Biphase continuous medical biodegradable magnesium alloy and TCP composite stent material and preparation method thereof

Abstract

The invention discloses a biphase continuous medical biodegradable magnesium alloy and TCP composite scaffold material and a preparation method thereof.The preparation method comprises the following steps that TCP powder and resin are mixed to obtain a TCP-resin mixture, a photoinitiator and a dispersing agent are added into the TCP-resin mixture, and ball milling is conducted to obtain slurry; performing 3D printing on the slurry through a resin light curing method to obtain a TCP-resin bracket; sintering the TCP-resin scaffold, and performing furnace cooling to room temperature to obtain a TCP porous scaffold; the preparation method comprises the following steps: calcining a TCP porous scaffold in a mold, adding a degradable magnesium alloy melt into the mold, and enabling the magnesium alloy melt to fully flow into pores of the TCP porous scaffold by adopting a pressure permeation method, so as to obtain the biphase continuous medical biodegradable magnesium alloy and TCP composite scaffold material. The obtained biphase continuous medical biodegradable magnesium alloy and TCP composite scaffold material magnesium alloy is degraded preferentially, enough space is provided for growth of new bones, along with degradation of the magnesium alloy, the TCP porous scaffold can still keep certain mechanical support, and it is ensured that bone tissue obtains continuous structural support in the repairing process.

Description

Technical Field

[0001] The invention belongs to the technical field of biomedical metal material preparation, and in particular relates to a two-phase continuous medical biodegradable magnesium alloy and TCP composite stent material and a preparation method thereof. Background Art

[0002] At present, the widely used degradable implants mainly include degradable polymer materials, degradable ceramic materials and degradable metal materials, but these materials all have certain defects.

[0003] The degradation rate of degradable polymer materials is difficult to control, their biocompatibility is poor, and their degradation products may trigger severe inflammatory responses. In addition, the mechanical properties of degradable polymer materials are difficult to meet the load requirements of degradable implants.

[0004] Among degradable metal materials, degradable zinc alloys and degradable magnesium alloys have attracted attention due to their mechanical properties similar to natural bones, biodegradability, and excellent biocompatibility. However, degradable zinc alloys suddenly release a large amount of zinc ions in the initial stage of degradation, and excessive zinc ion concentrations are toxic to cells, limiting their application in clinical medicine. Compared with degradable zinc alloys, degradable magnesium alloys degrade too quickly, resulting in insufficient mechanical support in the early stage of implantation.

[0005] Although the mechanical properties of degradable ceramic materials are improved compared to degradable polymer materials, they still cannot meet the requirements of mechanical strength and hardness, resulting in insufficient bearing capacity, such as hydroxyapatite (HA) and tricalcium phosphate (TCP); in terms of degradation, the composition of HA and TCP is similar to that of natural bone, with excellent biocompatibility, bioactivity and osteoinductivity, but their degradation rate is slow and difficult to control, which does not conform to the bone growth cycle and cannot allow new bone to grow. Summary of the invention

[0006] In order to solve the problems of insufficient mechanical properties and slow degradation rate of TCP, the purpose of the present invention is to provide a dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material, which introduces a degradable magnesium alloy into the TCP porous scaffold. The addition of the degradable magnesium alloy significantly improves the overall mechanical properties of the material, and preferentially degrades to provide sufficient space for the growth of new bones. As the magnesium alloy degrades, the TCP porous scaffold can still maintain a certain mechanical support to ensure that the bone tissue obtains continuous structural support during the repair process. That is, the dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material of the present invention can not only adjust its degradation rate, but also is conducive to bone repair at the same time.

[0007] Another object of the present invention is to provide a method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material. The preparation method utilizes resin photocuring (LCD) technology in 3D printing to accurately control and obtain a TCP porous stent with a high-precision structure, and utilizes pressure infiltration technology to tightly combine the degradable magnesium alloy and the TCP porous stent to obtain a dual-phase continuous medical biodegradable magnesium alloy and TCP stent composite material.

[0008] The purpose of the present invention is achieved through the following technical solutions.

[0009] A two-phase continuous medical biodegradable magnesium alloy and TCP composite stent material comprises a TCP porous stent and a degradable magnesium alloy composited in the TCP porous stent.

