CN114129772A

CN114129772A - Preparation method of composite scaffold for bone tissue engineering

Info

Publication number: CN114129772A
Application number: CN202111159225.8A
Authority: CN
Inventors: 禹路; 魏瑞平; 胡旭华
Original assignee: Shanghai Haobo Medical Technology Co ltd
Current assignee: Shanghai Haobo Medical Technology Co ltd
Priority date: 2021-09-30
Filing date: 2021-09-30
Publication date: 2022-03-04

Abstract

The invention relates to a preparation method of a composite scaffold for bone tissue engineering, which comprises the following steps: preparing a PCL bracket, and performing electrostatic spinning on the PCL bracket to form the PCL-gelatin bracket. In the composite scaffold prepared by the preparation method, the PCL scaffold is used for simulating a cancellous bone structure, the PCL-gelatin scaffold is used for simulating a subchondral bone structure, and the interface bonding strength between the PCL scaffold and the PCL-gelatin scaffold is remarkably improved. The results of corresponding physical and chemical property detection and mechanical analysis show that the composite bracket has more excellent performance. In vivo and in vitro experimental results show that the composite scaffold is beneficial to the synchronous repair of osteochondral defects and has potential clinical application value.

Description

Preparation method of composite scaffold for bone tissue engineering

Technical Field

The invention relates to a preparation method of a composite scaffold for bone tissue engineering.

Background

Osteoarthritis and osteochondritis dissecans are a type of disease caused by degeneration or lesion of cartilage or subchondral bone, and are clinically manifested as pain, disability and joint stiffness, and the destruction of osteochondral tissue of patients with osteoarthritis cannot perform effective self-repair.

Tissue engineering scaffolds are thought to allow for the repair and reconstruction of defective tissue through cell-scaffold interactions and associated signaling molecules. Today, various technologies can achieve biomimetics of the extracellular matrix (ECM) and provide suitable 3D network microstructure for cell adhesion, proliferation and osteogenic differentiation. One of the main problems of the existing composite scaffold in bone tissue engineering is that the bonding strength of each layer of interface is insufficient, and the risk of peeling between the interfaces is increased.

Disclosure of Invention

The invention mainly aims to provide a preparation method of a composite scaffold with strong interface binding force for bone tissue engineering.

In order to achieve the above object, the present invention provides a method for preparing a composite scaffold for bone tissue engineering, comprising the steps of:

preparing a PCL bracket, namely preparing the PCL bracket,

and (3) performing electrostatic spinning on the PCL scaffold to form the PCL-gelatin scaffold.

In some embodiments of the invention, the method further comprises the step of covering the surface of the PCL-gelatin scaffold with a thin layer of BMP-2 loaded gelatin.

In some embodiments of the invention, the thin layer of gelatin is coated on the PCL-gelatin scaffold by plasma surface modification.

In some embodiments of the invention, the gelatin is crosslinked via a silane coupling agent.

In some embodiments of the invention, the PCL-gelatin scaffold is obtained by electrospinning a mixed spinning solution containing PCL and gelatin on the surface of the PCL scaffold.

In some embodiments of the invention, the PCL: the weight ratio of gelatin is 8: 2.

In some embodiments of the invention, the mixed spinning solution further comprises GPTMS.

In some embodiments of the invention, the GPTMS is 20% by weight of the mixed spinning solution.

In some embodiments of the invention, the PCL scaffold is modified with PDA and then electrospun.

In some embodiments of the invention, the PCL scaffold is printed by an FDM process.

In the composite scaffold prepared by the preparation method, the PCL scaffold is used for simulating a cancellous bone structure, the PCL-gelatin scaffold is used for simulating a subchondral bone structure, and the interface bonding strength between the PCL scaffold and the PCL-gelatin scaffold is remarkably improved. The results of corresponding physical and chemical property detection and mechanical analysis show that the composite bracket has more excellent performance. In vivo and in vitro experimental results show that the composite scaffold is beneficial to the synchronous repair of osteochondral defects and has potential clinical application value.

Drawings

FIG. 1 is an FE-SEM micrograph of PC scaffolds (a-c) and PCD scaffolds (d-f). (vertical (a, b, d, e), cross-sections (c and f)).

FIG. 2 is an FE-SEM micrograph of a multilayer scaffold. Electrostatic spinning of PCL-gelatin on PCD scaffolds (a, b), multilayer PCDE scaffolds (c-e) vertical orientation (c) and cross section (d, e).

Fig. 3 shows the average pore size (a) and average fiber diameter (b) of the 3D printing scaffold.

FIG. 4 is an FTIR spectrum (a) of the starting material and the holder; and (b) a synthetic process and an idea of the scaffold.

FIG. 5 shows stress-strain curves (a), compressive strength (b), elastic modulus (c) and bonding strength (d) between different types of stent interfaces for PC stents, single-layer PCD stents and multi-layer PCDE stents.

Fig. 6 is the water contact angle (a) of the PC scaffold, the water contact angle (b) of the single layer PCD scaffold, the water contact angle (c) of the PCDE scaffold, and the absorption capacity (d) of the PBS solution at different test time points for different types of scaffolds.

FIG. 7 is an FM-SEM micrograph of the surface of the scaffolds after soaking the different scaffolds in PBS solution and PBS-pancreatin solution for 6 weeks. In FIG. 7, a, b, e, f, i, and j represent water biodegradation; c, d, g, h, k, and l represent enzymatic biodegradation. PC scaffolds (a-d), single-layer PCD (e-h) and multilayer PCDE scaffolds (i-l).

FIG. 8 is the rate of aqueous biodegradation (a) and enzymatic biodegradation (b) of PC scaffolds, PCD scaffolds, PCDE scaffolds in PBS solution and PBS-pancreatin solution.

Fig. 9 shows the surface mineralization deposition form of hydroxyapatite on the surface of a PC scaffold (a, b), a single-layer PCD scaffold (c, d) and a multi-layer PCDE scaffold (e, f) after being soaked in a 10 xSBF solution for 12 hours, and fig. g shows the schematic diagram of the surface mineralization deposition form of hydroxyapatite on the surface of the scaffold.

FIG. 10 shows FTIR spectrum (a) and XRD spectrum (b) of surface hydroxyapatite of PC scaffold, PCD scaffold and PCDE scaffold after being soaked in 10 XSBF solution for 12 hours.

FIG. 11 shows adhesion of rBMSCs to the surface of a PC scaffold (a), a single layer PCD scaffold (b), and multiple layers of PCDE (c) after 3 days of co-culture with the scaffold.

