and on day 5, the values ​​of Examples 16 and 18 increased slightly to approximately 20.39 and 21.81 ng / mL; only the type II collagen concentration of Example 17 decreased to 14.19 ng / mL, but there was no statistical difference between the other groups.

[0244] As shown in Figure 24, the type II collagen concentrations of Examples 13 to 15 were much higher than those of Examples 16 to 18. On day 1, the type II collagen of Example 13 was slightly less than that of the other two groups. However, after 5 days of culture, the type II collagen production of the cells seeded in Example 14 was the highest, reaching 11.17 μg / mL. The type II collagen levels of Examples 13 and 15 were quantified at 10.94 and 9.73 μg / mL, respectively.

[0245] Example 19

[0246] The HRP-precursor hydrogel and the H2O2-precursor hydrogel were respectively injected into a Teflon mold and allowed to stand for cross-linking and bridging to complete the first layer of hydrogel GAHPA; the above steps were repeated 10 times on the first layer of hydrogel GAHPA to complete the bionic tissue implant (100); in Example 19, the HRP-precursor hydrogel contained 5 units / mL of HRP, and the H2O2-precursor hydrogel contained 0.03 wt% of H2O2.

[0247] Example 20

[0248] Please refer to Figure 25. HRP-precursor hydrogel and H2O2-precursor hydrogel are injected into the Teflon mold respectively, and ACS is covered on the first layer of the precursor hydrogel mixture for conjugation; HRP-precursor hydrogel and H2O2-precursor hydrogel are again laid on the ACS to form a second layer of the precursor hydrogel mixture; then a second piece of ACS is covered on the second layer of the precursor hydrogel mixture, and HRP-precursor hydrogel and H2O2-precursor hydrogel are again laid on the second piece of ACS, and the steps after the second piece of ACS are repeated until the tenth piece of ACS is covered on the tenth layer of the precursor hydrogel mixture, and the eleventh layer of the precursor hydrogel mixture is laid on the tenth piece of AC to complete the biomimetic tissue implant (100); in Example 19, the HRP-precursor hydrogel contains 5 units / mL of HRP, the H2O2-precursor hydrogel contains 0.03 wt% of H2O2, and the biomimetic tissue implant (100) is marked as implant GAHPA-0.03 (10).

[0249] Example 21

[0250] The preparation method of the biomimetic tissue implant (100) is the same as that of Example 20, except that 6x10 5 Primary infrapatellar fat pad-derived mesenchymal stem cells (IFPSCs) are cultured and the biomimetic tissue implant (100) is labeled as implant IFPSCs.

[0251] Morphological assessment

[0252] The macrostructures of Examples 19 and 20 were captured using a digital camera, and their microstructures were evaluated using SEM after freeze-drying. As shown in Figures 26A and 26B, Examples 19 and 20 were prepared in a Teflon mold with a diameter of 7 mm and a height of 4 mm. After static cross-linking, both Examples 19 and 20 maintained a cylindrical appearance.

[0253] As shown in Figures 26C and 26D , SEM images show that Example 19 presents a pure hydrogel with a three-dimensional porous network structure, while Example 20 is a composite sandwich scaffold whose microstructure is composed of multiple parallel ACSs. ​​Due to the smaller volume of the hydrogel between the ACS layers in Example 20, some obvious gaps are generated in the microstructure after freeze-drying.

[0254] However, as shown in Figure 26E , at a higher magnification, a tight cross-linking bridge structure (red arrow) can be seen between the hydrogel of Example 19 and the upper and lower ACS layers. As previously mentioned, after enzyme catalysis, the phenolic groups on the ACS surface co-crosslink with the gelatin hydrogel. Freeze-drying SEM imaging may cause the hydrogel structure between the ACS to be pulled and destroyed, and the original state of the scaffold cannot be directly observed. However, it can still be determined that the appearance of these composite scaffolds and subsequent compression tests are stable. In other words, this sandwich composite scaffold with a stacked cross-linking pattern has a complementary effect of co-crosslinking and mutual fusion.

[0255] Mechanical properties testing

[0256] The compression modulus was calculated within the linear range of about 20-35% of the stress-strain curve, and the statistical values ​​are listed in Table 5. The average compression modulus of Example 19 was 25.35 kPa, while the average value of Example 20 reached 72.43 kPa. As shown in Figure 27, according to statistics, the mechanical properties of Example 20 stacked with ACS were significantly improved. In other words, based on the mechanical property evaluation of the previous Examples 7 to 9 and 13, it can be seen that the addition of ACS can improve the mechanical properties lacking in the hydrogel in the bionic tissue implant (100), and in the future, users can adjust the number of layers of hydrogel and ACS stacking according to needs to obtain the desired mechanical properties.

