Multi-scale graded magnesium alloy stent as well as preparation method and application thereof

Abstract

The invention belongs to the technical field of biomedical materials, and discloses a multi-scale graded magnesium alloy stent as well as a preparation method and application thereof. Magnesium alloy is used as a base material, and a laser powder bed melting technology is adopted to obtain a base body support; the method comprises the following steps: chemically polishing a matrix support, performing high-temperature oxidation treatment to form an oxide layer on the surface, and performing hydrothermal reaction to form a hydrotalcite layer on the surface to obtain the multi-scale graded magnesium alloy support. Precise control of a macrostructure of the stent is realized by using an L-PBF technology, optimal design of mesoscopic pores of the stent is realized through structural design, a micron-sized coating formed by an oxide layer and a hydrotalcite layer on the surface of the stent can inhibit rapid degradation, and hydrotalcite nanosheets can improve surface roughness and hydrophilicity, so that adhesion and differentiation of osteoblasts are facilitated. The multi-scale design (macroscopic shape, mesoscopic pore, microcosmic coating and nano surface) of the stent effectively simulates the hierarchical structure of natural bone, and the biological adaptability of the stent is improved.

Description

Technical Field

[0001] The present invention relates to the technical field of biomedical materials, and in particular to a multi-scale graded magnesium alloy stent and a preparation method and application thereof. Background Art

[0002] Repair of critical bone defects has always been a major challenge in clinical orthopedics. Although considered the "gold standard", autologous transplantation requires an additional surgical site and increases patient morbidity, while allografts carry the risk of transmitting infectious diseases from the donor. In light of these issues, there is a great deal of effort to develop alternative approaches to treat bone defects. Bone implants can promote bone tissue repair without the problems associated with bone transplantation, and are expected to provide more effective and safer treatment for patients with bone defects.

[0003] Commonly used biomaterials for bone implants include polymers, ceramics, and metals. Metals are widely used as bone implant materials due to their excellent mechanical properties. Currently, commonly used metal materials for bone repair include titanium and its alloys, magnesium and its alloys, etc. However, titanium alloys are non-degradable in the human body and require secondary surgery to remove, which increases costs and patient morbidity. The release of toxic metal ions can cause inflammation, and excessive elastic modulus can lead to stress shielding effects. Magnesium alloys have become a promising alternative material due to their biodegradability and excellent osteoinductive properties.

[0004] In order to better match the shape of bone defects and promote bone tissue ingrowth, bone implants need to be designed with customized geometric shapes and porous complex internal structures. Traditional manufacturing technologies, such as milling and casting, are difficult to manufacture products with complex geometric shapes, while traditional porous material production methods, such as powder metallurgy, foaming, sintering, and chemical vapor deposition, cannot achieve stable and controllable pore shape and distribution, which usually leads to low mechanical properties, poor permeability, poor biocompatibility, and is not conducive to bone repair and regeneration. Additive manufacturing, also commonly known as "3D printing", is a manufacturing method that uses three-dimensional computer-aided design (CAD) data to stack materials layer by layer in sequence into three-dimensional objects. It has become an effective technology for developing customized bone implants. This technology can achieve precise design and manufacturing of implants that match the shape of the patient's bone defect, and can also produce complex internal structures such as porous scaffolds. Laser powder bed fusion (L-PBF) technology is an important additive manufacturing technology that is widely used in the production of orthopedic metal implants due to its high forming quality and forming accuracy.

[0005] The standard corrosion potential of magnesium is -2.37V, which is much lower than that of Zn (-0.76V) and Fe (-0.44V). Therefore, one of the main challenges facing magnesium-based metal bone implants is that the degradation rate in the body fluid environment exceeds the healing rate of bone tissue. The rapid degradation of magnesium in the human body can lead to a large amount of hydrogen accumulation and local pH increase, triggering inflammation and even tissue damage. The preparation of magnesium alloys by L-PBF usually involves repeated heating and remelting, which leads to a large amount of precipitation of secondary phases in the magnesium alloy, causing galvanic corrosion, further accelerating degradation. At the same time, the porous structure brings a larger specific surface area, which further shortens the degradation cycle of porous bone implants in the body fluid environment. Therefore, it is crucial to control the degradation rate of L-PBF magnesium scaffolds. Existing magnesium alloy bone repair scaffolds still face the following key issues in practical applications:

[0006] (1) Rapid degradation problem: Magnesium alloys are prone to rapid corrosion in body fluid environments, causing the scaffold to collapse before bone healing is complete. In addition, the release of a large amount of hydrogen and magnesium ions during the corrosion process can cause local tissue reactions and affect the stability of the bone regeneration microenvironment.