FIG. 12 shows the staining of BMSCs after 1, 3, 5 days of co-culture (a, d, g, j), control (a-c), PC scaffolds (d-f), single-layer PCD scaffolds (g-i) and multi-layer PCDE scaffolds (j-l), and details of the proliferation of rBMSCs on different types of scaffolds are shown in (m).

FIG. 13 shows ALP activity (a) and OC protein expression (b) after 3, 7, and 14 days of co-culture of rBMSCs with different types of scaffolds.

Detailed Description

For patients who cannot perform effective self-repair of bone tissue, in the related art, the growth and repair of bone tissue are assisted by implanting a bone tissue engineering scaffold. In order to be able to repair and reconstruct bone tissue in different parts, composite scaffolds of multilayer composite materials can be prepared with different techniques. However, the interface bonding strength between different composite material layers of the existing composite support made of multiple layers of composite materials is insufficient, and the risk of interface peeling is increased due to the insufficient interface bonding strength.

In view of the above problems, the present application provides a method for preparing a composite scaffold for bone tissue engineering, comprising the steps of:

preparing a PCL bracket, namely preparing the PCL bracket,

Experiments prove that the composite scaffold obtained by forming the PCL-gelatin scaffold on the PCL scaffold in an electrostatic spinning mode has the bonding strength of about 5.4Mpa (figure 5d) between internal interfaces and high overall strength, and the bonding strength between the interfaces cannot be reduced even if the surface of the composite scaffold is subjected to plasma surface modification. At the same time, the composite scaffold can still provide appropriate hydrophilicity, absorption capacity, hydrolytic and enzymatic biodegradation capacity and bioactivity, and simultaneously promote the repair and reconstruction of bone tissues.

The present invention will be further described with reference to examples.

2-1-materials

Polycaprolactone (PCL) fibers (diameter 1.75mm) were purchased from Esun, inc (shenzhen, china). Dopamine hydrochloride (dopa-HCl, Mw 189.64g/mol), (3-glycidopropyl) trimethoxysilane (GPTMS, Mw 236.34s g/mol), sodium hydroxide (Mw 40g/mol) and 2,2, 2-trifluoroethanol (TFE, Mw 100.04g/mol) were purchased from Rhawn, Inc. (Shanghai, China). Bone morphogenetic protein-2 (BMP-2) was purchased from Peprotech, Inc. (New Jersey, USA). Gelatin (Mw 40-50kDa), isopropanol (99.99%, Mw 60.10g/mol), ethanol (Mw 46.07), glutaraldehyde (25% Mw 100.12 gr/mol), tris (Mw ═ 121.14g/mol), phosphate buffered saline (PBS, tablet, pH 7.2-7.4), sodium chloride (NaCl, Mw 58.44g/mol), potassium chloride (KCl, Mw 74.55g/mol), sodium bicarbonate (NaHCO) (NaHCO, Mw 60.10g/mol), sodium bicarbonate (NaHCO)₃Mw 84.01 g/mol), calcium chloride dihydrate (CaCl)₂·2H₂O, Mw 147.01g/mol), magnesium chloride hexahydrate (MgCl)₂·6H₂O, Mw 203.30 g/mol), monobasic sodium phosphate (NaH)₂PO₄·H₂O, Mw 137.99g/mol) and p-nitrophenyl phosphate (p-NPP, tablets) were purchased from Sigma-Aldrich, Inc. (Mo., USA). Hydrochloric acid (HCl, 37%, Mw 36.46g/mol) was purchased from Sinopharm, Inc. (Shanghai, China). DMEM medium and Fetal Bovine Serum (FBS) and penicillin-streptomycin solutions were purchased from Gibco-BRL Life technologies, Inc. (New York, USA). The alkaline phosphatase activity detection kit was purchased from Beyotime biotechnology limited (shanghai, china). BCA (bisquinolinecarboxylic acid) protein assay kit was purchased from Thermo Fisher technologies, Inc. (Massachusetts, USA)Country). The OC protein enzyme-linked immunoassay kit is purchased from Sino Biotechnology Ltd (Beijing, China). Deionized water is used for preparing the relevant solution. All reagents were used directly without further purification.

2-2-preparation of Single-layer 3D printing support

The scaffold modeling was performed using simple 3D software (version 4.0.1) and PCL scaffolds were prepared using an FDM-3D printer model HORI Z500. The molten filaments were printed in a lattice-like form using a nozzle of 200 μm size, and the filling rate was set to 0.14 mm/s. The temperature of the printer bed was set to 20 deg.C and the temperature of the nozzles was set to 115 deg.C. The printing angles were set at 0 ° and 90 °, and PCL holders were printed with a layer thickness of 0.2 mm. Finally, the prepared stent was subjected to a drying treatment in an oven (DHG-9050A) set at 36.0. + -. 0.5 ℃ for 24 hours. The single-layer PCL bracket is called a PC bracket for short. And (3) soaking the printed single-layer PC stent in a dopamine hydrochloride solution (working solution with the concentration of 2mg/ml is prepared by using a 10mM Tris solution), and fully stirring at room temperature to prepare the PDA modified single-layer PCL stent, which is called a PCD stent for short.

2-3-preparation of multilayer 3D printing support

And covering a PCL-gelatin layer prepared by an electrostatic spinning technology on the bottom PCD support. A mixed solution was prepared by adding PCL and gelatin to a 10% w/v TFE solution at a ratio of 8: 2. After being sufficiently dissolved, GPTMS was added to the above polymer solution in a proportion of 20 wt% and sufficiently stirred for 2 hours. Printing was performed at a rate of 3ml/h and a voltage of 5kV using the above mixed solution as a raw material via an electrospinning technique, with a distance between the nozzle and the PCD holder set to 13 cm. All samples prepared were dried at 36 + -0.5 deg.C and overnight. A three-layered scaffold was prepared by coating a thin layer of gelatin-BMP-2 on the surface of the scaffold by oxygen plasma modification technique. The stent was exposed to oxygen (0.9 l/h) at a voltage of 100W and a pressure of 0.76mbar for 90 seconds. After 45 minutes, the stent was immersed in the gelatin-BMP-2 solution for 30 minutes. Finally, the scaffolds were rinsed with deionized water and dried overnight at 36 ± 0.5 ℃. The gelatin-BMP-2 solution had 2% w/v gelatin and 1. mu.g/ml of BMP-2 dissolved therein. The above-described stent is simply referred to as a PCDE stent.

2-4-characterisation

A layer of thin gold is sprayed on the surface of the support, and the microstructure of the support is scanned by using an FE-SEM under the condition of 3kV voltage. The pore size and the fiber diameter of the stent are obtained by counting FE-SEM scanning images through software.

