Method for preparing a multiphasic "core-shell" biomimetic tissue engineering scaffold

Abstract

The application discloses a preparation method of a multi-phase "core-shell" biomimetic tissue engineering support, and constructs a multi-phase "core-shell" in-situ biomimetic support to promote the reconstruction of a rotator cuff tendon-bone interface; the support is made of PCL / PEO biological materials and mixed with a plurality of bioactive factor microspheres, the support fiber is a "core-shell" structure, the "shell" contains SDF-1 chitosan slow-release microspheres, the "inner core" is a bone tissue layer, a fibrocartilage layer and a tendon layer, the layers are continuously printed, the spatial distribution of various factors is fine and controllable, the tendon-bone interface is reconstructed in a biomimetic manner, the sequential controllable release of SDF-1 in the outer layer and various differentiation induction factors in the inner core is realized by using the "core-shell" structure fiber, the endogenous MSCs recruitment and the ordered directional differentiation of each layer are promoted, and the application develops an electrohydrodynamic coaxial 3D printing technology on the basis of the electrohydrodynamic printing technology, and constructs a multi-phase "core-shell" in-situ biomimetic support to promote the reconstruction of the rotator cuff tendon-bone interface.

Description

Technical Field

[0001] This invention relates to the field of bionics, and in particular to a method for preparing a multiphase "core-shell" bionic tissue engineering scaffold that promotes tendon-bone interface healing. Background Technology

[0002] Synthetic biodegradable materials are commonly used scaffold materials in tendon-bone tissue engineering. They possess good biocompatibility and suitable mechanical strength, serving as a primordium for protection, accommodating seed cell proliferation, and generating extracellular matrix before tendon-bone interface tissue reconstruction. These materials are inexpensive, the printing technology is mature, and no special maintenance is required after printing, resulting in good overall economic efficiency. Polymerized fibrous tissue (PCL) is a common material used in tendon-bone implantation. It has an elastic modulus suitable for tendon-bone tissue engineering, good biocompatibility, and is suitable for tissue growth. However, previous studies have shown that PCL scaffold fibers are relatively coarse, and their degradation rate is too slow, with significant residues remaining even after two years. Furthermore, the lack of sufficient stem cells around the implanted PCL scaffold severely limits the bioavailability of the compounds released from the scaffold. Therefore, to maximize the biological effects of PCL tissue engineering scaffolds, which cannot carry cells and serve as a primordium for inducing tendon-bone interface reconstruction, it is first necessary to recruit MSCs from the surrounding environment and circulatory system to the rotator cuff tear site, followed by the induction of directed differentiation of MSCs from different layers of the scaffold. The sequential chemotaxis-differentiation behavior enables the orderly regulation of MSC recruitment and differentiation, making in-situ reconstruction of the rotator cuff tendon-bone interface more effective. However, the reconstruction of the complex multilayered structure of the tendon-bone interface requires not only the sequential release of chemokines and differentiation factors, but also the controlled and sustained release of various factors to adapt to the time window of tendon-bone interface tissue reconstruction. Therefore, it is not an ideal scaffold for rotator cuff tissue engineering.

[0003] Existing engineering scaffolds utilize electrospinning technology to construct biphasic films, spinning polycaprolactone into two disordered and dense layers. The lower layer is loaded with hydroxyapatite to promote osteogenic formation, while the upper layer, pure PCL, does not promote osteogenic formation. The biphasic film is then sutured at the ruptured rotator cuff to promote healing. For example, patent application CN111214709A discloses a novel electrospun biphasic nerve conduit, consisting of an electrospun decellularized porcine neural matrix conduit layer and a PLCL layer disposed on the outer layer of the electrospun decellularized porcine neural matrix conduit. However, this method has drawbacks: the dense structure prevents cells from penetrating and growing; it lacks a 3D structure, leading to inaccurate spatial positioning of materials and drugs.

[0004] Existing engineering scaffolds also include decellularized scaffold structures, which involve specially treating animal rotator cuff tissue to remove cells, rendering it non-immunogenic while maintaining its basic structure and extracellular matrix. These scaffolds are then sutured to the torn rotator cuff to promote repair. For example, patent application CN116271243A discloses a decellularized matrix composite rotator cuff patch, which includes a decellularized dermis layer and a decellularized pericardium layer, with the decellularized dermis layer as the upper layer and the decellularized pericardium layer as the lower layer. Its disadvantages include: cumbersome preparation, poor reproducibility, and inability to standardize production. Summary of the Invention