[0010] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material, comprising the following steps:

[0011] Step 1, mixing TCP powder and resin to obtain a TCP-resin mixture, adding a photoinitiator and a dispersant to the TCP-resin mixture, and ball milling to obtain a uniformly dispersed slurry;

[0012] In the step 1, the ratio of the mass fraction of the TCP powder, the volume fraction of the resin, the mass fraction of the photoinitiator and the volume fraction of the dispersant is (52-53): (48.5-51.5): (0.7822-0.7832): (0.9-1.3), the unit of the mass fraction is g, and the unit of the volume fraction is mL.

[0013] In step 1, the resin is a mixture of 1,6-hexanediol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate, and the ratio of 1,6-hexanediol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate is (29.5-30.5): (14.5-15.5): (4.5-5.5) by volume.

[0014] In the step 1, the photoinitiator is (2,4,6-trimethylbenzoyl)phosphine oxide, and the dispersant is a silane coupling agent.

[0015] In the step 1, the rotation speed of the ball mill is 160 to 200 r / min, and the ball milling time is 19 to 21 hours.

[0016] Step 2, 3D printing the slurry by LCD photocuring method to obtain a TCP-resin scaffold, wherein the pore size of the TCP-resin scaffold is 600-1200 μm and the porosity is 75-85%;

[0017] In step 2, the TCP-resin stent is formed by connecting a plurality of unit cells, the shape of each unit cell is defined by a Gyroid function, the unit cell is 2 mm×2 mm×2 mm, and the wall thickness is -0.01 mm.

[0018] In the step 2, the TCP-resin bracket model is designed by the Gyroid function in the nTopology software, a three-periodic minimal surface (TPMS) structure is constructed, and then the TCP-resin bracket model is imported into the 3D printer in STL format for the 3D printing.

[0019] In the above technical solution, the bias parameter of the Gyroid function model in the nTopology software is 0.9.

[0020] Step 3, sintering the TCP-resin scaffold at 310-340° C. for 55-65 min, then sintering at 630-680° C. for 115-125 min, and finally sintering at 1100-1200° C. for 115-125 min, and cooling the furnace to room temperature to obtain a TCP porous scaffold;

[0021] In the step 3, the heating rate of the sintering is 0.2-1°C / min.

[0022] Step 4: Place the TCP porous scaffold in a mold and calcine it at 590-610° C. for 40-50 min. Then, add a 720-740° C. biodegradable magnesium alloy melt into the mold at 590-610° C., and use a pressure infiltration method to allow the magnesium alloy melt to fully flow into the pores of the TCP porous scaffold to obtain a dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material.

[0023] In step 4, the degradable magnesium alloy melt is a magnesium-zinc alloy, a magnesium-zinc-calcium alloy or a magnesium-zinc-calcium-manganese alloy.

[0024] In step 4, the method for preparing the degradable magnesium alloy melt comprises: mixing pure magnesium with a metal, smelting at 720-740° C. for at least 30 minutes, the metal being one or more of zinc, calcium and manganese;

[0025] When the degradable magnesium alloy melt is a magnesium-zinc alloy, the metal is zinc, and the content of zinc in the degradable magnesium alloy melt is 1.3-1.7 wt %;

[0026] When the degradable magnesium alloy melt is a magnesium-zinc-calcium alloy, the metals are zinc and calcium, the zinc content in the degradable magnesium alloy melt is 1.3-1.7wt%, and the calcium content in the degradable magnesium alloy melt is 0.09-0.11wt%;

[0027] When the degradable magnesium alloy melt is a magnesium-zinc-calcium-manganese alloy, the metals are zinc, calcium and manganese, the zinc content in the degradable magnesium alloy melt is 1.3-1.7wt%, the calcium content in the degradable magnesium alloy melt is 0.09-0.11wt%, and the manganese content in the degradable magnesium alloy melt is 0.18-0.22wt%.

[0028] In step 4, the pressure infiltration method includes: introducing argon gas into the mold until the pressure inside the mold is 0.1-0.3 MPa, so that the degradable magnesium alloy melt can fully flow into the pores of the TCP porous scaffold.