The stress-strain curve and the compressive strength of the bracket are detected by a crosshead of an instrument under the conditions of a set speed of 0.5mm/min and a pressing size of 100N. The bonding strength between the interfaces in the multilayer scaffold was measured by a tensile strength test system, and the displacement was set at 1 mm/min.

The hydrophilicity of the scaffold was detected by the sitting drop method under room temperature conditions (SL200KS, USA). The scaffold was evaluated for its ability to absorb PBS solution after immersion in 50ml PBS solution (37. + -. 0.5 ℃). The dry weight of the scaffold is W₀And the wet weight of the scaffolds after 2, 4, 6 and 24 hours soaking in PBS was recorded. And the absorption capacity of the scaffold to the PBS solution was calculated by the following formula 1.

Swelling ratio(％)＝[W-W₀/W₀]×100 (1)

The in vitro degradability of the scaffolds was tested by immersing the scaffolds in 50ml PBS solution (37. + -. 0.5 ℃) for 6 weeks, recording the initial weight of the scaffolds, and weighing and recording after 1 sample was taken every week. The water biodegradability of the scaffold was calculated using the following equation 2.

Biodegradation ratio(％)＝|[W-W₀/W₀]|×100 (2)

The enzyme biodegradability of the scaffold is determined by immersing the scaffold in 50ml PBS (37. + -. 0.5 ℃) and adding 4mg/ml trypsin thereto for 6 weeks. The initial weight of the scaffold was recorded and 1 sample taken every week was weighed and recorded and its enzymatic biodegradability evaluated using equation 2 above.

After the scaffold was immersed in 50ml of 10 XSBF F solution (37. + -. 0.5 ℃) for 12 hours, the scaffold was scanned using FE-SEM to confirm that a layered structure formed by hydroxyapatite mineralized deposition could occur on the surface of the scaffold, and the components of the layered structure on the surface of the scaffold were detected using ATR-FTIR technique and XRD technique. The incidence angle of X-ray in XRD is selected to be 2 theta (10-80 degrees), Cu-Ka (lambda 1.5418A degrees) ray is selected, and the detection result is compared with JCPDS standard to determine the crystal type.

2-5-cell-scaffold interaction

Bone marrow mesenchymal stem cells (rBMSCs) were used in vitro experiments. rBMSCs were seeded on the surface of sterile scaffolds at a density of 104/ml and in DMEM medium containing 10% FBS, 100U/ml penicillin-streptomycin at 37 + -0.5 deg.C, 5% CO₂After 3 days of culture in the environment, washing with PBS solution for 3 times, fixing with glutaraldehyde solution, performing gradient dehydration with ethanol solution of different concentrations, and soaking in isoamyl acetate-ethanol (1:1) for 10 minutes. The adhesion of cells to the surface of the scaffolds was evaluated by FE-SEM.

After 1, 3, 5 days of CO-culture of the cells with the scaffolds, live and dead cells were stained using PBS containing 24. mu.M calcein and 4.5. mu.M propidium iodide at 37. + -. 0.5 ℃ with 5% CO₂After incubation at 95% humidity for 30 minutes, the cells were assessed for viability using a fluorescence microscope after washing with PBS solution. After the rBMSCs cells were co-cultured with the scaffolds for 1, 3 and 5 days, the proliferation of the cells was evaluated using CCK-8 kit. At 5X 10⁴BMSCs were seeded on the surface of scaffolds at a cell density of/ml, and after incubation of cells with CCK-8 working solution (10% CCK-8 solution, diluted in serum-free DMEM) for 4 hours, 100ul of the reaction solution was pipetted and added to a 96-well plate and absorbance was measured at 450 nm.

Osteogenic behavior of the rbmscs cells ALP activity was assessed by ALP activity assay kit. After 3, 7 and 14 days of co-culture of the cells and the scaffold, the cells were lysed using RIPA lysate, centrifuged at 2000rpm for 10min, added with p-NPP solution for 30 min, and absorbance was measured at 405 nm. ALP protein concentration was detected using BCA protein detection kit.

After the cells and the scaffold are cultured for 3 days, 7 days and 14 days, the concentration of osteocalcin is detected by using an OC enzyme linked immunosorbent assay kit according to the use instruction, and an OC protein standard curve is drawn through the standard substance. The control group contained no scaffold sample.

2-6-statistical analysis

The results were repeated 5 times and recorded in EXCEL 2016 software. The results of 5 tests were presented as mean ± standard deviation and recorded in EXCEL 2016 software. And judging whether the difference between different results has statistical significance by using SPSS software, and considering that the difference has statistical significance when P is less than or equal to 0.05.

3-1-surface morphology characterization

The microstructure of the scaffold is critical to its physicochemical properties, mechanical properties, biological behavior and may even affect the regenerative reconstruction of damaged tissues. To date, 3D printed stents can achieve fine control over their geometry and microstructure, including pore size, pore shape, porosity, internal connectivity, etc., to meet the needs of target tissue repair and reconstruction. Figure 1 shows a microscopic image of a PDA-decorated PCL multilayer stent processed with FDM technology. The visible fibers in the image are interwoven perpendicular to each other to form a network, which is formed by printing a fused filament structure layer by layer on a machine tool.

In addition, the microscopic image shows that interconnected pore structures exist inside the stent, and meanwhile, the 3D printing stent well restores the size of the previous 3D design file without obvious shrinkage. In FIG. 1(d, e) we see the PDA modified PCL scaffold, during which the scaffold was first soaked in Tris-containing alcoholic solution, followed by dopamine hydrochloride addition. After 24 hours, dopamine hydrochloride is coated on the surface of the stent to form a thin layer structure due to the self-polymerization effect of dopamine hydrochloride, and the smooth surface of the PCL is fully covered. The PC stand in this application serves as a substrate to which the PDA is attached.

The porosity of the polymer scaffold is a critical factor for cell adhesion, migration, differentiation, and vascularization. The porosity is controlled to be 50-60%, and a proper condition is provided for the repair and reconstruction of bone tissues on the premise of ensuring the mechanical stability. Other parameters that may affect the 3D printing support include the size, distribution of the pores and fibers. The average pore size of the PC scaffold and the single layer PCD scaffold were 448.2 + -21.37 μm and 429 + -44.02 μm, respectively, while the diameter of the fiber in the PC scaffold was 355.8 + -22.76 μm, and the fiber diameter increased to 388.4 + - μm through PDA modification, as a result, it can be seen in FIG. 3. However, the above-mentioned differences between pore size and fiber diameter are not statistically significant, and the increase in fiber diameter may be attributed to coating of the PDA on the fiber surface. Other research results show that the pore size of 500 mu m in 100-. In the application, FE-SEM microscopic image results show that pores of the PC scaffold and the single-layer PCD scaffold are beneficial to proliferation and differentiation of stem cells and expression of genes and proteins, and finally the repair and reconstruction of osteochondral defects are realized.