[0005] To overcome the shortcomings of the prior art, the present invention aims to provide a method for preparing a multiphase "core-shell" biomimetic tissue engineering scaffold that promotes tendon-bone interface healing. Based on electrohydrodynamic printing technology, an electrohydrodynamic coaxial 3D printing technology was developed to construct a multiphase "core-shell" in-situ biomimetic scaffold to promote the reconstruction of the rotator cuff tendon-bone interface. The scaffold is made of PCL / PEO biomaterial mixed with microspheres loaded with various bioactive factors. The scaffold fibers have a "core-shell" structure. The "outer shell" contains chemokine SDF-1 chitosan sustained-release microspheres. The "inner core" is a layered and continuously printed structure consisting of bone tissue layer, fibrocartilage layer, and tendon layer. Through precise and controllable spatial distribution of various factors, biomimetic reconstruction of the tendon-bone interface is achieved. At the same time, the "core-shell" structure fibers enable the sequential and controllable release of SDF-1 in the outer layer and various differentiation-inducing factors in the inner core, promoting the recruitment of endogenous MSCs and the orderly directional differentiation of each layer.

[0006] To achieve the above objectives, the technical solution of the present invention is as follows:

[0007] The method for fabricating a multiphase "core-shell" biomimetic tissue engineering scaffold includes the following steps:

[0008] (1) Preparation of chitosan nanospheres

[0009] (1.1) Add chitosan to a dilute glacial acetic acid solution with a specification of 1v%-5v% at a ratio of 2w / v%-5w / v%, and stir until it is completely dissolved to obtain a chitosan glacial acetic acid solution.

[0010] (1.2) Add BMP-2, TGF-β, bFGF and SDF-1 to a sodium tripolyphosphate solution with a specification of 1w / v%-5w / v% at a concentration of 2ug / mL, respectively. Stir magnetically to make the distribution of each factor in the sodium tripolyphosphate solution uniform, and obtain four kinds of sodium tripolyphosphate mixed liquids.

[0011] (1.3) Add the four sodium tripolyphosphate mixtures to the prepared chitosan glacial acetic acid solution at a volume ratio of 1:(50-100) and stir magnetically until homogeneous to obtain the four mixtures.

[0012] (1.4) The four mixtures obtained in step (1.3) were sonicated for 5 min-10 min at a power of 60-70 W, with sonication for 2-4 s and pause for 4-6 s until completely dispersed, and then centrifuged to obtain the initial microspheres.

[0013] (1.5) Wash with dH2O and centrifuge again to obtain the final four types of microspheres, namely chitosan nanoparticles loaded with BMP-2, TGF-β, bFGF and SDF-1;

[0014] (1.6) Resuspend using dH2O and freeze overnight at -80°C or below;

[0015] (1.7) Store at -20°C or below after freeze-drying;

[0016] (2) Preparation of core materials and shell printing materials for osteogenic, chondrogenic and tendonogenic materials

[0017] (2.1) 1 w / v%-2 w / v% PCL and 4 w / v%-6 w / v% PEO were mixed at a volume ratio of 1:1, and then added to chitosan nanospheres loaded with BMP-2, TGF-β and bFGF at a concentration of 1 w / v%-5 w / v% respectively to construct three printing materials: osteogenic, cartilaginous, and tendon-forming cores. They were named as follows: PCL / PEO / BMP-2 as the osteogenic core material, PCL / PEO / TGF-β as the cartilaginous core material and PCL / PEO / bFGF as the tendon-forming core material.

[0018] (2.2) Mix 5-10 w / v% PCL and 4 w / v%-6 w / v% PEO at a volume ratio of 1:1; then add chitosan nanospheres loaded with SDF-1 at a concentration of 1 w / v%-5 w / v% to construct the shell printing material, which is a shell material PCL / PEO / SDF-1 with stem cell chemotaxis function.

[0019] (3) Adjusting the parameters of electrohydrodynamic coaxial 3D printing

[0020] (3.1) The moving speed of the XYZ triaxial displacement system, the voltage of the metal coaxial needle connected to the high voltage power supply, and the receiving distance are fixed at 10mm / s-150mm / s, 1kV-5kV, and 1.5mm-4mm, respectively.

[0021] (3.2) Add the core material obtained in step (2.1) to Rhodamine for staining. The Rhodamine is added at a concentration of 0.5-1 w / v%. Adjust the flow rate ratio of the core and shell materials to (1-9): (9-1). Set the total flow rate of the core and shell materials to 50-100 μL / h and then print.

[0022] (3.3) Select the optimal ratio of inner and outer layer material flow rate observed under an optical microscope and keep it constant. Increase the total flow rate from 10 μL / h to 100 μL / h in sequence. Observe and measure the diameter of the inner and outer layers of the fiber under an optical microscope.