[0257] Table 5

[0258] Cell compatibility testing

[0259] The cell viability test is to ensure the non-toxicity of the material. The final extract is extracted according to ISO10993-12 and the extraction time is 24 hours. This test uses IFPSCs that have been passaged 4 times, and the experimental groups are Example 19 and Example 20.

[0260] As shown in Figure 28, on the first day of culture, both experimental groups showed satisfactory cell activity. Although the cell activity showed a downward trend on the third day, it recovered significantly on the fifth day. Overall, both experimental groups performed well in this test and the activity percentage was still higher than the cell activity standard (75% of the control group), indicating that Example 19 and Example 20 are both suitable for the growth and proliferation of IFPSCs.

[0261] Experimental Example 5-Animal Experiment

[0262] Construction of Invivo New Zealand white rabbit osteoarthritis animal model

[0263] All animal experimental procedures were performed at the Laboratory Animal Center of the College of Medicine, National Cheng Kung University; the experimental protocols were reviewed and approved by the Institutional Animal Care and Use Committee (IACUC) with approval number 110243.

[0264] In Experimental Example 5, 24 New Zealand white rabbits from the Tainan Animal Husbandry Research Institute were used, with an average weight ranging from 2.5 to 3.0 kg. As shown in FIG29(A), the New Zealand white rabbits were divided into the following four groups: a control group, a hydrogel group (Gel), a hydrogel + ACS group (Gel + ACS), and a hydrogel + ACS + IFPSCs group (Gel + ACS + Cell).

[0265] Before surgery, general anesthesia and sedation were performed with 5 mg / kg of alfaxalone (Jurox) and 2 mg / kg of xylazine (Bayer). Postoperatively, anesthesia was maintained by intubation with 1.5 to 5% isoflurane (Panion). The superficial hair of the knee skin was shaved and disinfected. The patella was laterally dislocated, and the skin and muscle were incised to expose the patellofemoral groove of the knee joint. A portable electric drill (CANTA) was used to drill a circular full-thickness osteochondral defect in the weight-bearing area of ​​both knees. The defect had a diameter of 3 mm and a depth of 3 mm. As shown in FIG29(C), the defects were subsequently treated or not treated according to the aforementioned experimental groups. The control group received no treatment, the Gel group was treated with Example 19, the Gel+ACS group was treated with Example 20 containing 10 layers of ACS, and the Gel+ACS+Cell group was treated with 6×10 5 The method for preparing the biomimetic tissue implant (100) of Example 21 is shown in FIG29(B).

[0266] Postoperatively, the patella was repositioned, the joint capsule was sutured with absorbable Vicryl 4-0 sutures (SH-2 4-0, ETHICON), and the skin was sutured with nylon 4-0 sutures (NC194L 4-0) (UNIK). During all surgical procedures, the New Zealand white rabbits' body temperature was maintained with a warming pad until they were awake and able to stand independently. Three days after surgery, the rabbits were injected with the antibiotic enrofloxacin (Livisto) 5 mg / kg and the analgesic nalbuphine (Yunxin) 1.5 mg / kg to prevent infection and postoperative pain. Four and 12 weeks after surgery, the animals were sacrificed, and both knee joints were removed for subsequent analysis.

[0267] Macromorphological Assessment

[0268] The New Zealand white rabbits were sacrificed 4 and 12 weeks after surgery. The knee joint specimens from both sides were cleaned and photographed, and then evaluated by at least three technicians using the Wayne scoring system (Table 6). The Wayne scoring system is a modification of the International Cartilage Repair Society (ICRS) visual histology scale, with a maximum score of 12. Three technicians performed blind testing on the same specimen, and the three scores obtained were averaged and rounded to one decimal place to form the final score, which was then plotted.