The 3D printed scaffold can be used to promote the regeneration of cancellous bone, and for the repair and reconstruction of osteochondral defects, the scaffold should also be able to simulate subchondral bone as well as cartilage simultaneously. Therefore, the PCL-gelatin fiber layer is used for simulating subchondral bone, and the effect of increasing the bonding strength between interfaces is achieved. The electrospinning technique is a technique commonly used for preparing biomimetic extracellular matrices, and can achieve porosities in excess of 90%. Fibers on a nanometer or micrometer scale facilitate cellular infiltration, proliferation, and migration. In addition, hydrogels based on gelatin form a hydrophilic polymer network, and are largely similar in composition to the natural cartilage component, and thus can be used as a simulated cartilage layer. According to related researches, the surface modification and the loading of macromolecular protein can be beneficial to the directional growth, proliferation and differentiation of cells. In addition, treatment of hydrogels by gas plasma technology is an excellent strategy to improve the durability and stability of hydrogels. Therefore, the gelatin-BMP-2 hydrogel is directly fixed to the surface of the substrate by the oxygen plasma technique. FIG. 2 shows the microscopic morphology of the electrospun layer before and after coating the gelatin-BMP-2 hydrogel layer by FE-SEM images. Microscopic images revealed filaments of uniform morphology and no abnormal beaded structure with an average diameter of 2.27. + -. 0.46. mu.m. While the diameter of the fibers increased to 3.36 + -0.80 μm after the gelatin-BMP-2 hydrogel was fixed, the increase in diameter was favorable for increasing the mutual communication between the fibers and decreasing the ineffective area inside the stent (see FIG. 3 for details). The fixation of gelatin-BMP-2 hydrogel on the surface of the base material has an adverse effect on the uniformity of the fiber size. In addition, cross-sectional images of multiple layers of PCDE show that the fibers form a complete bond with the 3D printed stent (fig. 2(D, e)). This phenomenon may be closely related to the chemical composition structure of the fiber. The PCL layer can be used for simulating a cancellous bone layer, and the gelatin layer can be used for simulating a cartilage layer.

3-2-chemical characterization

The raw materials and the scaffold were tested by FTIR technique and the results are shown in fig. 4 a. The idea and process of scaffold synthesis are shown in detail in fig. b.

In the PCL material, CH₂The peak generated by the group stretching vibration can be seen at 2946cm^-1And 2865cm^-1Is observed. And at 1721cm^-1，1294cm^-1，and 1238cm^-1The peaks observed here are due to the stretching vibrations of C-O, C-C and C-O-C, respectively. For PDA, 3200-3500cm-1 peak resulted from stretching vibration of O-H group and N-H group was observed. The absorption peak generated by the stretching vibration of the N-H group can be 1603cm^-1Is observed. The peak values generated by stretching vibration of the N-H group and the C-O group in the amide group can be respectively 1511 cm and 1119cm^-1Is observed. The characteristic peak generated by the stretching vibration of the C-O-H group in the phenol group can be 1344cm^-1And 1285cm^-1Is observed. Characteristic peaks in gelatin can be observed at 3200-3600cm-1, suggesting stretching vibration of N-H groups and O-H groups. At 1633 and 1525cm^-1The peaks appear at (a) are due to stretching vibrations of the C ═ O and N — H groups in

amide groups

1 and 2. In addition, at 1230cm^-1The peaks appear in relation to the stretching vibrations of C-N and N-H in amide group 3. For GPTMS, at 2944 and 2812cm^-1Peak value at and CH in methoxy₃The stretching vibration of the group is related. Furthermore, deoxygenated water and CH in propyl chain₂The peak value of stretching vibration of the group can be 1195 and 1081cm^-1Is observed. GPTMSThe peak values of the epoxy resin in (1) are located at 1254, 910, 859 and 436cm^-1。

The characteristic peak of the PC bracket and the characteristic peak of the pure PCL raw material have better similarity, which indicates that the processing technology used by the application can not generate adverse effect on the chemical composition of the PCL raw material. In the application, a thin coating film is formed on the surface of a PCL support through spontaneous self-oxidation reaction of PDA in an alkaline medium. Dopamine hydrochloride forms dopamine quinone through autoxidation in an alkaline medium, then deprotonates at an amine group and further forms dopamine chromium through a subsequent Michael reaction. Finally, the interaction of o-quinone with o-catechol in 5, 6-dihydroxyindole was terminated to form PDA. Nielsen et al teach that the polymerization process of PDA enhances binding forces, including hydrogen bonding, bidentate chelation, coordination, and mixtures of monodentate and bridged bidentate bonds, to interact with the substrate. The spectrum of the PCD bracket shows a peak value related to PDA, which indicates that PDA is coated on the surface of a PCL substrate, the intensity of the peak value related to C-O, C-O, C-C is reduced on the PCD bracket, and the peak value generated at 1630cm < -1 > is related to the stretching vibration of aromatic rings and NH groups of the PDA. This phenomenon is probably due to a chemical interaction between the carbonyl group of PCL and the amine group of PDA.

In the multilayer PCDE scaffold, the GPTMS crosslinked electrospun gelatin was at 1160950 cm each^-1The peaks generated in (A) are associated with Si-O-Si groups and Si-OH groups. In this application, the amino group of gelatin interacts with the epoxy ring in GPTMS, which terminates the ring opening reaction of the epoxy group. After the phenomenon occurs, trimethoxy in the chemical structure of GPTMS undergoes hydration reaction, silanol groups are finally formed in the last step under the action of acid catalysis, and covalent bonds between Si-O-Si bonds are formed, so that condensation reaction of the silanol groups and mutual crosslinking are caused to form gelatin. In the present application, the cartilage simulant layer is immobilized on the PCDE scaffold after oxygen plasma surface treatment. Oxygen plasma treatment affects the functional groups on the surface of the cartilage simulant layer by bombarding the chemical groups with free radicals or ionized species, resulting in the production of carbon radicals and unstable hydroperoxides. The peroxides interact with each other to form oxygen-containing functional groupsGroups including carbonyl, carboxylic acid, hydroxyl, and the like. Therefore, the peak intensity of the hydrophobic CH group decreases due to the cleavage of the CH bond, and the peak corresponding to the oxygen functional group (C ═ O) contained in the oxygen-rich surface increases, and the spectrum shows 2946cm^-1Decrease in near peak intensity and 1721cm^-1The above view is also confirmed by the peak intensity in the vicinity. In addition, the active free radicals on the surface of the fiber and the gelatin quickly form interaction, and the fixation condition of the outermost layer on the surface of the stent is improved. Covalent coupling of methyl groups of the immobilized solution and hydroxyl groups of the electrospun layer and hydrogen bonding of the oxygen-rich surface to the gelatin-based solution is possible, 3300cm in the spectrum^-1The peaks appearing here (in relation to the hydroxyl groups) also confirm the above view.