[0023] (3.4) Select parameters that meet the product requirements as observed under an optical microscope.

[0024] (4) 3D bioprinting of multiphase “core-shell” in situ biomimetic scaffold

[0025] (4.1) Prepare a dilute agarose solution using ultrapure water, heat and stir until the agarose solvent is completely dissolved, and then spin-coat it onto the conductive surface of indium tin oxide glass.

[0026] (4.2) Load the inner core material and outer shell material into the corresponding injectors respectively;

[0027] (4.3) Connect the tendon core material to the metal coaxial needle core through a polytetrafluoroethylene tube, connect the shell material with stem cell chemotaxis function to the outer layer of the metal coaxial needle, and continuously print according to the target requirements to form the top tendon region with fiber spacing <200μm.

[0028] (4.4) Then switch the core material to the cartilaginous core material, keep the outer shell printing material unchanged, and continue printing according to the target requirements to form the cartilaginous region in the middle with a fiber spacing of 200μm-400μm;

[0029] (4.5) Finally, switch the inner core material to the osteogenic inner core material, while keeping the outer shell printing material unchanged; continue printing according to the target requirements to form the bottom osteogenic region with a fiber spacing of 400μm-800μm;

[0030] (4.6) Following steps (4.3)-(4.6), a complete three-phase "core shell" in-situ bionic support is obtained and stored in a refrigerator.

[0031] Compared with the prior art, the advantages of the present invention are:

[0032] (1) By applying a high voltage to the metal coaxial needle, the material is extruded from the metal coaxial needle under the combined action of electrostatic force, extrusion force, shear force, etc. By adjusting the distance between the metal coaxial needle and the receiving substrate, the material is uniformly deposited on the receiving substrate, thus realizing the controllable printing of three-dimensional structures. This method can print fibers that are much smaller than the inner diameter of the metal coaxial needle, thereby improving the printing accuracy and realizing the precise spatial positioning of biodegradable materials loaded with different substances in the scaffold. Furthermore, the extremely fine diameter of the fibers can significantly accelerate their degradation rate and avoid becoming an obstacle in the tissue repair process, providing a new method for constructing an ultrathin gradient structure of the natural rotator cuff tendon-bone interface.

[0033] (2) This invention develops "electrohydrodynamic coaxial 3D printing technology". This 3D printing technology uses a metal coaxial needle to extrude a scaffold fiber "shell" containing stem cell chemokines and a scaffold fiber "core" containing cell differentiation factors of different functional layers. Chemokines are placed in the fiber shell and are preferentially released under the action of body fluids and cell degradation; differentiation factors are located in the fiber core and are sequentially released as the scaffold fibers degrade after the chemokines are recruited to MSCs. By utilizing the spatial distribution of chemokines and differentiation factors in the "shell-core" structure, the time problem of sequential release of chemokines and differentiation factors is solved, which has the characteristics of controllable release and orderly differentiation.

[0034] (3) Growth factors are mostly proteins with short half-lives and relatively low stability. They are easily inactivated by various enzymes in the in vivo environment. Therefore, the application of various growth factors needs to be carried out under the protection of special carriers. Chitosan often carries a positive charge, while growth factors mostly carry a negative charge. Therefore, chitosan nanospheres are more capable of effectively loading factors such as bFGF and BMP-2. Chitosan nanospheres are used to encapsulate chemokine SDF-1 and growth factors BMP-2, TGF-β and bFGF to play a protective and sustained-release role.

[0035] (4) Polyethylene oxide (PEO) is a high molecular weight polymer with molecular weights ranging from 100,000 to 8 million. It is commonly used as a thickener in water-soluble adhesives. A 1%–5% (w / v%) aqueous solution has high viscosity. Mixing PCL with PEO can significantly increase the viscosity of the material, giving it good stringability in EHDP technology and increasing the stability of the printing process. After printing, the PEO material inside the scaffold is washed away, and a loose porous structure is formed on the scaffold surface, which not only promotes the degradation of PCL material but also makes it more suitable for cell adhesion and growth. The slow degradation rate of scaffolds obtained by electrohydrodynamic printing using pure PCL material is solved by using a sacrificial material in solution-state electrohydrodynamic printing.

[0036] In summary, this invention, based on electrohydrodynamic printing technology, developed "electrohydrodynamic coaxial 3D printing technology" to construct a multiphase "core-shell" in-situ biomimetic scaffold to promote the reconstruction of the rotator cuff tendon-bone interface. The prepared scaffold is made of PCL / PEO biomaterial mixed with microspheres loaded with various bioactive factors. It has the characteristics of good repeatability, can achieve standardized production, and has a significant promoting effect on interface structure remodeling. Attached Figure Description

[0037] Figure 1 This is a flowchart of the in-situ biomimetic scaffold fabrication method for multiphase "core-shell".