[0007] (2) Inadequate structural design: Natural bone tissue has a multi-scale hierarchical structure from nano to macro scales. Traditional scaffold design is mostly focused on a single scale, which makes it difficult to simultaneously meet multiple requirements such as mechanical strength, cell adhesion, nutrient transport, and angiogenesis, limiting the scaffold's comprehensive support for the bone repair process.

[0008] (3) Limitations of coating technology: Although studies have been conducted to improve the degradation performance of magnesium alloy stents through surface coating technology, such as high temperature oxide film (HTO) and hydrotalcite (LDH) coating, the protective effect of a single coating is limited, especially in complex body fluid environments where it is difficult to maintain structural integrity for a long time. At the same time, the poor uniformity and adhesion of the coating also limit its practical application.

[0009] (4) Insufficient personalized adaptation: Traditional processing methods make it difficult to prepare scaffolds that precisely match the shape of bone defects, resulting in poor contact between the scaffold and the bone defect area after implantation, affecting bone integration and repair effects.

[0010] Therefore, there is an urgent need for a bone repair scaffold that can combine multi-scale structural design, regulated degradation performance and personalized manufacturing capabilities, which can not only provide early mechanical support, but also gradually degrade to promote bone regeneration and achieve effective repair of complex bone defects. Summary of the invention

[0011] The purpose of the present invention is to provide a multi-scale graded magnesium alloy bracket and a preparation method and application thereof, so as to solve the above-mentioned problems existing in the prior art.

[0012] In order to achieve the above-mentioned object of the invention, the present invention provides the following technical solutions:

[0013] The present invention provides a method for preparing a multi-scale graded magnesium alloy stent, comprising the following steps:

[0014] (1) Macrostructure design: The shape of the stent is designed using CAD software based on the anatomical characteristics of the patient's bone defect;

[0015] Mesoscopic pore design: The internal pores of the scaffold are designed using a triple-periodic minimum surface structure;

[0016] (2) According to the macroscopic structure design and mesoscopic pore design of step (1), magnesium alloy powder is used as the substrate and 3D printing is performed by laser powder bed fusion technology to obtain a matrix bracket;

[0017] (3) chemically polishing the substrate support to obtain a polished support;

[0018] (4) subjecting the polished stent to high temperature oxidation treatment to form an oxide layer on the surface to obtain an oxidized stent;

[0019] (5) The oxidized stent is subjected to a hydrothermal reaction to form a hydrotalcite layer on the surface to obtain the multi-scale graded magnesium alloy stent.

[0020] Preferably, the triple periodic minimum surface structure in step (1) is a sheet-like diamond structure; the curvature of the triple periodic minimum surface structure is -10 to 2 mm. -1 .

[0021] Preferably, the mesoscopic pores in step (1) have a pore size of 400 to 800 μm and a porosity of 50 to 80%.

[0022] Preferably, the particle size of the magnesium alloy powder in step (2) is 5 to 100 μm; the composition of the magnesium alloy powder, by mass fraction, includes: 0.1 to 10% yttrium, 1 to 5% neodymium, 0.1 to 3% gadolinium, 0.01 to 0.15% zirconium, 0.001 to 0.5% zinc, 0.001 to 0.5% manganese, and the balance is magnesium.

[0023] Preferably, the conditions for the 3D printing in step (2) are: laser power of 20 to 200 W; scanning speed of 200 to 2000 mm / s; layer thickness of 10 to 100 μm; and scanning spacing of 20 to 200 μm.

[0024] Preferably, the temperature of the high temperature oxidation treatment in step (4) is 425-575° C.; and the time of the high temperature oxidation treatment is 1-24 hours.

[0025] Preferably, the temperature of the hydrothermal reaction in step (5) is 60 to 180° C.; the time of the hydrothermal reaction is 2 to 24 hours; and the solution of the hydrothermal reaction is an alkaline aqueous solution of aluminum salt.

[0026] Preferably, the thickness of the oxide layer in step (4) is 1 to 30 μm; and the thickness of the hydrotalcite layer in step (5) is 1 to 50 μm.

[0027] The present invention also provides a multi-scale graded magnesium alloy stent prepared by a method for preparing the multi-scale graded magnesium alloy stent.

[0028] The present invention also provides an application of a multi-scale graded magnesium alloy scaffold in the preparation of bone tissue engineering materials.