3-3-mechanical properties

A range of parameters, including porosity, pore size, pore shape, and internal interactions and chemical properties, all affect the mechanical properties, stability, and integrity of the scaffold. In order to ensure that the scaffold can bear the external force action while adhering, migrating, proliferating and vascularizing cells on the scaffold, a balance point needs to be found between the porosity and the mechanical property. Therefore, the stress-strain curve, compressive strength, elastic modulus, and bonding ability between interfaces of the stent need to be measured, and the results are detailed in fig. 5. The stress-strain curve of the PCDE scaffold is significantly higher than other scaffold samples, so the energy it can absorb before the PCDE scaffold breaks also matches its toughness strength. Compared with the PC scaffold, the compressive strength and the elastic modulus of the PCD scaffold are respectively 1.07 times and 1.06 times of those of the PC scaffold, however, the difference between the PCD scaffold and the PC scaffold in the above indexes has no statistical significance.

Although PCL is itself a polymeric material with high strength, the improvement in strength of the single layer PCD scaffold in this application is more relevant to nano-scale PDA, smaller pore size, thicker fibers at the surface of the scaffold. In a similar study, PDA modifications may enhance stent durability and stability. In contrast, the multilayer PCDE scaffold exhibited a more significant compressive strength than the PC and PCD scaffolds, providing 1.87 and 1.74 times compressive strength, respectively, and the single layer scaffold was significantly less resistant to deformation and durability than the multilayer scaffold. The mechanical strength of the scaffold is similar to that of natural cartilage tissue (0.24-1MPa) and subchondral bone (30-50 MPa). The elastic modulus of the material can generate a stress shielding effect in the process of bone tissue regeneration, and in order to solve the problem, the compressive elastic modulus of the multilayer PCDE scaffold is 1.80 times that of the PC scaffold and 1.72 times that of the single-layer PCD scaffold respectively. The higher mechanical stability of the multilayered PCDE scaffold is a factor affecting cell morphology, cell activity and tissue remodeling. One of the main problems of the multi-layer stent is that the interface bonding strength of each layer is insufficient, increasing the risk of peeling between the interfaces. The bond strength between the stent interfaces is presented in fig. 5 (d). In the application, the bonding strength between the stent interfaces before and after the plasma modification technology treatment of the 3D stent processed by electrostatic spinning is evaluated. The results show that the gelatin-BMP-2 gel layer after fixation does not adversely affect the bonding strength between the cancellous bone interface and the subchondral bone interface. The stent of the present application can have a bond strength between the internal interfaces of the stent about 250 times higher for obtaining suitable in vivo biomechanical stability. Therefore, the multilayer composite scaffold in the application is more beneficial to the regeneration and reconstruction of osteochondral defects.

3-4-absorption and biodegradation capabilities

The interaction between the scaffold and water molecules is influenced by the microstructure and chemical components of the scaffold, good hydrophilicity and the pore structure with the communicated inner parts are beneficial to the full absorption of water, and the interaction between the scaffold and the water molecules influences the regeneration process of tissues. Therefore, due to the improvement of water-scaffold interaction, the sufficient formation of capillary vessels inside the scaffold is more beneficial to promote the improvement of cell function, the increase of body fluid flow and nutrient transfer, the secretion of ECM, and finally the formation of new tissues. In addition, poor fluid flow results in cellular dystrophy and a lack of adequate vascularization, ultimately resulting in poor tissue repair and reconstruction. Previously published results suggest that the hydrophilicity of such structures, particularly the surface hydrophilicity, leads to the adhesion of proteins such as laminin. Therefore, the protein-rich surface can provide sites for cells to adhere to the surface of the scaffold, and promote cell adhesion, proliferation and differentiation, which has important significance in the regeneration of injury.

The hydrophilicity of the scaffold was evaluated by measuring the contact angle by the sessile drop method, and the results are shown in detail in FIG. 6 (a-c). The water contact angles of the PC stent, the single-layer PCD stent and the multi-layer PCDE stent are 86.41 +/-1.05, 52.63 +/-1.48 and 84.02 +/-0.67 degrees respectively. Thus, the reduction in water contact angle of the PCD scaffold was due to the improvement in non-hydrophilicity of the PCL scaffold after PDA modification. In other studies it was found that the hydrophilicity of PLA or PCL scaffolds modified by PDA was significantly increased. Needless to say, the hydroxyl and amine groups in PDA may explain the above phenomenon. Furthermore, the increase in contact angle of the multilayer structure may be due to the hydrophobicity of GPTMS as a modifier of gelatin. Furthermore, the reduced number of hydrophilic functional groups (amine and hydroxyl groups) in gelatin due to their involvement in GPTMS interactions may be another possible reason for the slight decrease in hydrophilicity of the multi-layered scaffold. However, other benefits of GPTMS, including bioactivity or osteogenic properties, are sufficient to counteract the adverse effects of GPTMS.

The water absorption capacity test result of the stent is shown in detail in FIG. 6 (d). Different scaffolds showed varying degrees of propensity to absorb water, consistent with the hydrophilicity of the different scaffolds themselves. The results show that each of the different scaffolds exhibited time-dependent behavior, with increasing experimental time, and the water uptake of each scaffold material gradually increased up to 24 hours. Thus, all scaffolds showed high water absorption (12-28%) in the first 2 hours. After 24 hours, the water absorption capacity of the PC stent, the single-layer PCD stent and the multi-layer PCDE stent is respectively 20.78 +/-3.53%, 33.67 +/-3.67% and 25.20 +/-1.05%. Although both PCL and GPTMS are hydrophobic in nature, the scaffold's own good specific surface area, and the presence of interconnected tubular pores between the 3D printed scaffold interfaces can significantly increase the ability of the scaffold to absorb PBS. Other efforts to study unidirectional pore structures have also suggested the above. Thus, this microstructure may increase moisture absorption by facilitating body fluid exchange. This phenomenon is better for cell nutrition and promotes the repair and reconstruction of damaged tissues. From the results obtained, it can be concluded that the water uptake capacity of the PC scaffold after modification with nano-scale PDA increases significantly, and therefore the water uptake at different test time points is significantly higher than that of the other groups. This phenomenon is associated with the abundance of hydrophilic functional groups, including catecholamines and hydroxyl groups, in surface modified PDA's, which can form hydrogen bonds with water molecules. Because GPTMS is contained in the chemical components of the scaffold, the water absorption capacity of the multi-layer scaffold is reduced compared with that of a single-layer scaffold. GPTMS itself is not cytotoxic and has improved ability to promote bone differentiation and increased gelatin internal crosslinking to improve hydrogel stability with minimal adverse effects on hydrophilicity. Therefore, the increased hydrophilicity of the PCDE scaffold is closely related to the hydrophilic groups abundant in PDA, gelatin, compared to the PC scaffold. Compared with the PC scaffold, the PCDE scaffold treated by the oxygen plasma has the advantages that the enrichment of oxygen functional groups and the improvement effect of the hydrophilicity of the PCDE scaffold can be realized on the surface of the PCDE scaffold. In short, both single-layered and multi-layered scaffolds have good hydrophilicity, and the optimal hydrophilicity of the PCDE scaffold determines its superiority in clinical applications.