[0038] Figure 2(a) is a schematic image of the 3D printing platform, and Figure 2(b) is a magnified image of the needle during printing.

[0039] Figure 3 The ratio of the inner and outer material flow rates is set to 1:1, the total material flow rate is 50 μ / h, and the core-shell structure image is shown.

[0040] Figure 4 The ratio of the inner and outer material flow rates is set to 1:1, the total material flow rate is 10 μ / h, and the core-shell structure image is shown.

[0041] Figure 5 This is a general image of the stent (2cm×2cm×300μm).

[0042] Figure 6(a) shows the gradient structure of the support structure observed by scanning electron microscopy.

[0043] Figure 6(b) shows the gradient structure of the support structure observed by a laser confocal microscope.

[0044] Figure 7 It refers to the release rate of factors in each functional layer of the "core-shell" structure.

[0045] Figure 8 Image of the core-shell structure with a fiber internal and external material flow ratio of 1:1 and a total material flow rate of 100 μ / h.

[0046] Figure 9 This is a schematic diagram of a rabbit rotator cuff injury stent implantation surgery. Figure 9 (a) in the image shows the site of the rabbit rotator cuff injury where the stent was implanted; the yellow arrow indicates the implanted stent. Figure 9 (b) shows the use of sutures to fix the tendon to the footprint area of ​​the humerus; Figure 9 (c) in the figure represents the cross-sectional area of ​​the rotator cuff insertion point after surgery; Figure 9 (d) in the figure represents the maximum failure load for repairing the rotator cuff. Figure 9 (e) in the figure represents the maximum stress value for rotator cuff repair, and * indicates P<0.05. Detailed Implementation

[0047] The present invention will now be described in further detail with reference to the accompanying drawings and specific embodiments.

[0048] Example 1

[0049] The method for fabricating a multiphase "core-shell" biomimetic tissue engineering scaffold includes the following steps:

[0050] (1) Preparation of chitosan nanospheres

[0051] (1.1) Chitosan was added to a 1v% dilute glacial acetic acid solution at a ratio of 2w / v% and stirred until completely dissolved to obtain a chitosan glacial acetic acid solution.

[0052] (1.2) Add BMP-2, TGF-β, bFGF and SDF-1 to a 1 w / v% sodium tripolyphosphate solution at a concentration of 2 ug / mL, respectively, and stir magnetically to make the components in the mixture evenly distributed, thus obtaining four sodium tripolyphosphate mixed liquids.

[0053] (1.3) Add the sodium tripolyphosphate mixture to the prepared chitosan glacial acetic acid solution at a volume ratio of 1:50 and stir magnetically until uniform to obtain the mixture.

[0054] (1.4) The four mixed liquids were sonicated for 5 min at 60 W, with sonication for 2 s and pause for 4 s until completely dispersed, and then centrifuged to obtain the initial microspheres.

[0055] (1.5) Wash with dH2O and centrifuge again to obtain the final four types of microspheres.

[0056] (1.6) Resuspended in dH2O overnight in a refrigerator at -80°C.

[0057] (1.7) After freeze-drying, store at -20℃ for later use.

[0058] (2) Preparation of core and shell printing materials for osteogenic, chondrogenic and tendon-forming materials

[0059] (2.1) PCL with a specification of 1 w / v% and PEO with a specification of 4 w / v% were first mixed as base materials, and then added to chitosan nanospheres loaded with BMP-2, TGF-β and bFGF respectively at a concentration of 1 w / v% to construct three printing materials for osteogenic, cartilaginous and tendon-forming cores respectively; namely: PCL / PEO / BMP-2 as osteogenic core material, PCL / PEO / TGF-β as cartilaginous core material and PCL / PEO / bFGF as tendon-forming core material.

[0060] (2.2) First, mix PCL with a specification of 5 w / v% and PEO with a specification of 4 w / v%, and then add chitosan nanospheres loaded with SDF-1 at a concentration of 1 w / v% to construct the shell printing material, which is a shell material PCL / PEO / SDF-1 with stem cell chemotaxis function.