[0029] Compared with the prior art, the present invention has the following beneficial effects:

[0030] (1) The dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained by the present invention not only has a good bonding surface, but also maintains the integrity and connectivity of the TCP porous stent and the magnesium alloy.

[0031] (2) The degradation rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained by the present invention is slow, thereby reducing the degradation rate of the magnesium alloy.

[0032] (3) In the initial stage of degradation of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material obtained by the present invention, the magnesium alloy is preferentially degraded, the TCP porous scaffold is retained and a certain mechanical support is provided, so that new bone can continue to grow in the pores of the TCP porous scaffold.

[0033] (4) The preparation method of the present invention can utilize resin photocuring (LCD) technology to precisely adjust the pore structure, pore size, porosity and morphology of the TCP porous scaffold by changing parameters according to the specific clinical needs and anatomical characteristics of the patient, construct a TCP porous scaffold, prepare a dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material, improve the degradation rate of the magnesium alloy, and thus achieve personalized treatment.

[0034] (5) The preparation method of the present invention has a wide range of materials for selection. BRIEF DESCRIPTION OF THE DRAWINGS

[0035] Figure 1 The flowchart of the preparation method of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Examples 1 to 3;

[0036] Figure 2 This is a physical picture of the TCP porous scaffold obtained in Example 3;

[0037] Figure 3 The actual picture of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Examples 1 to 3;

[0038] Figure 4 The weight loss rate and degradation rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 and the cast alloy obtained in Comparative Example 1;

[0039] Figure 5 The scanning electron microscope (SEM) images of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 after immersion for 1, 7 and 14 days;

[0040] Figure 6 This is a real picture of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 after the 14th day of in vitro immersion experiment. DETAILED DESCRIPTION

[0041] The technical solution of the present invention is further described in detail below in conjunction with the accompanying drawings and embodiments.

[0042] Examples 1 to 3

[0043] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material, comprising the following steps:

[0044] Step 1: Place TCP powder and resin in a ball mill and mix them to form a TCP-resin system to obtain a TCP-resin mixture; add a photoinitiator and a dispersant to the TCP-resin mixture; add agate ball media to the ball mill; and mill for 20 hours in a planetary ball mill at a speed of 180 r / min to obtain a uniformly dispersed slurry, wherein the photoinitiator is (2,4,6-trimethylbenzoyl)phosphine oxide, the dispersant is a silane coupling agent HK-560, and the resin is 1,6-hexanediol. A mixture of alcohol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate, wherein, by volume, the ratio of 1,6-hexanediol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate is 30:15:5, and the ratio of the mass fraction of TCP powder, the volume fraction of resin, the mass fraction of photoinitiator and the volume fraction of dispersant is 52.5:50:0.7827:1.1, the unit of mass fraction is g, and the unit of volume fraction is mL.

[0045] Step 2, 3D printing the slurry by LCD photocuring method to obtain a TCP-resin bracket, the pore size of the TCP-resin bracket is W, and the porosity is 80%; the TCP-resin bracket is connected by a plurality of unit lattices, the shape of each unit lattice is defined by a Gyroid function, the TCP-resin bracket model is designed by nTopology software, a three-period minimal surface (TPMS) structure is constructed, and the TCP-resin bracket model is imported into a 3D printer (model UNIZ LCDSH 2PLUS) in STL format for 3D printing to obtain the TCP-resin bracket, wherein the bias parameter λ of the Gyroid function in the nTopology software is 0.9, the unit lattice size is set to 2mm×2mm×2mm, and the wall thickness is -0.01mm in the nTopology software, and the pore size and porosity of the TCP-resin bracket model are the same as those of the TCP-resin bracket, the pore size is W, and the porosity is 80%.

[0046] Step 3, place the TCP-resin bracket in a muffle furnace, heat it to 330°C at a heating rate of 1°C / min and sinter it at 330°C for 60min, then heat it to 650°C at a heating rate of 0.2°C / min and sinter it at 650°C for 120min, continue to heat it to 1150°C at a heating rate of 0.5°C / min and sinter it at 1150°C for 120min to remove the resin in the TCP-resin bracket, cool the furnace to room temperature, and obtain a TCP porous bracket.

