Collagen antiseptic coating with tissue repair performance, method of preparation and use

Abstract

The application discloses a collagen anticorrosion coating with tissue repair performance, a preparation method and application, and belongs to the technical field of medical coatings. The application designs a novel self-repairing anticorrosion coating, and collagen is loaded on the coating through an excipient, so that the anticorrosion protection and biological function of the collagen anticorrosion coating are deeply integrated. The prepared collagen anticorrosion coating is firmly combined with a magnesium alloy substrate through a covalent bond, and no cracking and falling-off phenomenon occurs after being implanted in vivo for 4 weeks, the scaffold structure has good integrity, can provide sustained radial support, provides a sufficient time window for blood vessel repair, and provides a feasible scheme for clinical transformation of a new generation of biodegradable magnesium-based blood vessel stents.

Description

Technical Field

[0001] This invention belongs to the field of medical coating technology, specifically relating to a collagen anti-corrosion coating with tissue repair properties, its preparation method, and its uses. Background Technology

[0002] Percutaneous coronary intervention (PCI) combined with stent implantation has become the standard clinical strategy for restoring blood flow to occluded vessels. However, permanent metallic stents can cause complications such as in-stent restenosis, chronic inflammation, and late thrombosis, requiring long-term antiplatelet therapy and limiting their long-term clinical efficacy. Therefore, developing biodegradable vascular stents is the direction of development to overcome the above limitations. Biodegradable vascular stents provide temporary mechanical support and gradually degrade as the vessel heals.

[0003] Among numerous biodegradable stent materials, magnesium alloys stand out due to their unique advantages: magnesium is an essential trace element for the human body, participating in various physiological processes such as heart rhythm regulation and platelet aggregation inhibition, and exhibits excellent biocompatibility; the density of magnesium alloys is comparable to that of human bone and its Young's modulus, providing stable support for vascular stents to promote vascular healing. However, magnesium alloys have several limitations that urgently need to be addressed: magnesium has high chemical reactivity, making it prone to rapid and uneven corrosion in physiological environments, leading to premature loss of mechanical strength and collapse failure of the stent; hydrogen bubbles generated during degradation may interfere with the vascular healing process, triggering local inflammatory reactions; and its biocompatibility still needs further optimization to reduce thrombus formation and promote endothelial cell adhesion and proliferation.

[0004] Repair of vascular endothelial injury following stent implantation is a complex and dynamic process closely related to the spatiotemporal interactions of chemokines, growth factors, extracellular matrix (ECM) proteins, and transmembrane receptor-ligand interactions. Therefore, introducing specific functional factors capable of inducing endothelial regeneration onto the stent surface can effectively induce and promote endothelial cell adhesion and proliferation, thereby promoting endothelial repair and regeneration. However, most of the biomolecules used in this strategy lack high specificity, making it difficult to selectively regulate multiple pathological responses in the complex physiological environment of blood vessels. For example, extracellular matrix proteins can simultaneously promote the adhesion and growth of both ECs and SMCs. Furthermore, the strategy of endothelialization of the vascular stent surface can be directly applied to magnesium alloy vascular stent materials to achieve endothelial repair and regeneration.

[0005] However, as a reactive metal, magnesium alloys are susceptible to loss and degradation when the aforementioned methods are applied to their surface modification to promote endothelial regeneration. Biomolecules fixed on the surface are easily lost or degraded due to corrosion. Therefore, improving the corrosion resistance of magnesium alloys is essential for promoting in-situ regeneration and repair of vascular endothelium. The corrosion degradation behavior of magnesium alloys is closely related to the performance of their anti-corrosion coatings, while biocompatibility is closely related to their surface properties. Thus, simultaneously improving the corrosion resistance and biocompatibility of magnesium alloys through surface modification to promote in-situ repair and regeneration of damaged vascular endothelium is an important direction for improving magnesium alloy stents. Summary of the Invention

[0006] In view of the above-mentioned shortcomings in the prior art, the present invention provides a collagen anti-corrosion coating with tissue repair properties, a preparation method and applications.