[0269] Table 6

[0270] In Experimental Example 5, the repair ability of the composite sandwich scaffold used in the New Zealand white rabbit cartilage defect model was evaluated. First, the overall observation performance was evaluated. As shown in Figure 30, at 4 weeks after surgery, the defects in each group were still obvious, and the surface still showed vacant or rough new tissue. Although the difference in appearance was not significant, the Gel+ACS group or Gel+ACS+Cell group using the composite scaffold was significantly filled with more repair tissue than the Gel group using only hydrogel. Obviously, the repair performance of the Gel+ACS+Cell group was better than that of the other groups, but the surface of the repair tissue was still rough.

[0271] Twelve weeks after surgery, the control group still showed incomplete tissue coverage, while the Gel group, Gel+ACS group, or Gel+ACS+Cell group was filled with new cartilage with a color similar to that of the host tissue; all experimental groups showed a more complete and smoother tissue appearance 12 weeks after surgery than 4 weeks after surgery, among which the Gel+ACS+Cell group had the smallest color difference and appearance.

[0272] The Wayne scoring system shown in Table 6 was used to score the regenerated tissue based on its macroscopic morphology, and the statistical results are shown in Figure 31. Specifically, 4 weeks after surgery, the total scores of each group were the control group (5.33±1.21), the Gel group (6.17±1.72), the Gel+ACS group (6.17±1.17), and the Gel+ACS+Cell group (6.33±1.75). 12 weeks after surgery, the total scores of each group were the control group (8.0±0.63), the Gel group (8.33±1.86), the Gel+ACS group (9.0±1.55), and the Gel+ACS+Cell group (10.5±1.64), with a statistically significant difference between the Gel+ACS+Cell group and the control group.

[0273] From 4 to 12 weeks after surgery, the restorative appearance and scores of all groups (including the control group) were significantly improved; based on the overall morphology and a simple scoring system, the therapeutic ability of the composite scaffolds, including the Gel+ACS group and the Gel+ACS+Cell group, on animal models was preliminarily verified. Whether 4 or 12 weeks after surgery, the Gel+ACS+Cell group in the experimental group showed better repair effects than the Gel group.

[0274] CT scan bone assessment

[0275] After macromorphological assessment, the specimens were fixed in 10% formalin and further assessed qualitatively and quantitatively for bone tissue regeneration using micro-computed tomography (Micro-CT) (DCT-110CC). Scanning parameters were as follows: power supply voltage: 75 kV, image pixel size: 7.5 μm, aluminum filter: 0.5 mm. During CT data analysis, a cylindrical region of interest (ROI) with a diameter of 3 mm was selected within the repaired area using CT-An software. Bone volume and thickness of the regenerated bone tissue were measured as bone volume per tissue volume (BV / TV) and trabecular thickness (Tb / Th).

[0276] Micro-CT analysis was used to qualitatively and quantitatively assess the extent of bone regeneration. Representative 2D Micro-CT images are shown in Figure 32A. Four weeks after surgery, bone regeneration was suboptimal in all groups, with no mineral matrix present within the defect cavity. Twelve weeks after surgery, defects in the experimental group showed significant filling, while the control group still exhibited complete cavities. Consistent with the repair performance observed in previous macromorphological assessments, the Gel+ACS group showed more significant filling than the Gel group at week 12, and the addition of IFPSCs to the Gel+ACS+Cell group further accelerated bone regeneration.

[0277] As shown in Figure 32B, the degree of bone regeneration was quantified by calculating the bone volume per tissue volume (BV / TV) and trabecular thickness (Tb.Th). Bone density was used to analyze whether the bone tissue had osteoporosis, and trabecular thickness was used to determine whether the new bone tissue was abnormal. In subsequent analyses, normal bone tissue data (n=2) was introduced as a standard control group and a reference for analysis. The bone volume BV / TV (%) value was 57.28±0.44%, and the trabecular thickness Tb.Th (mm) value was 0.27±0.01 mm.

[0278] Four weeks after surgery, the BV / TV values ​​of the groups were control group (24.20±1.39), Gel group (23.88±5.54), Gel+ACS group (23.67±4.58), and Gel+ACS+Cell group (20.46±9.80), respectively. There was no significant difference between the groups, and there was an obvious decrease compared with normal bone tissue. Twelve weeks after surgery, the BV / TV values ​​reached control group (32.29±15.57), Gel group (40.56±3.12), Gel+ACS+Cell group (41.40±3.03), and Gel+ACS+Cell group (43.37±4.09). The S group (39.14±7.79) and the Gel+ACS+Cell group (49.83±6.54) showed significant improvement compared with 4 weeks after surgery, among which the Gel+ACS+Cell group had a better repair effect than the other groups; the bone density analysis results were consistent with the 2Dmicro-CT observation results; the regeneration of the control group did not improve significantly 12 weeks after surgery, while the experimental groups showed more bone regeneration effects, among which the Gel+ACS+Cell group had the most obvious effect, but it was still slightly lower than normal bone tissue.