[0047] Step 4: fix the TCP porous stent with gypsum and place it in a mold, calcine it at 600°C for 40 minutes, add a 730°C degradable magnesium alloy melt (a method for preparing a degradable magnesium alloy melt, comprising: using pure magnesium as a matrix, mixing it with a metal, and smelting it at 730 for 30 minutes) into the mold at 600°C, and introduce argon gas into the mold by pressure infiltration until the pressure in the mold is 0.1MPa, so that the degradable magnesium alloy melt can fully flow into the pores of the TCP porous stent, demold and remove the gypsum to obtain a two-phase continuous medical biodegradable magnesium alloy and TCP composite stent material.

[0048] Reference Figure 1 The preparation method flow chart, the numbers of the two-phase continuous medical biodegradable magnesium alloy and TCP composite stent materials obtained in Examples 1 to 3, W, the degradable magnesium alloy melt, the metals in the degradable magnesium alloy melt and their contents are shown in Table 1.

[0049] Table 1

[0050]

[0051]

[0052] Comparative Example 1

[0053] A method for preparing a cast alloy (number: cast Mg-1.5Zn-0.1Ca-0.2Mn), comprising the following steps: placing pure magnesium, a magnesium-manganese alloy and a magnesium-calcium alloy in a crucible and smelting at 780°C for 30 minutes, cooling to 730°C after smelting to obtain a magnesium-zinc-calcium-manganese alloy melt, adding the 730°C magnesium-zinc-calcium-manganese alloy melt into a mold preheated to 270°C, cooling and demolding to obtain a cast alloy, wherein the zinc content in the magnesium-zinc-calcium-manganese alloy melt is 1.5wt%, the calcium content in the magnesium-zinc-calcium-manganese alloy melt is 0.1wt%, and the manganese content in the magnesium-zinc-calcium-manganese alloy melt is 0.2wt%.

[0054] Comparative Example 2

[0055] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material is basically the same as Example 3, the only difference being that the pore size of the TCP-resin stent is different, and the pore size of the TCP-resin stent in Comparative Example 2 is 600 μm.

[0056] Comparative Example 3

[0057] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material is basically the same as Example 3, the only difference being that the pore size of the TCP-resin stent is different, and the pore size of the TCP-resin stent in Comparative Example 3 is 1200 μm.

[0058] Comparative Example 4

[0059] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material is basically the same as Example 1, the only difference being that the pore size of the TCP-resin stent is different, and the pore size of the TCP-resin stent in Comparative Example 4 is 800 μm.

[0060] Comparative Example 5

[0061] A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material is basically the same as Example 2, the only difference being that the pore size of the TCP-resin stent is different, and the pore size of the TCP-resin stent in Comparative Example 5 is 800 μm.

[0062] Figure 2 This is a physical picture of the TCP porous stent obtained in Example 3. Figure 3 The actual picture of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Examples 1 to 3 is as follows: Figure 3 It can be seen that the dual-phase continuous medical biodegradable magnesium alloy obtained in Examples 1 to 3 has a good bonding surface with the TCP composite stent material.

[0063] An in vitro immersion experiment was carried out according to ASTM-G31-72 standard: a sample with a mass of m0 was immersed in a simulated body fluid (SBF) at 37°C and a pH value of 7.4 for 14 days, and then immersed in a chromic acid solution (the chromic acid concentration in the chromic acid solution was 200 g / L) for 14 days to remove corrosion products, and weighed, recorded as m1. The weight loss rate and degradation rate were calculated based on m1 and m0. In order to ensure the accuracy of the experiment, the simulated body fluid was replaced every 2 days. The sample was one of the two-phase continuous medical biodegradable magnesium alloy and TCP composite stent materials obtained in Examples 1 to 3 and Comparative Examples 2 to 5 and the cast alloy obtained in Comparative Example 1. The ratio of the volume of the simulated body fluid (SBF) to the surface area of ​​the sample was 20 mL cm -2 .

[0064] In the in vitro immersion experiment, the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Comparative Examples 2 to 3 produced a large amount of gas (the gas in the gas was hydrogen) after being immersed for one day, and after being taken out, it was found that it was broken under the action of simulated body fluid (SBF) and could not maintain structural integrity.