[0007] To achieve the above objectives, the technical solution adopted by the present invention to solve its technical problem is as follows:

[0008] The purpose of this invention is to provide a method for preparing a collagen anti-corrosion coating with tissue repair properties, which includes the following steps:

[0009] (1) The magnesium alloy substrate material is pretreated to obtain a magnesium alloy with surface hydroxylation;

[0010] (2) The surface-hydroxylated magnesium alloy is immersed in solution I and dried. This is one cycle. After three cycles, it is dried under vacuum and high temperature to form an anti-corrosion coating (K / P). The solute in solution I includes copolymer P and copolymer K with a molar ratio of 1:1 to 1:2. The copolymer P is obtained by polymerization reaction of siloxane, phenylboronic acid and reducing agent. The copolymer K is obtained by polymerization reaction of siloxane coupling agent, long-chain alkyl and initiator.

[0011] (3) Immerse the magnesium alloy with anti-corrosion coating obtained in step (2) into solution II and dry it. This is one cycle. Repeat three cycles to obtain collagen anti-corrosion coating (K / PC). The solutes in solution II are excipients and collagen.

[0012] Further, in step (1), the hydroxylation process involves immersing the magnesium alloy substrate material in a 0.5-2 mol / L sodium hydroxide ethanol solution at 50-70℃ for 0.5-2 h, or placing the magnesium alloy substrate material in a plasma cleaner for 3-6 min.

[0013] Furthermore, in step (1), the magnesium alloy substrate material is pure magnesium, magnesium-iron alloy, magnesium-zinc alloy, magnesium-rare earth alloy, or magnesium-aluminum alloy.

[0014] Furthermore, in step (2), the reaction temperature for preparing copolymer P is 15~25℃ and the reaction time is 2-4 days; the reaction temperature for preparing copolymer K is 60~80℃ and the reaction time is 10-12 h.

[0015] Further, in step (2), the molar ratio of siloxane to phenylboronic acid is 1:2~5; wherein, the siloxane is at least one of bis(3-aminopropyl)-terminated polydimethylsiloxane and 3-mercaptopropyltriethoxysilane; the phenylboronic acid is at least one of 3-aminophenylboronic acid and 2-formylphenylboronic acid; and the reducing agent is at least one of sodium cyanoborohydride and sodium borohydride.

[0016] Furthermore, the silane coupling agent, long-chain alkyl group and initiator used in step (2) are all liquid phases with a volume ratio of 2~3:18~30:1;

[0017] The silane coupling agent is at least one of 3-(trimethoxysilyl)propyl methacrylate, methacryloyloxypropyltriethoxysilane, and acryloyloxypropyltrimethoxysilane; the long-chain alkyl group is at least one of laurate methacrylate, tridecyl methacrylate, and isodecanyl methacrylate; and the initiator is at least one of dimethyl azobisisobutyrate, azobisisoheptanenitrile, and azobisisobutyronitrile.

[0018] Furthermore, in step (2), the reaction solvent is at least one of anhydrous ethanol, methanol, isopropanol and dichloromethane.

[0019] Furthermore, in step (2), the magnesium alloy with surface hydroxylation is immersed in solution I for 5-25 min, and the drying conditions after immersion are 70-100℃ under vacuum for 2-24 h.

[0020] Further, the mass ratio of collagen to excipient in solution II is 0.5 to 1:1, and the final concentration of collagen is 1 to 5 mg / mL; the excipient is at least one of polylactic acid, polycaprolactone, polylactic-glycolic acid copolymer, silk fibroin, and poly(benzyl-L-glutamic acid).

[0021] Furthermore, the collagen was recombinant human type III collagen (rhCOL III), purchased from Shanxi Jinbo Biopharmaceutical Co., Ltd.

[0022] Furthermore, in step (3), the reaction solvent is at least one of isopropanol, hexafluoroisopropanol, methanol, acetonitrile, and dichloromethane.