[0029] It can be seen from the above technical solution that compared with the prior art, the present invention has the following beneficial effects:

[0030] (1) The multi-scale graded magnesium alloy stent of the present invention effectively solves the problems of rapid degradation, single structure and function, and insufficient personalized adaptation of magnesium alloy stents in the prior art through structural design, manufacturing process and surface coating technology, significantly improves the degradation performance and mechanical properties of the stent during bone repair, and can significantly promote osteogenic differentiation of cells and angiogenesis.

[0031] (2) The present invention utilizes L-PBF technology to achieve precise control of complex macroscopic geometric shapes and internal pore structures, so that the scaffold can fit seamlessly with the bone defect area, effectively improving the bone integration effect after implantation. The present invention forms a double-layer structure of a high-temperature oxide film and a hydrotalcite layer on the surface of the scaffold through high-temperature oxidation treatment and hydrothermal treatment in sequence. The micron-thick coating formed by the oxide layer and the hydrotalcite layer can significantly inhibit the rapid degradation of the scaffold; the hydrotalcite nanosheets in the hydrotalcite layer can significantly improve the surface roughness and hydrophilicity of the scaffold, which is beneficial to the attachment and differentiation of osteoblasts. The multi-scale design of the scaffold of the present invention (macroscopic shape, mesoscopic pores, microscopic coating, nano-surface) effectively mimics the hierarchical structure of natural bone and significantly improves the bioadaptability of the scaffold.

[0032] (3) The scaffold of the present invention is used as a bone repair material, which helps to reduce complications of bone defect treatment and improve the quality of life of patients. Compared with traditional materials, it reduces the need for secondary surgery and reduces medical costs. At the same time, it expands market application prospects through customized manufacturing. BRIEF DESCRIPTION OF THE DRAWINGS

[0033] In order to more clearly illustrate the embodiments of the present invention or the technical solutions in the prior art, the drawings required to be used in the description of the embodiments or the prior art are briefly introduced below.

[0034] Figure 1 Schematic diagram of the synthesis process of the multi-scale graded magnesium alloy stent of Example 1; wherein A is the macroscopic structure design and mesoscopic pore design of the stent, and B is a schematic diagram of the synthesis process;

[0035] Figure 2 The microscopic morphology of WE43 magnesium alloy powder; A is the microscopic morphology at a scale of 150 μm, and B is the microscopic morphology at a scale of 10 μm;

[0036] Figure 3 The appearance structures and microstructure morphology of the printed stent, AP stent, HTO stent, and HTO-LDH stent obtained in Example 1; wherein A is an appearance structure diagram, B is a cross-sectional microstructure morphology diagram, C is a surface microstructure morphology diagram, and D is a surface energy spectrum analysis diagram of the HTO-LDH stent;

[0037] Figure 4 The degradation test results of the AP stent, HTO stent, and HTO-LDH stent obtained in Example 1; wherein, (a) is the potential polarization curve, (b) is the weight loss rate of the stent before and after immersion in the r-SBF solution, (c) is the amount of hydrogen released before and after immersion in the r-SBF solution, (d) is the surface morphology of the stent after immersion in the r-SBF solution, (e) is the elastic modulus of the stent before and after immersion in the r-SBF solution, and (f) is the yield strength of the stent before and after immersion in the r-SBF solution;

[0038] Figure 5 The in vitro biocompatibility and osteogenic performance test results of the AP scaffold, HTO scaffold, and HTO-LDH scaffold obtained in Example 1 are shown in Figure 1; wherein (a) is a schematic diagram of the experimental process, (b) is the live-dead staining result, (c) is the CCK-8 test result, (d) is the alkaline phosphatase test result of BMSC cells, and (e) is the Alizarin Red staining result;

[0039] Figure 6 The atomic force surface topography and contact angle results of the AP scaffold, HTO scaffold, and HTO-LDH scaffold obtained in Example 1; wherein (a) is the atomic force surface topography, and (b) is the contact angle result;

[0040] Figure 7 The in vitro cell adhesion test results of the AP scaffold, HTO scaffold, and HTO-LDH scaffold obtained in Example 1; wherein (a) is a schematic diagram of the experimental process, (b) is the CLSM observation result, (c) is the result of the number of adhered cells, (d) is the result of the area of ​​adhered cells, and (e) is the microscopic morphology of the adhered cells;