[0065] After soaking for two days, the two-phase continuous medical biodegradable magnesium alloy and TCP composite stent materials obtained in comparative examples 4 to 5 all broke under the action of simulated body fluids. After soaking for two days, the two-phase continuous medical biodegradable magnesium alloy and TCP composite stent materials obtained in examples 1 to 2 and comparative examples 4 to 5 broke and could not maintain a complete structure. The reason may be that the degradation rate is too fast, and the degradation of the degradable magnesium alloy produces a large amount of gas (the gas is hydrogen), resulting in a large stress inside.

[0066] The dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 can still maintain its complete structure after being immersed in simulated body fluid for 14 days, indicating that its structure is not damaged due to internal stress during the degradation process.

[0067] The weight loss rate and degradation rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 and the cast alloy obtained in Comparative Example 1 after immersion in simulated body fluid (SBF) for 14 days are as follows: Figure 4 As shown by Figure 4It can be seen that the weight loss rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 is 16.77%, which is significantly lower than the weight loss rate of the cast alloy obtained in Comparative Example 1 (31.72%), and the degradation rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 is 0.039 mm / y, and the degradation rate of the cast alloy obtained in Comparative Example 1 is 0.052 mm / y, indicating that the degradation rate of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material obtained in Example 3 is significantly lower than that of the degradable magnesium alloy (magnesium-zinc-calcium-manganese alloy), so that it still retains certain mechanical properties in the early stage of implantation.

[0068] The corrosion morphology of the sample (the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3) was characterized on the 1st, 7th and 14th day of the in vitro immersion experiment. Figure 5 As shown, Figure 5 a1-a3 are SEM images of the corrosion morphology of the degradable magnesium alloy (magnesium-zinc-calcium-manganese alloy) in the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3 at a scale of 100 μm. Figure 5 a4 to a6 are SEM images of the corrosion morphology of the degradable magnesium alloy (magnesium-zinc-calcium-manganese alloy) in the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3 at a scale of 10 μm. Figure 5 b1-b3 are SEM images of the corrosion morphology of the TCP porous stent in the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3 at a scale of 100 μm. Figure 5 b4 to b6 are SEM images of the corrosion morphology of the TCP porous stent in the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3 at a scale of 10 μm. Figure 5 It can be seen that compared with the TCP porous support part, the corrosion rate of the degradable magnesium alloy (magnesium-zinc-calcium-manganese alloy) part is faster, so the corrosion products first produce a corrosion product layer on the degradable magnesium alloy (magnesium-zinc-calcium-manganese alloy) part, and then the corrosion products are enriched on the TCP porous support part, and a layer of corrosion product layer is generated to cover the TCP porous support part.

[0069] Figure 6 This is a photo of the dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material prepared in Example 3 after the 14th day of in vitro immersion experiment. Figure 6 It can be seen that the TCP porous scaffold part in the dual-phase continuous medical biodegradable magnesium alloy and TCP composite scaffold material prepared in Example 3 still retains a complete structure, while Figure 6The degradable magnesium alloy (Mg-Zn-Ca-Mn alloy) marked by the middle circle has been partially degraded, achieving the expected goal of preferentially degrading the Mg-Zn-Ca-Mn alloy in the early stage of degradation, retaining the TCP porous scaffold and providing certain mechanical support.

[0070] The present invention is described above by way of example. It should be noted that, without departing from the core of the present invention, any simple deformation, modification or other equivalent replacement that can be made by those skilled in the art without inventive effort falls within the protection scope of the present invention.

Claims

1. A dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material, characterized in that: include: A TCP porous scaffold and a degradable magnesium alloy composited in the TCP porous scaffold, wherein the pore size of the TCP-resin scaffold is 600-1200 μm, the porosity is 75-85%, and the degradable magnesium alloy melt is a magnesium-zinc alloy, a magnesium-zinc-calcium alloy or a magnesium-zinc-calcium-manganese alloy.