[0061] (3) Optimization of electrohydrodynamic coaxial 3D printing technology

[0062] As shown in Figures 2(a) and 2(b), the electrohydrodynamic coaxial 3D printing equipment used in this embodiment mainly consists of three components: an XYZ triaxial displacement system (Alio Industries, USA), two injection pumps (Harvardapparatus, USA), and a metal coaxial needle connected to a high-voltage power supply (inner and outer needle sizes are 26G / 18G respectively); the specific printing steps are as follows:

[0063] (3.1) Fix the moving speed of the XYZ triaxial displacement system, the voltage of the metal coaxial needle connected to the high voltage power supply, and the receiving distance to 100mm / s, 2kV, and 2mm, respectively;

[0064] (3.2) Rhodamine was added to the core material for staining. The Rhodamine was added at a concentration of 0.5 w / v%. The total flow rate of the inner and outer layers was set to 10 μL / h and the flow rate ratio of the inner and outer layers was adjusted to 1:1 for printing.

[0065] (3.3) The optimal ratio of inner and outer layer material flow rates observed under an optical microscope was kept constant, and the total flow rate was increased sequentially from 10 μL / h to 100 μL / h. The diameters of the inner and outer layers of the fiber were observed and measured under an optical microscope. Figure 3 , Figure 4 As shown, the diameters of the inner and outer layers of the fiber were observed and measured under an optical microscope to observe the printing effect;

[0066] (3.4) Select the most suitable parameters observed under an optical microscope for printing.

[0067] (4) 3D bioprinting of multiphase “core-shell” in situ biomimetic scaffold

[0068] (4.1) Prepare a dilute agarose solution using ultrapure water, heat and stir until the solvent is completely dissolved, and then spin-coat it onto the conductive surface of indium tin oxide glass.

[0069] (4.2) Load all printing materials into different injectors.

[0070] (4.3) Connect the tendon core material to the metal coaxial needle core through a polytetrafluoroethylene tube, and connect the shell material with stem cell chemotaxis function to the outer layer of the metal coaxial needle. Print 30 layers continuously to form the top tendon region with a fiber spacing of 100 μm.

[0071] (4.4) Switch the inner core to the cartilage core material, keep the outer printing material unchanged, and continue printing 30 layers to form the central cartilage region with a fiber spacing of 200 μm.

[0072] (4.5) Switch the inner core material to the osteogenic inner core material, keep the outer printing material unchanged, and continue printing 30 layers to form the bottom osteogenic region with a fiber spacing of 400μm.

[0073] (4.6) Following steps (4.3)-(4.6), a complete multiphase “core shell” in-situ biomimetic support is obtained and stored in a refrigerator.

[0074] Example 2

[0075] The method for fabricating a multiphase "core-shell" biomimetic tissue engineering scaffold includes the following steps:

[0076] (1) Preparation of chitosan nanospheres

[0077] (1.1) Specification is 2 v Chitosan was added to a dilute glacial acetic acid solution at a ratio of 4 w / v and stirred until completely dissolved to obtain a chitosan glacial acetic acid solution.

[0078] (1.2) Add BMP-2, TGF-β, bFGF and SDF-1 to a 2 μg / mL container at a concentration of 2 μg / mL. w / v% Four sodium tripolyphosphate mixtures were obtained by magnetic stirring in a sodium tripolyphosphate solution to ensure uniform distribution of all factors within the mixture.

[0079] (1.3) Add the sodium tripolyphosphate mixture to the prepared chitosan glacial acetic acid solution at a volume ratio of 1:75 and stir magnetically until uniform to obtain the mixture.

[0080] (1.4) The four mixed liquids were sonicated for 7 min at 60 W, with sonication for 4 s and pause for 6 s until completely dispersed, and then centrifuged to obtain the initial microspheres.

[0081] (1.5) Wash with dH2O and centrifuge again to obtain the final four types of microspheres.

[0082] (1.6) Resuspended in dH2O overnight in a -85°C refrigerator.

[0083] (1.7) After freeze-drying, store at -25℃ for later use.

[0084] (2) Preparation of core and shell printing materials for osteogenic, chondrogenic and tendon-forming materials

[0085] (2.1) PCL with a specification of 1.5 w / v% and PEO with a specification of 5 w / v% were first mixed as base materials, and then added to chitosan nanospheres loaded with BMP-2, TGF-β and bFGF respectively at a concentration of 2 w / v% to construct three printing materials for osteogenic, cartilaginous and tendon-forming cores respectively; namely: PCL / PEO / BMP-2 as osteogenic core material, PCL / PEO / TGF-β as cartilaginous core material and PCL / PEO / bFGF as tendon-forming core material.

[0086] (2.2) First, mix PCL with a specification of 7.5 w / v% and PEO with a specification of 5 w / v%, and then add chitosan nanospheres loaded with SDF-1 at a concentration of 2 w / v% to construct the shell printing material, which is a shell material PCL / PEO / SDF-1 with stem cell chemotaxis function.