The degradability of a scaffold and the biocompatibility of its degradation products are important parameters in the field of tissue engineering. A suitable scaffold should provide a matrix for cell attachment and function. Thus, a rapidly degradable or non-degradable structure does not achieve the end purpose. The microstructure (pore size, pore shape, internal connectivity, surface morphology and crystallinity, etc.), chemical composition (hydrophilicity, glass transition temperature, molecular mass, etc.), environmental conditions (PH, temperature, pressure, etc.), and additive components (acidity, basicity, monomers, solvents, etc.) affect the degradation rate of the material. FIG. 7 is an FE-SEM micrograph of different types of scaffolds after being soaked in PBS and PBS-pancreatin culture solution for 6 weeks, respectively.

As shown in fig. 7, in the enzyme-containing PBS solution, more additional porosity was present in the 3D printed scaffold as biodegradation occurred, while the integrity of the scaffold was destroyed as the fibers disintegrated. In addition, the arrangement of pores and fibers is disrupted. The Cai et al think that the degradation process of the enzyme biodegradable stent in the enzyme-containing PBS solution is similar to the in vivo degradation process, so the stent has better application prospect in human body.

Fig. 8(a, b) shows the biodegradation rate of the scaffold over 6 weeks. The degradation curve indicates that the stent is degraded at a relatively constant rate. It is known that the degradation of the scaffold in PBS solution and enzyme-containing PBS solution is less than 40%, reflecting that the scaffold can still maintain a certain stability. In the process of bone tissue repair and reconstruction, callus can be seen to appear after 2 weeks, and full callus formation can be seen after 4-7 weeks. In addition, the bridging callus and periosteum are involved in the process of bone tissue repair and reconstruction, and the process of shaping the new bone tissue begins within 2 months. In addition, the proliferation and growth of cells require 2 months, fibrocartilage is formed after 4 months, cartilage is completely repaired after 6 months, and extracellular matrix synthesized and secreted is gradually absorbed. This requires that the degradation rate of the scaffold should be matched to the rate of cell adhesion, proliferation, vascularization and formation of new bone tissue in the body.

Various factors can affect the durability of the stent. Generally, the interconnected pore structure inside the scaffold provides a plurality of binding sites for the degradation process of the scaffold, which facilitates the exchange of body fluid and reduces the mechanical strength of the scaffold. However, the stent internal tubular microstructure is one of the contributing factors affecting the durability of 3D printed stents. Another study shows that the integrity of the stent can be ensured by a regular tubular pore structure, so that the phenomenon of sudden stent disintegration caused by the existence of randomly distributed pores is prevented.

The chemical composition of the scaffold can have an effect on its rate of biodegradation. Whereas the biodegradation rate of hydrophilic materials is significantly higher than that of non-hydrophilic materials. The chemical composition of the constituent osteoid and subchondral osteoid layers is typically PCL material. While the PCL material itself is a semi-crystalline, semi-hydrophobic polymer. Related researches report that the aquatic degradability of plants in the PCL is low. In addition, the PCL-containing scaffold can maintain the stability and integrity of the scaffold for a long time due to the characteristics of the PCL-containing scaffold. Therefore, the crystalline phase component can effectively resist rapid disintegration. However, acidic degradation products generated when the material biodegrades in vivo may further increase the rate of scaffold mass loss compared to in vitro environments. As the hydrophilicity and water absorption capacity of the material increase, the biodegradation rate of the material also increases.

In the present application, the increased hydrophilicity of the single layer PCD scaffold due to the presence of PDA modification resulted in a PCD scaffold with a significantly higher biodegradation rate than the PC scaffold. This is due to the rich hydrophilic groups in PDA which leads to a significant increase in the propensity of the PCD scaffold to interact with water molecules.

Thus, the stronger the interaction between the scaffold and the water molecule, the more likely it is to cause damage to the hydrolysable groups in the material. Research on biodegradation of PDA in vivo shows that under the phagocytosis of cells, nicotinamide adenine dinucleotide phosphate oxidase can be produced in a human body, and further microorganisms, free radicals and oxygen are produced. All of the above factors will affect the biodegradation of PDA within 8 weeks. Other related studies indicate that degradation of PDA in vitro is dependent on a hydrolytic mechanism. Previous studies of PDA nanospheres in this application demonstrated that PDA disintegration within 4 weeks in vitro was due to disruption of the-NH group and disruption of C-C, C-O. Furthermore, the destruction of the-COH group can also be observed during the degradation of PDA.

In contrast, the multi-layered scaffold biodegrades more slowly than the single-layered scaffold, which leads to reduced biodegradation due to the hydrophobicity of GPTMS in subchondral bone and cartilage-like layers and the decreased density of hydrophilic functional groups of the polymer (gelatin) after crosslinking, thereby extending the durability of the scaffold. The in vitro hydrolysis of gelatin-GPTMS containing material also follows the mechanism of water biodegradation. When the nanofiber size is less than 300 microns, biological disintegration will occur gradually with the fiber direction. During biodegradation of the second and third layers of the scaffold, as the Si-O-Si and Si-OH bonds are broken, the release of silicon ion groups, amino acids and metal ions results. Hydrophilic functional groups in gelatin, including amine and hydroxyl groups, provide active sites for rapid hydrolysis and biodegradation.

Nevertheless, PCDE scaffolds showed higher mass loss rates than PC scaffolds. This phenomenon may be due to the presence of the uppermost layer containing gelatin. The amine and hydroxyl groups in gelatin can increase the rate of biodegradation by increasing hydrophilicity. Another possible reason is that oxygen-containing functional groups generated by plasma modification can increase the rate of biodegradation by further increasing hydrophilicity.