[0041] Figure 8The experimental results of bone regeneration behavior of the AP scaffold, HTO scaffold and HTO-LDH scaffold obtained in Example 1 in animals; wherein, (a) is a schematic diagram of the experimental process, (b) is a Micro-CT reconstruction of the femoral defect, (c) is the results of H&E and Masson histological staining, (d) is an analysis of the volume of space occupied by the scaffold and the new bone mass, (e) is the result of sequential fluorescent labeling, (f) is the Micro-CT new bone mineral density result, and (g) is the Micro-CT new bone mass result. DETAILED DESCRIPTION

[0042] The present invention provides a method for preparing a multi-scale graded magnesium alloy stent, comprising the following steps:

[0043] (1) Macrostructure design: The shape of the stent is designed using CAD software based on the anatomical characteristics of the patient's bone defect;

[0044] Mesoscopic pore design: The internal pores of the scaffold are designed using a triple-periodic minimum surface structure;

[0045] (2) According to the macroscopic structure design and mesoscopic pore design of step (1), magnesium alloy powder is used as the substrate and 3D printing is performed by laser powder bed fusion technology to obtain a matrix bracket;

[0046] (3) chemically polishing the substrate support to obtain a polished support;

[0047] (4) subjecting the polished stent to high temperature oxidation treatment to form an oxide layer on the surface to obtain an oxidized stent;

[0048] (5) subjecting the oxidized stent to a hydrothermal reaction to form a hydrotalcite layer on the surface, thereby obtaining the multi-scale graded magnesium alloy stent.

[0049] In the present invention, the triply periodic minimum surface (TPMS) structure in step (1) is preferably a sheet-diamond structure. This structural design has a uniform stress distribution and can improve the mechanical stability and the bearing capacity of the bracket.

[0050] In the present invention, the curvature of the triple periodic minimum surface structure in step (1) is preferably -10 to 2 mm. -1 , more preferably -7 to 1 mm -1 , more preferably -5 to -2 mm -1 The mesoscopic pore design of the present invention presents a relatively uniform average curvature distribution as a whole, avoiding drastic changes in curvature, which is beneficial to cell adhesion and differentiation.

[0051] In the present invention, the pore size of the mesoscopic pores in step (1) is preferably 400-800 μm, and the porosity is preferably 50-80%, more preferably 60-75%, and more preferably 70%. The pore size distribution of the present invention is beneficial to nutrient transport, cell migration and angiogenesis.

[0052] In the present invention, the particle size of the magnesium alloy powder in step (2) is preferably 5 to 100 μm, more preferably 10 to 70 μm, and even more preferably 20 to 63 μm.

[0053] In the present invention, the composition of the magnesium alloy powder in step (2) includes, by mass fraction, 0.1-10% yttrium, 1-5% neodymium, 0.1-3% gadolinium, 0.01-0.15% zirconium, 0.001-0.5% zinc, 0.001-0.5% manganese, and the balance is magnesium. The yttrium element in the magnesium alloy powder can form a dense oxide film on the surface of the magnesium alloy under high-temperature oxidation treatment, which can effectively slow down the degradation of the magnesium alloy. The density and elastic modulus of the magnesium alloy are close to those of human bones, and it has good mechanical properties, biocompatibility and degradability. The magnesium ions released when the scaffold degrades in the body can promote bone formation and angiogenesis.

[0054] In the present invention, the conditions of 3D printing in step (2) are as follows: the laser power is preferably 20-200W, more preferably 50-100W, and more preferably 60W; the scanning speed is preferably 200-2000mm / s, more preferably 500-1000mm / s, and more preferably 600mm / s; the layer thickness is preferably 10-100μm, more preferably 20-40μm, and more preferably 20μm; the scanning spacing is preferably 20-200μm, more preferably 60-90μm, and more preferably 70μm. By adjusting the laser energy, scanning speed, layer thickness, and scanning spacing, high-density and high-precision forming of magnesium alloys can be achieved, and a stent with a complex macroscopic geometry and an internal interconnected pore structure can be prepared.

[0055] In the present invention, the step (3) further includes ultrasonic cleaning of the substrate support before the chemical polishing. The present invention does not limit the conditions of ultrasonic cleaning, and the conventional selection by those skilled in the art can be adopted.

[0056] In the present invention, the time of the chemical polishing in step (3) is preferably 1 to 20 minutes, more preferably 1 to 10 minutes, and more preferably 1 minute; the temperature of the chemical polishing is preferably room temperature.

[0057] In the present invention, the polishing liquid for chemical polishing in step (3) is preferably an alcohol solution of 5% nitric acid + 5% hydrochloric acid.