[0023] Furthermore, in step (3), the magnesium alloy with the anti-corrosion coating is immersed in solution II for 10-120 s, and the drying conditions are 30-40℃ for 15-60 min.

[0024] Another object of the present invention is to provide a collagen anti-corrosion coating with tissue repair properties, which is prepared by the above method.

[0025] Another object of the present invention is to provide the use of the above-mentioned collagen anti-corrosion coating with tissue repair properties in the preparation of medical device materials.

[0026] The beneficial effects of this invention are:

[0027] 1. After the substrate material surface is hydroxylated, the terminal trimethoxysilyl group of the polymer K containing the silane coupling agent hydrolyzes to generate silanol group, which forms a stable Si-O-Mg covalent bond with the hydroxyl group on the substrate material surface. At the same time, adjacent silanol groups condense to form a Si-O-Si network. Meanwhile, the boric acid group in the copolymer P of silane coupling agent and long-chain alkane trimerizes during thermal crosslinking to form a reversible boron-oxygen six-membered ring (BO)3, thus constructing a K / P anti-corrosion self-healing coating with synergistic physical crosslinking and chemical bonding. This coating combines dense hydrophobic properties (water contact angle of 108.0°±0.9°) with water-triggered self-healing capabilities. On the one hand, it isolates body fluids from the substrate material through a physical barrier, significantly inhibiting the rapid corrosion of the substrate material (taking magnesium alloy as an example, the hydrogen evolution amount in 240 hours is only less than 1 / 3 of that of bare magnesium), avoiding inflammatory reactions caused by local alkalization and hydrogen release. On the other hand, the dynamic reversibility of the boron-oxygen six-membered ring bond can repair mechanical damage during the deployment of the stent, maintaining long-term corrosion protection and solving the problems of easy damage and protective failure of traditional static coatings.

[0028] 2. The recombinant human type III collagen (rhCOL III) selected in this invention is a customized functional material synthesized in biosynthesis. Its amino acid sequence is highly homologous to human collagen, contains the GER cell adhesion fragment and avoids platelet binding sites (without the GFOGER fragment), exhibiting extremely low immunogenicity and effectively reducing the risk of in vivo rejection of magnesium-based scaffolds. Unlike traditional collagen mixtures, this recombinant collagen primary structure does not contain hydroxyproline (O residue), avoiding immune responses and platelet aggregation problems caused by animal-derived amino acid groups. Simultaneously, its molecular structure has concentrated positive and negative charges (containing carboxyl, amino, and guanidine groups), exhibiting excellent water solubility and cell adhesion activity. It can promote the adhesion, proliferation, and migration of endothelial cells (ECs) while inhibiting excessive proliferation of smooth muscle cells (SMCs), achieving synergistic regulation of endothelialization and anti-inflammation. Within 3 days, the proliferation of endothelial cells is more than twice that of bare magnesium, and the smooth muscle cell density is significantly lower than that of the medical 316L stainless steel control group.

[0029] 3. This invention designs a novel self-healing anti-corrosion coating, and loads collagen onto the coating using excipients, achieving a deep integration of anti-corrosion protection and biological function. The K / P anti-corrosion self-healing coating maintains the mechanical integrity of the scaffold by inhibiting magnesium alloy degradation. The third functional layer optimizes biocompatibility through multiple mechanisms: First, it enhances blood compatibility, with a hemolysis rate of less than 5%, significantly reduces platelet adhesion and inhibits platelet activation (adhered platelets exhibit an inactive disc-like morphology), and no significant thrombus formation occurs under dynamic blood flow conditions. Second, it regulates the immune inflammatory response by inhibiting the NF-κB signaling pathway to reduce macrophage infiltration (decreased CD68 expression), promoting macrophage polarization towards the anti-inflammatory M2 phenotype, increasing the expression of the anti-inflammatory factor IL-10, and reducing the release of the pro-inflammatory factor TNF-α. The thickness of the fibrous capsule after subcutaneous implantation for 4 weeks is only 67.5±4.5μm, far lower than that of the bare magnesium group (152.5±11.4μm). Third, it accelerates vascular repair. After in vivo implantation, the CD31-positive endothelial coverage rate on the stent surface reaches 36.81±3.20%, and the eNOS expression level is as high as 84.77±5.60%, significantly better than bare magnesium and single K / P coating, achieving rapid and functional reendothelialization.