2. A method for preparing a dual-phase continuous medical biodegradable magnesium alloy and TCP composite stent material, characterized in that: The following steps are involved: Step 1, mixing TCP powder and resin to obtain a TCP-resin mixture, adding a photoinitiator and a dispersant to the TCP-resin mixture, and ball milling to obtain a uniformly dispersed slurry; Step 2, 3D printing the slurry by a resin photocuring method to obtain a TCP-resin scaffold, wherein the pore size of the TCP-resin scaffold is 600-1200 μm and the porosity is 75-85%; Step 3, sintering the TCP-resin scaffold at 310-340° C. for 55-65 min, then sintering at 630-680° C. for 115-125 min, and finally sintering at 1100-1200° C. for 115-125 min, and cooling the furnace to room temperature to obtain a TCP porous scaffold; Step 4: Place the TCP porous stent in a mold and calcine it at 590-610° C. for 40-50 min, then add a 720-740° C. degradable magnesium alloy melt into the mold at 590-610° C., and use a pressure infiltration method to make the magnesium alloy melt flow into the pores of the TCP porous stent to obtain a two-phase continuous medical biodegradable magnesium alloy and TCP composite stent material, wherein the degradable magnesium alloy melt is a magnesium-zinc alloy, a magnesium-zinc-calcium alloy, or a magnesium-zinc-calcium-manganese alloy.

3. The preparation method according to claim 2, characterized in that: The ratio of the mass fraction of the TCP powder, the volume fraction of the resin, the mass fraction of the photoinitiator and the volume fraction of the dispersant is (52-53): (48-52): (0.77-0.79): (0.9-1.3), the unit of the mass fraction is g, and the unit of the volume fraction is mL.

4. The preparation method according to claim 2, characterized in that: The resin is a mixture of 1,6-hexanediol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate, and the ratio of 1,6-hexanediol diacrylate, 2-hydroxyethyl methacrylate and trimethylolpropane triacrylate is (29.5-30.5) by volume: (14.5～15.5)：(4.5～5.5)。 5. The preparation method according to claim 2, characterized in that: The photoinitiator is (2,4,6-trimethylbenzoyl)phosphine oxide, and the dispersant is a silane coupling agent; In the step 1, the rotation speed of the ball mill is 160 to 200 r / min, and the ball milling time is 19 to 21 hours.

6. The preparation method according to claim 2, characterized in that: The TCP-resin bracket is formed by connecting a plurality of unit cells, the shape of each unit cell is defined by a Gyroid function, the unit cell is 2 mm×2 mm×2 mm, and the wall thickness is -0.01 mm.

7. The preparation method according to claim 2, characterized in that: The TCP-resin bracket model was designed by the Gyroid function in the nTopology software, a three-periodic minimal surface structure was constructed, and then the TCP-resin bracket model was imported into the 3D printer for the 3D printing. The bias parameter of the Gyroid function model in the nTopology software was 0.

9.

8. The preparation method according to claim 2, characterized in that: The heating rate of the sintering is 0.2-1°C / min.

9. The preparation method according to claim 2, characterized in that: The method for preparing the degradable magnesium alloy melt comprises: mixing pure magnesium with a metal, smelting at 720-740° C. for at least 30 minutes, wherein the metal is one or more of zinc, calcium and manganese; When the degradable magnesium alloy melt is a magnesium-zinc alloy, the metal is zinc, and the content of zinc in the degradable magnesium alloy melt is 1.3-1.7 wt %; When the degradable magnesium alloy melt is a magnesium-zinc-calcium alloy, the metals are zinc and calcium, the zinc content in the degradable magnesium alloy melt is 1.3-1.7wt%, and the calcium content in the degradable magnesium alloy melt is 0.09-0.11wt%; When the degradable magnesium alloy melt is a magnesium-zinc-calcium-manganese alloy, the metals are zinc, calcium and manganese, the zinc content in the degradable magnesium alloy melt is 1.3-1.7wt%, the calcium content in the degradable magnesium alloy melt is 0.09-0.11wt%, and the manganese content in the degradable magnesium alloy melt is 0.18-0.22wt%.

10. The preparation method according to claim 2, wherein the pressure osmosis method comprises: Argon gas is introduced into the mold until the pressure inside the mold reaches 0.1-0.3 MPa, so that the degradable magnesium alloy melt can fully flow into the pores of the TCP porous scaffold.

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