[0058] In the present invention, the temperature of the high temperature oxidation treatment in step (4) is preferably 425-575°C, more preferably 450-550°C, and more preferably 525°C; the time of the high temperature oxidation treatment is preferably 1-24h, more preferably 6-12h, and more preferably 8h.

[0059] In the present invention, the thickness of the oxide layer in step (4) is preferably 1 to 30 μm, more preferably 1.2 to 10 μm, and more preferably 1.5 μm. The oxide layer formed by the present invention is continuous and uniform, and is mainly composed of Y 2 O 3 and Nd 2 O 3 The composition can significantly inhibit the rapid degradation of the stent and enhance the adhesion of the coating.

[0060] In the present invention, the temperature of the hydrothermal reaction in step (5) is preferably 60-180°C, more preferably 80-150°C, and more preferably 120°C; the time of the hydrothermal reaction is preferably 2-24h, more preferably 4-10h, and more preferably 6h; the solution of the hydrothermal reaction is preferably an alkaline aqueous solution of aluminum salt.

[0061] In the present invention, the aluminum salt in the alkaline aqueous solution of aluminum salt is preferably aluminum nitrate or aluminum chloride, and more preferably aluminum nitrate; the concentration of the aluminum salt is preferably 0.002-0.2 mol / L, more preferably 0.01-0.1 mol / L, and more preferably 0.02 mol / L; the pH of the alkaline aqueous solution of aluminum salt is preferably 9-12, more preferably 9.5-11, and more preferably 10; the pH is preferably adjusted by 0.1-10 mol / L of sodium hydroxide solution, more preferably 0.5-3 mol / L, and more preferably 1 mol / L.

[0062] In the present invention, the thickness of the hydrotalcite layer in step (5) is preferably 1 to 50 μm, more preferably 2 to 20 μm, and more preferably 2 μm. The hydrotalcite layer formed by the present invention is chemically stable and has excellent biocompatibility, further enhancing the anti-corrosion performance, and capturing corrosive ions (such as Cl - ).

[0063] The present invention also provides a multi-scale graded magnesium alloy stent prepared by a method for preparing the multi-scale graded magnesium alloy stent.

[0064] The present invention also provides an application of a multi-scale graded magnesium alloy scaffold in the preparation of bone tissue engineering materials.

[0065] The technical solutions in the embodiments of the present invention are described clearly and completely below. Obviously, the described embodiments are only part of the embodiments of the present invention, not all of them. Based on the embodiments of the present invention, all other embodiments obtained by ordinary technicians in this field without creative work are within the scope of protection of the present invention.

[0066] Example 1

[0067] This embodiment provides a method for preparing a multi-scale graded magnesium alloy stent, and the schematic diagram of the synthesis process is shown in FIG. Figure 1 As shown, the following steps are included:

[0068] (1) Macrostructure design: The shape of the stent is designed using CAD software based on the anatomical characteristics of the patient's bone defect;

[0069] Mesoscopic pore design: The internal pores of the stent are designed using a triple-periodic minimum surface structure, specifically a Sheet-Diamond type, with a pore size controlled at 400-800 μm and a porosity of 70%;

[0070] (2) According to the macroscopic structure design and mesoscopic pore design of step (1), WE43 magnesium alloy powder is used as the substrate, and its composition is 3.50wt% yttrium, 2.28wt% neodymium, 1.18wt% gadolinium, 0.07wt% zirconium, 0.20wt% zinc, 0.02wt% manganese and the balance magnesium. Laser powder bed melting technology is used for 3D printing. The printer model is Platinum S210. The printing parameters are as follows: laser spot diameter 70μm, laser power 60W, scanning speed 600mm / s, layer thickness 20μm, scanning spacing 70μm, printing platform preheating temperature 200℃, scanning strategy is Z-shaped and the interlayer rotation is 67°, and a matrix bracket is obtained, which is recorded as the printed bracket;

[0071] (3) The base stent was cleaned by ultrasonic cleaning to remove residual powder on the surface, and then chemically polished in an alcohol solution of 5% nitric acid + 5% hydrochloric acid at room temperature for 1 min to obtain a polished stent, which was recorded as an AP stent;

[0072] (4) The polished stent was placed in a muffle furnace in an air atmosphere and subjected to high-temperature oxidation treatment at 525° C. for 8 h to form an oxide layer on the surface, thereby obtaining an oxidized stent, which was recorded as an HTO stent;

[0073] (5) The oxidized stent was placed in a 0.02 mol / L aluminum nitrate solution with a pH of 10, and subjected to a hydrothermal reaction at 120° C. for 6 h to form a hydrotalcite layer on the surface, thereby obtaining a double-layer coated stent, which is the multi-scale graded magnesium alloy stent, denoted as the HTO-LDH stent.