[0030] 3. The preparation process of this invention has the advantages of being simple, controllable, and scalable: the K / P coating is prepared through a "immersion-thermal crosslinking" cycle, which is simple to operate and produces a uniform coating thickness (approximately 20 μm), without affecting the gripping and expansion performance of the stent; rhCOLIII is loaded using ultrasonic spraying technology with mild process parameters (35-45℃), which avoids collagen denaturation, and achieves stable fixation through multiple spraying-drying cycles. Even after 7 days of PBS immersion and rinsing, it still maintains a significant fluorescence signal, ensuring long-lasting bioactivity. Furthermore, the coating is firmly bonded to the magnesium alloy substrate through covalent bonds, and no cracking or detachment was observed after 4 weeks of in vivo implantation. The stent structure maintains good integrity, providing continuous radial support and a sufficient time window for vascular repair, thus providing a feasible solution for the clinical translation of a new generation of biodegradable magnesium-based vascular stents. Attached Figure Description

[0031] Figure 1 Electron micrographs of magnesium alloy supports coated with different coatings;

[0032] Figure 2The following are performance characterization diagrams of the collagen anti-corrosion coating prepared in Example 1: A is the attenuated total reflectance Fourier transform infrared spectrum; B is the X-ray photoelectron spectroscopy full spectrum; C is the surface elemental composition analysis diagram of K, K / P, and K / PC coatings; D is the high-resolution carbon 1s spectrum of the K / PC coating; E is the water contact angle detection result of the coating; F is a typical scanning electron microscope image of the bare magnesium alloy substrate, K coating, K / P coating, and K / PC coating; G is the elemental distribution diagram of the energy dispersive spectroscopy; H is a cross-sectional scanning electron microscope image of the K / P coating; I is the fluorescence image of the K / PC coating labeled with fluorescein isothiocyanate after immersion in PBS for 0 days and 7 days.

[0033] Figure 3 The figures show the electrochemical corrosion resistance test results of the collagen anticorrosion coating prepared in Example 1; where A is the potentiodynamic polarization curve; B and C are the electrochemical impedance spectroscopy curves of bare magnesium and the anticorrosion coated magnesium alloy; D is the hydrogen evolution analysis diagram; E is the pH value analysis diagram; F is the optical image after immersion for 240 h; G is the SEM image after immersion; and H is the scanning electron microscope image after the immersion degradation experiment.

[0034] Figure 4 The images show the effects of the collagen antiseptic coating prepared in Example 1 on blood function; where A is a scanning electron microscope image of whole blood adhesion on the surface of different samples; B is a scanning electron microscope image of platelet-rich plasma adhesion on the surface of different samples; C is a statistical representation of the hemolysis rate of different samples; D is a platelet density analysis image; E is a schematic diagram of the in vitro arteriovenous shunting experiment; F is a cross-sectional image of different samples; and G is a graph showing the residual mass results of different samples.

[0035] Figure 5 The images show fluorescence and endothelial cell viability assays of the collagen antiseptic coating prepared in Example 1 after culturing endothelial cells for 1 day and 3 days. Specifically, A represents fluorescence after co-culturing with the extract; B represents fluorescence after direct contact culturing on different sample surfaces; C represents fluorescence after culturing smooth muscle cells directly on different sample surfaces for 24 h and 72 h; D represents viability assay after co-culturing with the extract; E represents viability assay after direct contact culturing on different sample surfaces; F represents viability assay after culturing smooth muscle cells for 1 day and 3 days; G represents fluorescence after direct culturing of mouse macrophages on different sample surfaces; and H represents viability assay after direct contact culturing on different sample surfaces.