3-5-biological Activity

Hydroxyapatite biomineralization deposition is the key point in the bone tissue repair and reconstruction process, the scaffold is soaked in a 10 xSBF solution for 12 hours, and the scanning evaluation of the hydroxyapatite biomineralization deposition condition on the surface of the scaffold is carried out by FE-SEM (figure 9), and the FTIR spectrum and the XRD spectrum are detailed (figure 10). In the application, the 10 xSBF solution is used for evaluating the biological activity of the scaffold, and compared with the conventional SBF solution, the 10 xSBF solution with the novel formula can improve the mineralization and deposition speed of hydroxyapatite on the surface of the scaffold. The present application utilizes NaH in the present application₂PO₄·H₂O instead of K₂HPO₄·3H₂O and changing HCO^3-And Cl^-The results are shown in Table 1.

TABLE 1 comparison of ion concentrations of 10 XSBF solution, regular SBF solution and human plasma

FE-SEM microscopic images prove that hydroxyapatite forms granular mineralized crystals on the surface of the scaffold after the scaffold is soaked in a 10 xSBF solution for 12 hours. Then, as the degree of mineralized deposition increases, the surface of the whole stent forms a layered mineralized deposition structure. Therefore, the nano-like mineralized particles uniformly cover the surface of the single-sided PCD stent and the multi-layered PCDE stent. However, only a small number of hydroxyapatite mineralized crystals were visible on the surface of the PC scaffold (FIG. 9 (a-f)).

The results obtained by using PDA and GPTMS in the chemical composition of the composite scaffold demonstrate that PDA and GPTMS improve the bioactivity of the scaffold. Other studies have shown that hydroxyapatite-like layers may be developed on scaffolds containing PDADipole-generating interactions and ion conduction. The surface of the PC scaffold is modified by using PDA, and the negative charges and catecholamine carried by the surface of the composite scaffold can further influence the mineralization and deposition of hydroxyapatite. Charge repulsion terminates with the absorption of metal ions in a 10 x SBF solution, thereby reducing the energy of the overall system. Soaking the PCD holder in a 10 xsbf solution resulted in the release of H + from the hydroxyl groups contained in the PDA. Thus, activated surfaces and nucleation sites are generated, available for transfer at HCO 3-and HPO 42-on the stent surface, and stimulate the deposition of calcium ions on the stent surface. The high hydrophilicity of the single layer PCD scaffold plays a crucial role in promoting the mineralised deposition of hydroxyapatite on the surface of the scaffold. In the second and third layers of gelatin containing GPTMS, silanol groups promote the nucleation of the calcium phosphate layer and reduce the energy consumed by the system. Thus, the silanol groups electrostatically interact with the ions in the 10 x SBF solution to promote the formation of a hydroxyapatite-like layer. Furthermore, the surface negative charge generated by hydroxyl and carboxylic acid groups provides a suitable substrate for the adsorption of calcium ions. The deposition of calcium ions on the surface of the PCDE scaffold disrupts the ionic balance in the 10 x SBF solution, changing the surface charge. This phenomenon is accompanied by H in the carboxyl group⁺And phosphate is adsorbed on the surface of the stent by calcium silicate. The result of this interaction is the formation of octacalcium phosphate. Finally, after the stent is soaked in the solution, different ions are deposited on the surface of the stent, and hydroxyapatite mineralized deposition is further formed. Notably, the increased hydrophilicity of the gelatin and the plasma modification of the scaffold may facilitate the absorption of the 10 x SBF solution and the mineralized deposition of hydroxyapatite. The hydroxyapatite formation schematic is shown in fig. 9 (b). It is contemplated that the formation of hydroxyapatite on the surface of the scaffold facilitates the binding of the scaffold to surrounding tissue and may contribute to the reparative reconstruction of the defective tissue.

ATR-FTIR spectrum (detailed in FIG. 10(a)) of the mineralized and deposited structure on the surface of the stent can show the existence of chemical bonds related to hydroxyapatite, 3100-^-1Broad peak and 1610cm^-1The peaks are related to the stretching vibration of O-H and water molecules in the mineralized sediment structure. 582cm^-1Peak of (d) and vibrational contraction of P-O in phosphateAre closely related. In addition, the peak value generated by C-O vibration contraction can be 819cm^-1Is observed. Another study demonstrated that the hydroxyapatite formed by the mineralized deposition of the material after soaking in a 10 x SBF solution was very close to the hydroxyapatite component formed under physiological conditions. The peak value generated by the stretching vibration of the carbonate group in the hydroxyapatite can be 1410-1455cm^-1Is observed.

The XRD pattern of hydroxyapatite (fig. 10(b)) confirmed the mineralized deposition of hydroxyapatite on the surface of the scaffold. A broad peak due to the amorphous phase of the polymer was observed at 2 theta angles (10-25 deg.). However, the deposition of the hydroxyapatite layer on the surface of the scaffold resulted in a decrease in the intensity of the amorphous peaks, and new peaks appeared at 27.27, 31.55, 39.65, 45.62, and 49.53 ° of the 2 θ angle, which correspond to the (210), (211), (310), (222), and (213) crystals, respectively. The increase of the peak intensity of the crystal indicates that the PDA and the GPTMS play a synergistic promotion role in the biomineralization process of the hydroxyapatite.

3-6-cell-scaffold interaction

It is necessary to perform in vitro cell experiments before the stent is implanted into a human or animal body. FE-SEM microscopic images of adhesion, proliferation, differentiation, activity, etc. of rBMSCs cells on the surfaces of PC scaffolds, PCD scaffolds, PCDE scaffolds after 3 days of co-culture of the monolayer scaffolds and the multilayer scaffolds are shown in detail in FIG. 11 (a-c). The results show that the scaffold provides suitable conditions for cell adhesion, reflecting that the porous microstructure of the scaffold in the present application is favorable for oxygenation, body fluid exchange and angiogenesis. In addition, scaffold biocompatibility plays a crucial role in cell behavior and its function. The related results show that the cells of the PC scaffold after the surface modification of the PDA can be fully extended on the surface of the scaffold. One possible reason for improving cell adhesion to the surface of the scaffold is that PDA helps prevent protein denaturation. In addition, catecholamine and hydroxyl functional groups in PDA can promote cell adhesion by interacting with integrins. In addition, it must be considered that PDA can provide a binding site for cell adhesion by binding to fibronectin in serum. Among them, the multilayered PCDE scaffold showed cells tightly adhered to the scaffold surface by the plating and filopodia, compared to other experimental groups. Hydrophilic functional groups in the gelatin chemical composition, such as amine, hydroxyl, and carboxylic acid, can promote cell migration. In addition, oxygen plasma treatment improves the hydrophilicity of the stent surface, thereby improving cell adhesion and migration. In addition, the modified scaffold surface can provide binding sites for integrin receptors, thereby facilitating cell spreading. In addition, BMP-2 may further improve cell adhesion, which results from improving the interaction between cells and the scaffold by binding to cell surface receptors.