[0279] Four weeks after surgery, the Tb.Th values ​​of each group were control group (0.17±0.04), Gel group (0.22±0.01), Gel+ACS group (0.23±0.02), and Gel+ACS+Cell group (0.21±0.07), and there were significant differences between the Gel group and the Gel+ACS group and the control group; 12 weeks after surgery, the Tb.Th values ​​of each group were control group (0.22±0.04), Gel group (0.27±0.03), Gel+ACS group (0.24±0.02), and Gel+ACS+Cell group (0.27±0.04), and there were significant differences between the Gel group and the control group. values, but all groups except the Gel+ACS group showed no significant differences from normal bone tissue (*p<0.05, **p<0.005, ***p<0.001, ****p<0.0001); the trabecular thickness showed a large difference at 4 weeks after surgery, but decreased at 12 weeks after surgery and was very close to that of normal bone tissue in terms of value; based on the aforementioned bone volume BV / TV and trabecular Tb.Th analysis, it was confirmed that after transplanting Example 19, Example 20 and Example 21 into the knee cartilage defect, the repaired bone tissue reached the healthy standard, and the bone coverage density of Example 21 (Gel+ACS+Cell group) performed the best among the experimental groups.

[0280] Histological staining

[0281] Hematoxylin and Eosin (HE) histological staining

[0282] After the aforementioned samples were fixed in formalin and subjected to Micro-CT analysis, they were further decalcified by immersion in 10% formic acid. After complete decalcification, the samples were cut and sent to the National Cheng Kung University Hospital Human Database for paraffin embedding and sectioning. The tissue sections were stained with HE and SO staining to obtain the internal appearance and glycosaminoglycan (GAG) content. Finally, the histological scoring system proposed by Wakitani et al. in Blocking of tumor necrosis factor activity promotes natural repair of osteochondral defects in rabbit knee (Acta Orthop, 80(5), 606-611) was used to score the performance of cartilage regeneration. The scoring criteria are shown in Table 7.

[0283] Table 7

[0284] The regenerated cartilage at the defect site was histologically evaluated using HE and SO staining; the HE staining results are shown in Figures 33A to 33B, and the SO staining results are shown in Figure 34; the boundary of the cartilage tissue defect is marked with a red arrow, and the red box in Figures 33A to 33B is the boundary of the tissue defect, and the approximate location of the image photographed at high magnification on the right is shown.

[0285] As shown in Figure 33A, 4 weeks after surgery, the surface and underlying structures of the control group were disorganized; the surface of the Gel group was slightly damaged, but the cartilage thickness and underlying layer showed regeneration; in the Gel+ACS group, obvious residual multilayer composite scaffold structure could be observed under the cartilage; the surface of the Gel+ACS+Cell group was relatively smooth, and the surface thickness was intact, but like the Gel+ACS group, residual scaffold structure could be observed under the cartilage.

[0286] As shown in Figure 33B, 12 weeks after surgery, the structure of the control group was improved compared with 4 weeks after surgery, but there were still large defects; the cartilage thickness in the Gel group was thicker than 4 weeks after surgery, but there were still some gaps and undulations on the surface; the defect in the Gel+ACS group retained the structure of the composite scaffold, while the cartilage thickness of the Gel+ACS+Cell group was sufficient and smooth, and the composite scaffold structure was also significantly reduced.

[0287] Safranin O (SO) histological staining

[0288] As shown in Figure 34, in the results of SO staining, the Gel+ACS group and the Gel+ACS+Cell group showed rich glycosaminoglycan (GAG) color 4 weeks after surgery; 12 weeks after surgery, not only the Gel+ACS group and the Gel+ACS+Cell group, but also the Gel group had a significantly improved coloring degree.

[0289] Cartilage regeneration was evaluated according to the histological scoring system listed in Table 7, with a maximum score of 16. Specific results are shown in Figure 35 . Four weeks after surgery, the scores of the groups were: control group (7.167±1.941), Gel group (8.167±2.229), Gel+ACS group (8.5±2.168), and Gel+ACS+Cell group (8.5±1.871). Twelve weeks after surgery, the scores of the groups were: control group (9.667±1.033), Gel group (11.83±1.472), Gel+ACS group (10.83±1.472), and Gel+ACS+Cell group (13.33±2.16). There were statistically significant differences between the Gel+ACS+Cell group and both the control group and the Gel+ACS group (*p<0.05, **p<0.005).