[0036] Figure 6The images show the results of a subcutaneous implantation experiment using the collagen antiseptic coating prepared in Example 1. A is a schematic diagram of the subcutaneous implantation experiment; B is a hematoxylin-eosin staining image of the fibrous capsules 4 weeks after subcutaneous implantation of different samples; C is a detection image of fibrous capsule thickness; D is an immunofluorescence staining image of CD68; E is an immunofluorescence staining image of IL-10 and TNF-α; F is a detection image of CD68 expression level; G is a detection image of IL-10 expression level; and H is a detection image of TNF-α expression level.

[0037] Figure 7 The images show the results of in vivo implantation experiments of stents containing the collagen antiseptic coating prepared in Example 1; where A is a schematic diagram of stent implantation in rabbits; B is a morphological image of the stent; C is a diagram showing the compression and expansion changes of the stent on a balloon; D is a scanning electron microscope image of the stent; E is a typical image of stent implantation in the abdominal aorta; F is an image of a vascular stent sample 4 weeks after stent implantation; G is a three-dimensional model of the stent reconstructed by high-resolution computed tomography microscopy; H is a hematoxylin-eosin staining image of the cross-section of the stent-containing vessel; I is a scanning electron microscope image of the inner surface of the vascular stent; J is an immunofluorescence image; K is a detection image of CD31 expression levels on the surface of coated stents and bare magnesium stents prepared by K / P and K / PC; L is a detection image of eNOS expression levels on the surface of coated stents and bare magnesium stents prepared by K / P and K / PC. Detailed Implementation

[0038] The specific embodiments of the present invention are described below to enable those skilled in the art to understand the present invention. However, it should be understood that the present invention is not limited to the scope of the specific embodiments. For those skilled in the art, various changes are obvious as long as they are within the spirit and scope of the present invention as defined and determined by the appended claims. All inventions utilizing the concept of the present invention are protected.

[0039] Example 1

[0040] A method for preparing a collagen anti-corrosion coating with tissue repair properties, comprising the following steps:

[0041] ①Preparation of copolymers for anti-corrosion coatings

[0042] (1) Synthesis of copolymer P: 50 parts by mass of diaminopropyl-terminated polydimethylsiloxane, 150 parts by mass of 2-formylphenylboronic acid, and 300 parts by mass of sodium cyanoborohydride were placed in a three-necked flask, and 25 mL of anhydrous ethanol was added as a solvent. After stirring continuously for 3 days, the organic phase was washed successively with sodium bicarbonate solution, deionized water, and saturated sodium chloride solution. After concentration under reduced pressure, it was redissolved in 100 mL of dichloromethane. The organic phase was washed again in the above manner. The obtained organic layer was dried with anhydrous magnesium sulfate, filtered, concentrated under vacuum, and finally dried in an oven at 80 °C for 12 h to obtain a colorless, transparent, viscous solid.

[0043] (2) Synthesis of copolymer K: Take 60 parts by mass of 3-(trimethoxysilyl)propyl methacrylate, 50 parts by mass of butyl methacrylate and 10 parts by mass of dimethyl azobisisobutyrate into a three-necked flask, add 25 mL of isopropanol as solvent, perform nitrogen purging 5 times, seal the three-necked flask, and then stir the reaction at 70°C for 12 h to obtain a viscous liquid, which is obtained by vacuum distillation.

[0044] ②Preparation of corrosion-resistant and collagen-functionalized magnesium-based metal coatings

[0045] S1: The cleaned magnesium-aluminum alloy support was immersed in a 1M sodium hydroxide ethanol solution and reacted at 70°C for 90 min. Then the reacted support was completely immersed in isopropanol for solvent replacement. The support was then placed in a 50°C oven to dry for 10 min. After drying, a magnesium-aluminum alloy support with surface hydroxylation was obtained.