[0074] The microstructure of the WE43 magnesium alloy powder used in this example is as follows: Figure 2 As shown. Figure 2 It can be seen that the magnesium alloy powder has high sphericity and less satellite powder is observed, so the powder has good fluidity, and the powder particle size is 20 to 63 μm.

[0075] The appearance and structure of the printed stent, AP stent, HTO stent, and HTO-LDH stent obtained in this embodiment are as follows: Figure 3 As shown in A, the cross-sectional microstructures of the AP scaffold, HTO scaffold, and HTO-LDH scaffold are shown in Figure 3 As shown in B, the surface microstructures of AP scaffolds, HTO scaffolds, and HTO-LDH scaffolds are shown in Figure 3 As shown in C, the surface energy spectrum analysis of HTO-LDH scaffold is as follows Figure 3 As shown in D. Figure 3 As shown in Figure A, the AP scaffold can be obtained by chemical polishing, the HTO scaffold can be obtained by high temperature oxidation treatment, and the HTO-LDH scaffold can be obtained by further hydrothermal treatment. Figure 3 From Figure B, we can see that the thickness of the oxide layer is about 1.5 μm, the thickness of the hydrotalcite layer is about 2 μm, and the surface of the HTO-LDH scaffold is evenly covered by a double-layer micron-scale coating formed by the oxide layer and the hydrotalcite layer. Figure 3 As shown in Figure 5C, a large amount of secondary phases were observed on the surface of the AP scaffold, while after HTO treatment, the surface was 2 O 3 and Nd 2 O 3 The surface of the HTO-LDH scaffold is a composite coating composed of an oxide film and a hydrotalcite coating, with a thickness of about 3.5 μm. At the same time, hydrotalcite nanosheets with a lateral size of 600 to 800 nm were observed on the surface of the scaffold. Figure 3 As shown in Figure 5D, according to the element distribution on the surface of the HTO-LDH scaffold, in addition to Mg, Y, O, and Nd elements, a uniform distribution of Al element was also observed, which suggests that Mg-Al hydrotalcite was uniformly formed on the surface.

[0076] The performance tests of the AP stent, HTO stent, and HTO-LDH stent obtained in this embodiment are as follows:

[0077] (1) Degradation performance test:

[0078] The scaffolds were immersed in simulated body fluid (r-SBF) and tested at 37°C for 28 days. Figure 4 (a) shows the potential polarization curve of the stent measured in r-SBF solution. The corrosion current density of AP, HTO and HTO-LDH stents is 2.51×10 -4 ,9.55×10 -6 and 1.57×10 -6 A.cm-2 , indicating that the high-temperature oxidation treatment effectively improved the early transient corrosion resistance, while the hydrotalcite layer further enhanced the early transient corrosion resistance. Figure 4 (b) and (c) show the results of long-term immersion experiments of the stents in r-SBF solution at 37°C. The weight loss rate of the AP stent after 1 day of immersion was as high as 31.5±6.1%, and it was completely degraded within 14 days. The weight loss rates of the HTO stent after 3 days and 28 days were 16.5±1.5% and 75.3±10.0%, respectively. The weight loss rate of the HTO-LDH stent after 28 days of immersion was only 23.6±3.6%, and the hydrogen release amount was 4.4±0.9mL·cm -2 , about 0.3mm·year -1 . Figure 4 (d) shows the surface morphology of the scaffold after immersion. The AP scaffold showed obvious degradation and surface cracking after 1 day of immersion, and its structural integrity was severely damaged. After 3 days, large pitting pits appeared on the surface. The HTO scaffold did not show obvious degradation after 3 days of immersion, and maintained its structural integrity before 14 days, but obvious cracks appeared on the surface. The surface of the HTO-LDH scaffold remained intact after 14 days, with few cracks. The nanosheets on the surface were almost intact after 3 days, but were covered by degradation products after 14 days. After 28 days, the HTO-LDH scaffold still maintained its structural integrity; in contrast, the HTO scaffold collapsed, with a weight loss rate of 75.3±10.0%.