[0087] (3) Optimization of electrohydrodynamic coaxial 3D printing technology

[0088] The electrohydrodynamic coaxial 3D printing technology used in this embodiment mainly consists of three components: an XYZ triaxial displacement system (Alio Industries, USA), two injection pumps (Harvard Apparatus, USA), and a metal coaxial needle connected to a high-voltage power supply (inner and outer needle sizes are 26G / 18G respectively); the specific printing steps are as follows:

[0089] (3.1) Fix the moving speed of the XYZ triaxial displacement system, the voltage of the metal coaxial needle connected to the high voltage power supply, and the receiving distance to 100mm / s, 3kV, and 2.5mm, respectively;

[0090] (3.2) Rhodamine was added to the core material for staining. The addition of rhodamine was based on a concentration of 1 w / v%. The total flow rate of the inner and outer layers was set to 50 μL / h and kept constant. The flow rate ratio of the inner and outer layers was adjusted to (1-9): (1-9) for printing. The diameter of the inner and outer layers of the fiber was observed and measured under an optical microscope to observe the printing effect.

[0091] (3.4) Select the most suitable parameters observed under an optical microscope for printing. In this example, the total flow rate of the inner and outer layers is selected as 50 μL / h, and the flow rate ratio of the inner and outer layer materials is 1:1. Figure 4 As shown.

[0092] (4) 3D bioprinting of multiphase “core-shell” in situ biomimetic scaffold

[0093] (4.1) Prepare a dilute agarose solution using ultrapure water, heat and stir until the solvent is completely dissolved, and then spin-coat it onto the conductive surface of indium tin oxide glass.

[0094] (4.2) Load all printing materials into different injectors.

[0095] (4.3) Connect the tendon core material to the metal coaxial needle core through a polytetrafluoroethylene tube, and connect the shell material with stem cell chemotaxis function to the outer layer of the metal coaxial needle. Print 50 layers continuously to form the top tendon region with a fiber spacing of 150 μm.

[0096] (4.4) Switch the inner core to the cartilage-forming inner core material, keep the outer printing material unchanged, and continue printing 50 layers to form the central cartilage-forming region with a fiber spacing of 300 μm.

[0097] (4.5) Switch the core material to the osteogenic core material, keep the outer printing material unchanged, and continue printing 50 layers to form the bottom osteogenic region with a fiber spacing of 600 μm.

[0098] (4.6) Following steps (4.3)-(4.6), the complete multiphase "core-shell" in-situ biomimetic support is obtained and stored in a refrigerator. Figure 5 As shown.

[0099] Figure 6 shows the release rates of factors in each functional layer of the scaffold gradient structure and the "core-shell" structure of the ELISA kit detector, observed using scanning electron microscopy and laser confocal microscopy. Figure 7 As shown.

[0100] Example 3

[0101] The method for fabricating a multiphase "core-shell" biomimetic tissue engineering scaffold includes the following steps:

[0102] (1) Preparation of chitosan nanospheres

[0103] (1.1) Chitosan is added to a 5v% dilute glacial acetic acid solution at a ratio of 5w / v% and stirred until it is completely dissolved to obtain a chitosan glacial acetic acid solution.

[0104] (1.2) Add BMP-2, TGF-β, bFGF and SDF-1 to a 5 w / v% sodium tripolyphosphate solution at a concentration of 2 ug / mL, respectively, and stir magnetically to make the components evenly distributed in the mixture to obtain four sodium tripolyphosphate mixed liquids.

[0105] (1.3) Add the sodium tripolyphosphate mixture to the prepared chitosan glacial acetic acid solution at a volume ratio of 1:100 and stir magnetically until uniform to obtain the mixture.

[0106] (1.4) The four mixed liquids were sonicated for 10 min at 60 W, with a sonication time of 2 s and a pause of 4 s until they were completely dispersed. Then they were centrifuged to obtain the initial microspheres.

[0107] (1.5) Wash with dH2O and centrifuge again to obtain the final four types of microspheres.

[0108] (1.6) Resuspended in dH2O overnight in a refrigerator at -80°C.

[0109] (1.7) After freeze-drying, store at -20℃ for later use.

[0110] (2) Preparation of core and shell printing materials for osteogenic, chondrogenic and tendon-forming materials

[0111] (2.1) PCL with a specification of 2 w / v% and PEO with a specification of 6 w / v% were first mixed as base materials, and then added to chitosan nanospheres loaded with BMP-2, TGF-β and bFGF at a concentration of 5 w / v respectively to construct three printing materials for osteogenic, cartilaginous and tendon-forming cores respectively; namely: PCL / PEO / BMP-2 as osteogenic core material, PCL / PEO / TGF-β as cartilaginous core material and PCL / PEO / bFGF as tendon-forming core material.