[0290] Chondrocyte morphology reveals not only the presence of fibrocartilage but also the number of chondrocytes and cell clusters, which can be used to assess the overall cartilage condition. Increased chondrocyte numbers and the presence of numerous cell clusters indicate mild osteoarthritis degeneration. Following treatment, the Gel group primarily exhibited some fibrocartilage phenotypes, while the Gel+ACS group exhibited numerous cell clusters 12 weeks post-surgery, indicating a degenerative state. However, the tissue surface in the Gel+ACS group was incomplete 4 weeks post-surgery, but cell morphology remained normal. In contrast, the Gel+ACS+Cell group displayed normal cell types and numbers 12 weeks post-surgery, demonstrating no degeneration.

[0291] In general, the above embodiments and experimental examples fully verify that the biomimetic tissue implant (100) provided by the present invention has good biocompatibility, mechanical properties and tissue repair ability; the present invention combines decellularized tissue and biocompatible materials in a stacked form to form a sandwich composite scaffold, and then adds target cells of the corresponding tissue to meet the elements such as cells, scaffolds and biological signals required in tissue engineering; the decellularized tissue layer retains the original three-dimensional structure of the tissue, enhances the mechanical properties of the biocompatible material, and at the microscopic level, the biocompatible material is tightly cross-linked with the matrix macromolecules retained by the decellularized tissue layer.

[0292] Animal experiments have shown that the use of sandwich composite scaffolds, including the combination of hydrogel and decellularized tissue, as well as the combination of hydrogel, decellularized tissue and cells, exhibited excellent cartilage regeneration function in cartilage defect repair experiments; the integration of the new cartilage tissue with the surrounding cartilage was also quite significant, and the addition of decellularized tissue reduced the degradation and integration rate of the scaffold; the added cells further promoted the regeneration and repair of the surface cartilage, and increased the recovery of the bone density of the subchondral layer, which was more conducive to improving bone density; in addition, the thickness of the trabeculae was restored to the level of normal bone tissue after the implementation of the bionic tissue implant (100), so that the bone tissue after the defect repair had stable compressive resistance; therefore, the bionic tissue implant (100) provided by the present invention can be used not only for the treatment of tissue defects, but also has the potential to improve the health of the defect tissue after repair.

[0293]

Explanation of symbols

[0294] 100 Bionic tissue implants

[0295] 1. First bioscaffold layer

[0296] 10 Decellularized tissue layer

[0297] 10a - Decellularized tissue layer

[0298] 10' Another decellularized tissue layer

[0299] 2. Second bioscaffold layer

[0300] (Steps S1 to S3)

[0301] (Steps S2a to S2b)

[0302] (Steps S3a to S3b)

[0303] (Steps S11 to S14)

[0304] (Steps S10a to S10b)

[0305] (Steps S1' to S2')

[0306] (Step S1", Steps S2a"-1 to 2a"-2, Step S2b")

[0307] (Steps S3a"-1 to S3a"-2, Step S3b)

Claims

1. A bionic tissue implant, characterized in that: include: Decellularized tissue layer (10); A first bioscaffold layer (1), disposed on one side of the decellularized tissue layer (10), comprising a first biocompatible polymer and target cells; and The second bioscaffold layer (2) is arranged on the other side of the decellularized tissue layer (10) relative to the first bioscaffold layer (1), and comprises a second biocompatible polymer, wherein the first biocompatible polymer is the same as or different from the second biocompatible polymer.

2. The bionic tissue implant according to claim 1, characterized in that: The first biocompatible high molecular polymer is cross-linked with the decellularized tissue layer (10) through multiple covalent bonds, and the second biocompatible high molecular polymer is cross-linked with the decellularized tissue layer (10) through multiple covalent bonds.

3. The bionic tissue implant according to claim 1 or 2, characterized in that: The first biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof; the second biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof.

4. The bionic tissue implant according to claim 1 or 2, characterized in that: The target cells include fibroblasts, glial cells, interstitial cells, embryonic stem cells, umbilical cord mesenchymal stem cells, bone marrow mesenchymal stem cells, adipose stem cells or lymphocytes.