[0079] (2) Mechanical properties test:

[0080] Compression testing is performed using a universal material testing machine. Figure 4 (e) and (f) show the elastic modulus and yield strength results of the stent before and after immersion in r-SBF solution. The yield strength of the HTO-LDH stent is 32.4MPa, which can effectively withstand the normal load of human bones. The damage caused by degradation to the stent structure greatly weakened the bearing capacity of the stent. After immersion for 3 days, the elastic modulus and yield strength of the AP stent decreased significantly, with a decrease of 95% and 98%, respectively. The HTO stent decreased by 73% and 66% after 14 days, while the HTO-LDH stent decreased by 46% and 43% after 28 days. In summary, the hydrotalcite layer composite oxide film can effectively regulate the degradation of magnesium alloy stents and help provide longer reliable mechanical support during implantation.

[0081] (3) Biological performance test:

[0082] Human bone marrow mesenchymal stem cells (BMSCs) were cultured in Transwell chambers to evaluate the biocompatibility and osteogenic properties of the scaffolds. The schematic diagram of the experimental process is shown in Figure 5 As shown in (a). Figure 5(b) shows the live-dead staining results. It can be observed that there are almost no dead cells in all groups, but the number of live cells in the AP scaffold group is significantly lower. It is speculated that the rapid degradation of the AP scaffold leads to excessive magnesium ion concentration, which is not conducive to osteoblast proliferation, but does not lead to massive cell death. Figure 5 (d) and (e) show the results of alkaline phosphatase and alizarin red staining of BMSC cells. The HTO-LDH scaffold group showed a significant enhancement of osteogenic differentiation, with ALP activity 2.08 times that of the HTO scaffold group and 6.4 times that of the AP scaffold group. Alizarin red staining was used to evaluate the osteogenic differentiation performance after 21 days of culture, and a large number of mineralized nodules were observed in the HTO-LDH scaffold group.

[0083] The surface morphology of the scaffold was observed by atomic force microscopy. Figure 6 As shown in (a), the surface roughness of the HTO-LDH scaffold was the highest, with Ra reaching 100.4±7.5 nm, while there was no obvious difference between the AP and HTO scaffolds. Figure 6 (b) is the contact angle of the scaffold. Compared with the AP scaffold, the contact angle of the HTO scaffold surface decreases due to the formation of an oxide layer, while the contact angle of the HTO-LDH scaffold is the smallest, which is 21.3±6.2°, indicating that the hydrotalcite nanosheets greatly improve the hydrophilicity of the scaffold surface.

[0084] In vitro cell adhesion experiments on scaffolds were performed using BMSC cells. Figure 7 The morphology and quantitative results of BMSC cells after 12 hours of direct cell adhesion. Figure 7 As shown in (b) and (c), the HTO-LDH scaffold exhibited the best cell adhesion effect, while almost no cell adhesion was observed on the surface of the AP scaffold. The number of cells adhering to the surface of the HTO-LDH scaffold was 3.26 times that of the HTO scaffold and 13.04 times that of the AP scaffold. The BMSC cells cultured on the HTO-LDH scaffold were stretched, and the extended pseudopodia firmly adhered to the scaffold surface. Figure 7 As shown in (d), the cell area covered on the surface of HTO-LDH scaffold is 4.25 times and 19.89 times that of HTO and AP scaffolds, respectively. Figure 7 As shown in (e), scanning electron microscopy showed that the morphology of cells adhering to the HTO-LDH scaffold was flat polygonal, and the cells extended a large number of pseudopodia anchored on the hydrotalcite nanosheets.

[0085] (4) Animal experiment verification:

[0086] The magnesium alloy stent was implanted into the cylindrical defect of the rabbit femoral condyle for a 12-week animal experiment. The animal experiment diagram is shown in Figure 8During the study, all rabbits maintained good health, with no obvious adverse reactions to their vital organs, including the heart, liver, spleen, lungs, and pancreas. Figure 8 Middle (b) shows Micro-CT images of the scaffolds after different implantation times. The AP scaffold group lost its structural integrity after 6 weeks, and a small amount of new bone was observed. The HTO scaffold group was slightly degraded after 6 weeks of implantation, but the overall structure was basically unaffected. A small amount of new bone tissue was found in the porous structure and close to the pore wall. The HTO-LDH scaffold was structurally intact after 6 weeks of implantation, and more new bone formation was observed. After 12 weeks of implantation, the AP scaffold was almost completely degraded, and new bone formation increased slightly compared to 6 weeks. The HTO scaffold lost its structural integrity, and a certain amount of new bone was observed, while the HTO-LDH scaffold still maintained good structural integrity. At the same time, the high porosity and high connectivity provided sufficient space for bone growth, and a large amount of new bone tissue was observed inside the scaffold. Figure 8 (d), (f), and (g) are quantitative results obtained based on Micro-CT. The amount of new bone formation in the HTO-LDH scaffold group was 3.57 times that of the AP scaffold group, and the bone mineral density was 2.3 times that of the AP scaffold group. Twelve weeks after implantation, the HTO-LDH scaffold only degraded by 11.13%, maintaining good structural integrity, while the HTO scaffold degraded by 63% and lost its complete structure. Figure 8 Middle (c) shows the results of histological staining of the sections, including H&E and Masson staining. New bone and fibrous tissue growth were observed in the pores of all scaffolds. Figure 8 Middle (e) is the result of sequential fluorescence labeling, indicating that the mineral deposition rate of the HTO-LDH scaffold group is the highest. In summary, the HTO-LDH scaffold with a multi-scale hierarchical structure promotes bone reconstruction and achieves high-quality bone repair.