[0112] (2.2) First, mix PCL with a specification of 10 w / v% and PEO with a specification of 6 w / v%, and then add chitosan nanospheres loaded with SDF-1 at a concentration of 5 w / v% to construct the shell printing material, which is a shell material PCL / PEO / SDF-1 with stem cell chemotaxis function.

[0113] (3) Optimization of electrohydrodynamic coaxial 3D printing technology

[0114] The electrohydrodynamic coaxial 3D printing equipment used in this embodiment mainly consists of three components: an XYZ triaxial displacement system (Alio Industries, USA), two injection pumps (Harvard Apparatus, USA), and a metal coaxial needle connected to a high-voltage power supply (inner and outer needle sizes are 26G / 18G respectively); the printing steps are as follows:

[0115] (3.1) Fix the moving speed of the XYZ triaxial displacement system, the voltage of the metal coaxial needle connected to the high voltage power supply, and the receiving distance to 150mm / s, 5kV, and 4mm, respectively.

[0116] (3.2) Rhodamine was added to the core material for staining. The addition of rhodamine was based on a concentration of 1 w / v%. The total flow rate of the inner and outer layers was set to a fixed value of 100 μL / h. The flow rate ratio of the inner and outer layers was adjusted to (1-9):(1-9) for printing. The diameter of the inner and outer layers of the fiber was observed and measured under an optical microscope to observe the printing effect.

[0117] (3.4) Select the most suitable parameters observed under an optical microscope for printing. In this example, the total flow rate of the inner and outer layers is selected as 50 μL / h, and the flow rate ratio of the inner and outer layer materials is 1:1. Figure 8 As shown.

[0118] (4) 3D bioprinting of multiphase “core-shell” in situ biomimetic scaffold

[0119] (4.1) Prepare a dilute agarose solution using ultrapure water, heat and stir until the solvent is completely dissolved, and then spin-coat it onto the conductive surface of indium tin oxide glass.

[0120] (4.2) Load all printing materials into different injectors;

[0121] (4.3) Connect the tendon core material to the metal coaxial needle core through a polytetrafluoroethylene tube, and connect the shell material with stem cell chemotaxis function to the outer layer of the metal coaxial needle. Print 100 layers continuously to form the top tendon region with a fiber spacing of 200 μm.

[0122] (4.4) Switch the inner core to the cartilage core material, keep the outer printing material unchanged, and continue printing 100 layers to form the central cartilage region with a fiber spacing of 400 μm.

[0123] (4.5) Switch the core material to the osteogenic core material, keep the outer printing material unchanged, and continue printing 100 layers to form the bottom osteogenic region with a fiber spacing of 800 μm.

[0124] (4.6) Following steps (4.3)-(4.6), a complete multiphase “core shell” in-situ biomimetic support is obtained and stored in a refrigerator.

[0125] Application Example 1

[0126] Animal experiments on tendon-bone interface repair in rabbit rotator cuffs, such as... Figure 9 As shown in Figure a, the scaffold can be easily sutured to the tendon-bone interface repair site. When the in-situ biomimetic scaffold for the tendon-bone interface prepared in Example 2 was applied to tendon-bone interface repair, there was no significant difference in the cross-sectional area of ​​the rotator cuff insertion region among the different groups at various postoperative time points. Figure 9 b). The maximum effective load and maximum stress values ​​in the multiphase core-shell scaffold group were significantly greater than those in the other three groups at 6 and 12 weeks post-operation (P<0.05). Figure 9 (cd). The tendon-bone junction has a unique structure and function, enabling it to evenly distribute stress passing through the junction and prevent tissue damage due to stress concentration. If proper remodeling is not achieved after rotator cuff repair surgery, irregular scar tissue can hardly withstand sufficient stress, leading to re-tears of the rotator cuff post-surgery. Therefore, the quality of rotator cuff repair depends primarily on its maximum stress tolerance, and postoperative mechanical properties depend mainly on the recovery of the transitional structure at the tendon-bone interface. The above results show that the multiphase "core-shell" in-situ biomimetic scaffold significantly promotes interfacial structural remodeling after rotator cuff injury surgery.

[0127] This multi-layered gradient scaffold can also be used for tendon-bone interface repair of other tendons (Achilles tendon, cruciate ligament, etc.).