[0046] S2: Dissolve copolymer P and copolymer K in ethanol to obtain solution I with a concentration of 5wt%. Then, immerse the surface hydroxylated magnesium-aluminum alloy support in solution I. After the solution I completely covers the support, soak for 5 min. Then, take out the support and place it in an 80℃ oven to dry for 5 h. This is one cycle. Repeat three cycles. After drying, an anti-corrosion coating is obtained on the surface of the magnesium-aluminum alloy support.

[0047] S3: Dissolve collagen and polylactic acid in hexafluoroisopropanol to obtain solution II with a collagen concentration of 1 mg / mL (the mass ratio of collagen to polylactic acid in the solution is 1:1). Immerse the magnesium-aluminum alloy scaffold with anti-corrosion coating from step S2 in solution II. After solution II completely covers the scaffold, soak for 30 seconds. Remove the scaffold and dry it in an oven at 35°C for 30 minutes. The dried scaffold is then obtained.

[0048] Example 2

[0049] A method for preparing a collagen anti-corrosion coating with tissue repair properties, comprising the following steps:

[0050] ①Preparation of copolymers for anti-corrosion coatings

[0051] (1) Synthesis of copolymer P: 50 parts by mass of 3-mercaptopropyltriethoxysilane, 120 parts by mass of 2-formylphenylboronic acid, and 300 parts by mass of sodium cyanoborohydride were placed in a three-necked flask, and 25 mL of anhydrous ethanol was added as a solvent. The mixture was stirred continuously overnight. The organic phase was washed successively with sodium bicarbonate solution, deionized water, and saturated sodium chloride solution. After concentration under reduced pressure, it was redissolved in 100 mL of dichloromethane. The organic phase was washed again using the above method. The resulting organic layer was dried with anhydrous magnesium sulfate, filtered, concentrated under vacuum, and finally dried in an oven at 80 °C for 12 h to obtain a viscous solid.

[0052] (2) Synthesis of copolymer K: Take 60 parts by mass of 3-(trimethoxysilyl)propyl methacrylate, 50 parts by mass of butyl methacrylate and 10 parts by mass of dimethyl azobisisobutyrate into a three-necked flask, add 25 mL of isopropanol as solvent, perform nitrogen purging 5 times, seal the three-necked flask, and then stir the reaction at 70°C for 12 h to obtain a viscous liquid, which is obtained by vacuum distillation.

[0053] ②Preparation of corrosion-resistant and collagen-functionalized magnesium-based metal coatings

[0054] S1: The cleaned magnesium-aluminum alloy support was immersed in a 1M sodium hydroxide ethanol solution and reacted at 70°C for 90 min. Then the reacted support was completely immersed in isopropanol for solvent replacement. The support was then placed in a 50°C oven to dry for 10 min. After drying, a magnesium-aluminum alloy support with surface hydroxylation was obtained.

[0055] S2: Dissolve copolymer P and copolymer K in ethanol to obtain solution I with a concentration of 5wt%. Then, immerse the surface hydroxylated magnesium-aluminum alloy support in solution I. After solution I completely covers the support, soak for 10 minutes. Take out the support and place it in an 80℃ oven to dry for 5 hours. This is one cycle. Repeat three cycles. After drying, an anti-corrosion coating is obtained on the surface of the magnesium-aluminum alloy support.

[0056] S3: Dissolve collagen and polylactic acid in hexafluoroisopropanol to obtain solution II with a collagen concentration of 1 mg / mL (the mass ratio of collagen to polylactic acid in the solution is 1:1). Immerse the magnesium-aluminum alloy scaffold with anti-corrosion coating from step S2 in solution II. After solution II completely covers the scaffold, soak for 15 seconds. Remove the scaffold and place it in a 35℃ oven to dry for 30 minutes. The dried scaffold is then obtained.