5. The bionic tissue implant according to claim 1 or 2, characterized in that: The decellularized tissue layer (10) comprises decellularized tissue, which includes decellularized liver tissue, decellularized fat, decellularized skin, decellularized cartilage, decellularized myocardial tissue, decellularized pericardium or decellularized blood vessels.

6. A method for preparing a bionic tissue implant, characterized in that: include: Taking a decellularized tissue and planting target cells on one side of the decellularized tissue; Taking a first biocompatible coating, which comprises a first biocompatible precursor polymer, and coating the first biocompatible coating on the side of the decellularized tissue implanted with the target cells to form a first bioscaffold layer (1), wherein the first biocompatible precursor polymer comprises a first biocompatible high molecular polymer; and A second biocompatible coating is obtained, which comprises a second biocompatible precursor polymer, and the second biocompatible coating is applied to the side of the decellularized tissue opposite to the side where the target cells are implanted to form a second bioscaffold layer (2) to obtain a biomimetic tissue implant (100), wherein the second biocompatible precursor polymer comprises a second biocompatible high molecular polymer, wherein the first biocompatible high molecular polymer is the same as or different from the second biocompatible high molecular polymer.

7. A method for preparing a bionic tissue implant, characterized in that: include: The decellularized tissue is taken as the decellularized tissue layer; Mixing a first biocompatible coating and target cells to obtain a first biocompatible cell mixture, and coating the first biocompatible cell mixture on one side of the decellularized tissue layer (10) to form the first bioscaffold layer (1), wherein the first biocompatible precursor polymer comprises a first biocompatible high molecular polymer; and A second biocompatible coating is obtained, which comprises a second biocompatible precursor polymer, and the second biocompatible coating is applied to the side of the decellularized tissue opposite to the side where the target cells are implanted to form a second bioscaffold layer (2) to obtain a biomimetic tissue implant (100), wherein the second biocompatible precursor polymer comprises a second biocompatible high molecular polymer, wherein the first biocompatible high molecular polymer is the same as or different from the second biocompatible high molecular polymer.

8. The method for preparing a bionic tissue implant according to claim 6 or 7, characterized in that: The first biocompatible coating comprises: A first precursor coating, comprising the first biocompatible precursor polymer and a first cross-linking enzyme; and A second precursor coating, comprising the first biocompatible precursor polymer and a first cross-linking agent; The second biocompatible coating comprises: A third precursor coating, comprising the second biocompatible precursor polymer and a second cross-linking enzyme; and A fourth precursor coating comprises the second biocompatible precursor polymer and a second cross-linking agent, wherein before forming the first biological scaffold layer (1) and forming the second biological scaffold layer (2), the preparation method of the bionic tissue implant (100) further comprises: Simultaneously or separately coating the first leading coating and the second leading coating on the side of the decellularized tissue seeded with the target cells to obtain a first leading scaffold layer; Simultaneously or separately coating the third leading coating and the fourth leading coating on the side of the decellularized tissue opposite to the side where the target cells are implanted, to obtain a second leading scaffold layer; The first cross-linking enzyme is brought into contact with the first cross-linking reagent for a first time so that the first biocompatible polymer and the decellularized tissue layer (10) are cross-linked, so that the first leading scaffold layer forms the first bioscaffold layer (1); the second cross-linking enzyme is brought into contact with the second cross-linking reagent for a second time so that the second biocompatible polymer and the decellularized tissue layer (10) are cross-linked, so that the second leading scaffold layer forms the second bioscaffold layer (2), wherein the first time is equal to or less than the second time.

9. The method for preparing a bionic tissue implant according to claim 8, characterized in that: The first cross-linking enzyme comprises horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxidoreductase, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase or catechol oxidase; the second cross-linking enzyme comprises horseradish peroxidase, glutathione peroxidase, haloperoxidase, myeloperoxidase, catalase, heme protein, peroxide, peroxidoreductase, animal heme-dependent peroxidase, thyroid peroxidase, vanadium bromoperoxidase, lactoperoxidase, tyrosinase or catechol oxidase.

10. The method for preparing a bionic tissue implant according to claim 9, characterized in that: The first biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof; the second biocompatible high molecular polymer includes fibrin, collagen gel, agar, agarose, gelatin, wood glue, guar gum, xanthan gum or any combination thereof.

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