[0087] The above is only a preferred embodiment of the present invention. It should be pointed out that for ordinary technicians in this technical field, several improvements and modifications can be made without departing from the principle of the present invention. These improvements and modifications should also be regarded as the scope of protection of the present invention.

Claims

1. A method for preparing a multi-scale graded magnesium alloy stent, characterized in that: The following steps are involved: (1) Macrostructure design: The shape of the stent is designed using CAD software based on the anatomical characteristics of the patient's bone defect; Mesoscopic pore design: The internal pores of the scaffold are designed using a triple-periodic minimum surface structure; (2) According to the macroscopic structure design and mesoscopic pore design of step (1), magnesium alloy powder is used as the substrate and 3D printing is performed by laser powder bed fusion technology to obtain a matrix bracket; (3) chemically polishing the substrate support to obtain a polished support; (4) subjecting the polished stent to high temperature oxidation treatment to form an oxide layer on the surface to obtain an oxidized stent; (5) The oxidized stent is subjected to a hydrothermal reaction to form a hydrotalcite layer on the surface to obtain the multi-scale graded magnesium alloy stent.

2. The method for preparing a multi-scale graded magnesium alloy stent according to claim 1, characterized in that: The triple periodic minimum curved surface structure in step (1) is a sheet-like diamond structure; the curvature of the triple periodic minimum curved surface structure is -10 to 2 mm. -1 .

3. The method for preparing a multi-scale graded magnesium alloy stent according to claim 1, characterized in that: The mesoscopic pores in step (1) are designed to have a pore size of 400 to 800 μm and a porosity of 50 to 80%.

4. The method for preparing a multi-scale graded magnesium alloy stent according to claim 1 or 3, characterized in that: The particle size of the magnesium alloy powder in step (2) is 5 to 100 μm; the composition of the magnesium alloy powder, by mass fraction, includes: 0.1 to 10% yttrium, 1 to 5% neodymium, 0.1 to 3% gadolinium, 0.01 to 0.15% zirconium, 0.001 to 0.5% zinc, 0.001 to 0.5% manganese, and the balance is magnesium.

5. The method for preparing a multi-scale graded magnesium alloy stent according to claim 4, characterized in that: The conditions for the 3D printing in step (2) are: laser power of 20 to 200 W; scanning speed of 200 to 2000 mm / s; layer thickness of 10 to 100 μm; and scanning spacing of 20 to 200 μm.

6. The method for preparing a multi-scale graded magnesium alloy stent according to claim 1 or 5, characterized in that: The temperature of the high temperature oxidation treatment in step (4) is 425-575° C.; the time of the high temperature oxidation treatment is 1-24 hours.

7. The method for preparing a multi-scale graded magnesium alloy stent according to claim 6, characterized in that: The temperature of the hydrothermal reaction in step (5) is 60 to 180° C.; the time of the hydrothermal reaction is 2 to 24 hours; and the solution of the hydrothermal reaction is an alkaline aqueous solution of aluminum salt.

8. The method for preparing a multi-scale graded magnesium alloy stent according to claim 7, characterized in that: The thickness of the oxide layer in step (4) is 1 to 30 μm; the thickness of the hydrotalcite layer in step (5) is 1 to 50 μm.

9. A multi-scale graded magnesium alloy stent prepared by the method for preparing a multi-scale graded magnesium alloy stent according to any one of claims 1 to 8.

10. Use of the multi-scale graded magnesium alloy scaffold according to claim 9 in the preparation of bone tissue engineering materials.

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