Claims

1. A method for preparing a multiphase "core-shell" biomimetic tissue engineering scaffold, characterized in that, Includes the following steps: (1) Preparation of chitosan nanospheres (1.1) Add chitosan to a dilute glacial acetic acid solution with a specification of 1v%-5v% at a ratio of 2w / v%-5w / v%, and stir until it is completely dissolved to obtain a chitosan glacial acetic acid solution. (1.2) BMP-2, TGF-β, bFGF and SDF-1 were added to a sodium tripolyphosphate solution with a specification of 1 w / v%-5 w / v% at a concentration of 2 μg / mL, and the mixture was magnetically stirred to make the components evenly distributed in the mixture to obtain four sodium tripolyphosphate mixed liquids. (1.3) Add the sodium tripolyphosphate mixture to the prepared chitosan glacial acetic acid solution at a volume ratio of 1:50-100, and stir magnetically until uniform to obtain a mixture; (1.4) The four mixtures obtained in step (1.3) were sonicated for 5 min-10 min at a power of 60-70W, with sonication for 2-4 s and pause for 4-6 s until completely dispersed, and then centrifuged to obtain the initial microspheres; (1.5) Wash with dH2O and centrifuge again to obtain the final four types of microspheres, namely chitosan nanospheres loaded with BMP-2, TGF-β, bFGF and SDF-1; (1.6) Resuspend using dH2O and freeze overnight in a refrigerator at -80 ℃ or below; (1.7) Store at -20 °C or below after freeze-drying; (2) Preparation of osteogenic, chondrogenic and tendon-forming core materials and shell printing materials; (2.1) PCL with specifications of 1 w / v% - 2 w / v% and PEO with specifications of 4 w / v% - 6 w / v% were first mixed as base materials, and then added to chitosan nanospheres loaded with BMP-2, TGF-β and bFGF respectively at a concentration of 1 w / v% - 5 w / v% to construct three printing materials: osteogenic, cartilaginous and tendon core. They were named as follows: PCL / PEO / BMP-2 as the osteogenic core material, PCL / PEO / TGF-β as the chondrogenic core material, and PCL / PEO / bFGF as the tendonogenic core material. (2.2) First mix PCL with a specification of 5-10 w / v% and PEO with a specification of 4 w / v%-6 w / v%, then add chitosan nanospheres loaded with SDF-1 at a concentration of 1 w / v%-5 w / v% to construct the shell printing material, which is a shell material PCL / PEO / SDF-1 with stem cell chemotactic function; (3) Electrohydrodynamic coaxial 3D printing; (4) 3D bioprinting of multiphase "core-shell" in situ biomimetic scaffolds; (4.1) Prepare a dilute agarose solution using ultrapure water, heat and stir until the solvent is completely dissolved, and then spin-coat it onto the conductive surface of indium tin oxide glass; (4.2) Load all printing materials into the corresponding injectors; (4.3) Connect the myotenoid core material to the metal coaxial needle core through a polytetrafluoroethylene tube, connect the outer shell material with stem cell chemotaxis function to the outer layer of the metal coaxial needle, and continuously print according to the target requirements to form the top myotenoid region with fiber spacing <200 μm. (4.4) Then switch the core material to the cartilaginous core material, keep the outer shell printing material unchanged, and continue printing according to the target requirements to form the cartilaginous region in the middle with a fiber spacing of 200 μm - 400 μm; (4.5) Finally, the core material is switched to osteogenic core material, while the outer shell printing material remains unchanged; Continue printing according to the target requirements to form the bottom osteogenic region with a fiber spacing of 400 μm-800 μm; (4.6) Following steps (4.3)-(4.5), a complete three-phase "core shell" in-situ bionic support is obtained and stored in a refrigerator.

2. The method for preparing a multiphase "core-shell" biomimetic tissue engineering scaffold according to claim 1, characterized in that, The aforementioned electrohydrodynamic coaxial 3D printing specifically refers to: (3.1) The moving speed of the XYZ triaxial displacement system, the voltage of the metal coaxial needle connected to the high voltage power supply, and the receiving distance are fixed at 10 mm / s - 150 mm / s, 1 kV - 5 kV, and 1.5 mm - 4 mm, respectively. (3.2) Rhodamine was added to the core material for staining. The addition of rhodamine was based on a concentration of 0.5-1 w / v%. The total flow rate of the inner and outer layers was set to a fixed value of 50-100 μL / h. The flow rate ratio of the inner and outer layers was adjusted to (1-9):(9-1) for printing. (3.3) Select the optimal ratio of inner and outer layer material flow rate observed under an optical microscope and keep it constant. Increase the total flow rate from 10 μL / h to 100 μL / h in sequence. Observe and measure the diameter of the inner and outer layers of the fiber under an optical microscope. (3.4) Select the most suitable parameters observed under an optical microscope for printing.

Images