Therefore, the scaffold has good biocompatibility because the cells are uniformly distributed on the surface of the scaffold after 1 day of culture. Here, a cluster of cells is gradually formed to increase the culture time, which is the first sign of cellular cartilage formation. It is speculated that the phenomenon can be influenced by a nutrient rich medium, lack of stress, and good cellular respiration. However, the density of viable cells decreased after 5 days of co-culture, and according to the results of CCK-8, the cell proliferation on the scaffold in the experimental group was not significantly different from that of the control group, while the cell proliferation on the scaffold in each experimental group was significantly increased after 3 days. After 5 days, the cell proliferation on the scaffold of each experimental group is reduced, but the survival rate of the whole cells is still over 80%, and other related researches also obtain similar results. There were significant differences between the sets of measurements at all time points tested. However, the cell-scaffold interaction of the monolayer PCD scaffold was significantly improved compared to the PC scaffold, thus confirming the supporting effect of PDA on cell growth. BMSCs were co-cultured with multi-layered PCDE scaffolds without a significant increase in cell number at 24 hours, while the number of viable cells increased significantly with prolonged culture time. One possible reason may be that the affinity of the scaffold for interaction with water molecules is reduced due to the addition of GPTMS. Thus, the decrease in scaffold hydrophilicity resulted in no significant increase in cell proliferation within the first 1 day. In addition, Si ions released by GPTMS are known to retard cell growth and proliferation. However, after 3 days of co-culture with the multi-layered PCDE scaffold, the multi-layered PCDE scaffold showed the highest cell survival rate and cell proliferation compared to other groups. This may be related to the interaction that occurs between the scaffold and the integrin. Thus. The cell behavior is altered, thereby promoting cell proliferation.

When properly stimulated, rBMSCs have the potential to differentiate into osteoblasts or chondrocytes. Lays a good foundation for the in vitro research of the bone cartilage regeneration. Accordingly, the present application evaluated the effect of different types of scaffolds on ALP and osteogenesis-related gene expression, as shown in fig. 13(a, b). ALP is a major marker expressed in osteogenic processes. The results of this application show that ALP activity levels were low after 3 days of co-culture of BMSCs with scaffolds, but ALP activity increased significantly with the increase of culture time. However, after 14 days of co-culture, the expression of ALP was rather decreased, but the activity was still significantly higher than that of the control group. Thus, the activity level of ALP was higher than that of PC scaffolds, whether single-layer PCD or multilayer PCDE scaffolds. Since ALP is expressed during ECM formation and maturation, the presence of bioactive factors may promote ALP activity. Therefore, PDA and GPTMS and their synergy affect cell differentiation and increase ALP activity. Other related studies have also demonstrated the potential of PDA and GPTMS to promote osteogenic differentiation. In addition, Si ions are contained in the chemical structure of GPTMS, although release of Si ions inhibits cell proliferation and activity of ALP.

OC protein is a non-collagenous protein and is the major biomarker for osteogenesis. It is closely related to new bone formation. Thus, the present application determines the expression level of OC protein. In the present application, the OC protein expression level in all experimental groups (PCD scaffold, PCDE scaffold) was significantly higher than that of the control group. In this application, OC protein measurements were higher in monolayer PCD scaffolds than multilayer PCDE scaffolds after 7 days of co-culture, although the difference between the measured concentrations was not statistically significant. But after 14 days of co-culture, the multilayer PCDE scaffold expressed OC protein significantly higher than the monolayer PCD scaffold. This phenomenon may be caused by the release of Si ions from the PCDE scaffold. In addition, the role of the osteoinductive signaling protein BMP-2 in improving the capacity of the scaffold to promote bone should not be overlooked. Another study showed that new bone formation increased significantly with BMP-2 release in the nanofiber scaffold.

According to in vitro experiment results, the application can think that the multilayer PCDE support can improve cell behaviors and promote the expression of osteogenesis related proteins, fully simulates osteochondral epimatrix, and is beneficial to the repair and reconstruction of bone defect parts.

The multilayer PCL-PDA-gelatin composite bracket loaded with BMP-2 is successfully constructed by combining a 3D printing technology, an electrostatic spinning technology and a low-temperature oxygen plasma surface modification layer fixing technology. The multi-layer composite scaffold is proved to be suitable for simultaneously promoting the repair and reconstruction of osteochondral defects by detecting the compressive strength, the interface bonding strength, the hydrophilicity, the absorption capacity, the hydrolytic and enzyme biodegradation capacity and the biological activity of the composite scaffold. The adhesion condition, proliferation condition and other results of the BMSCs on the scaffold prove that the multilayer composite scaffold has better biocompatibility. At the same time, ALP activity and OC protein expression also reflect the bone-promoting ability of the scaffold during tissue repair reconstruction. Therefore, the application provides a multilayer composite scaffold for integrated repair and reconstruction of osteochondral defects.

The embodiments of the present invention are merely illustrative, and not restrictive, of the scope of the claims, and other substantially equivalent alternatives may occur to those skilled in the art and are within the scope of the present invention.

Claims

1. A preparation method of a composite scaffold for bone tissue engineering is characterized by comprising the following steps:

preparing a PCL bracket, namely preparing the PCL bracket,

2. The method according to claim 1, further comprising the step of coating a thin layer of BMP-2-loaded gelatin on the surface of the PCL-gelatin scaffold.

3. The method according to claim 2, wherein the thin gelatin layer is coated on the PCL-gelatin scaffold by plasma surface modification.

4. The method according to claim 2, wherein the gelatin is crosslinked by a silane coupling agent.

5. The method according to claim 1, wherein the PCL-gelatin scaffold is obtained by electrospinning a mixed spinning solution containing PCL and gelatin on the surface of the PCL scaffold.

6. The method of claim 5, wherein the PCL: the weight ratio of gelatin is 8: 2.

7. The method of claim 5, wherein the mixed spinning solution further comprises GPTMS.

8. The method of claim 7, wherein the GPTMS is 20% by weight of the combined spinning solution.

9. The method of claim 1, wherein the PCL scaffold is modified with PDA and then electrospun.

10. The method of claim 1, wherein the PCL scaffold is printed by FDM process.