[0057] Example 3

[0058] A method for preparing a collagen anti-corrosion coating with tissue repair properties, comprising the following steps:

[0059] ①Preparation of copolymers for anti-corrosion coatings

[0060] (1) Synthesis of copolymer P: 50 parts by mass of 3-mercaptopropyltriethoxysilane, 100 parts by mass of 3-aminophenylboronic acid, and 400 parts by mass of sodium borohydride were placed in a three-necked flask, and 25 mL of anhydrous ethanol was added as a solvent. The mixture was stirred continuously overnight. The organic phase was washed successively with sodium bicarbonate solution, deionized water, and saturated sodium chloride solution. After concentration under reduced pressure, it was redissolved in 100 mL of dichloromethane. The organic phase was washed again using the above method. The resulting organic layer was dried with anhydrous magnesium sulfate, filtered, concentrated under vacuum, and finally dried in an oven at 80 °C for 12 h to obtain a colorless, transparent, viscous solid.

[0061] (2) Synthesis of copolymer K: Take 60 parts by mass of 3-(trimethoxysilyl)propyl methacrylate, 90 parts of methyl laurate and 10 parts of dimethyl azobisisobutyrate in a three-necked flask, add 25 mL of isopropanol as solvent, perform nitrogen purging 5 times, seal the three-necked flask, and then stir the reaction at 70 °C for 12 h to obtain a viscous liquid, which is then obtained by vacuum distillation.

[0062] ②Preparation of corrosion-resistant and collagen-functionalized magnesium-based metal coatings

[0063] S1: The cleaned magnesium-aluminum alloy support was immersed in a 1M sodium hydroxide ethanol solution and reacted at 70°C for 90 min. Then the reacted support was completely immersed in isopropanol for solvent replacement. The support was then placed in a 50°C oven to dry for 10 min. After drying, a magnesium-aluminum alloy support with surface hydroxylation was obtained.

[0064] S2: Dissolve copolymer P and copolymer K in ethanol to obtain a 5wt% solution I. Then, immerse the surface-hydroxylated magnesium-aluminum alloy support in solution I. After the solution I completely covers the support, soak for 10 min. Take out the support and place it in an 80℃ oven to dry for 60 min. This is one cycle. Repeat three cycles. After drying, an anti-corrosion coating is obtained on the surface of the magnesium-aluminum alloy support.

[0065] S3: Dissolve collagen and polylactic acid in hexafluoroisopropanol to obtain solution II with a collagen concentration of 1 mg / mL (the mass ratio of collagen to polylactic acid in the solution is 1:1). Immerse the magnesium-aluminum alloy scaffold with anti-corrosion coating from step S2 in solution II. After solution II completely covers the scaffold, remove it and place it in a 35°C oven to dry for 5 min. The dried product is then obtained.

[0066] Example 4

[0067] A method for preparing a collagen anti-corrosion coating with tissue repair properties, comprising the following steps:

[0068] ①Preparation of copolymers for anti-corrosion coatings

[0069] (1) Synthesis of copolymer P: 50 parts by mass of bis(3-aminopropyl)-terminated polydimethylsiloxane, 120 parts by mass of 2-formylphenylboronic acid, and 400 parts by mass of sodium borohydride were placed in a three-necked flask, and 25 mL of anhydrous ethanol was added as a solvent. The mixture was stirred continuously overnight. The organic phase was washed successively with sodium bicarbonate solution, deionized water, and saturated sodium chloride solution. After concentration under reduced pressure, it was redissolved in 100 mL of dichloromethane. The organic phase was washed again using the above method. The resulting organic layer was dried with anhydrous magnesium sulfate, filtered, concentrated under vacuum, and finally dried in an oven at 80 °C for 12 h to obtain a colorless, transparent, viscous solid.

[0070] (2) Synthesis of copolymer K: Take 60 parts by mass of 3-(trimethoxysilyl)propyl methacrylate, 90 parts of methyl laurate and 10 parts of dimethyl azobisisobutyrate in a three-necked flask, add 25 mL of isopropanol as solvent, perform 5 nitrogen purgings, seal the three-necked flask, and then stir the reaction at 70°C for 12 h to obtain a viscous liquid, which is then obtained by vacuum distillation.

[0071] ②Preparation of corrosion-resistant and collagen-functionalized magnesium-based metal coatings

[0072] S1: The cleaned magnesium-aluminum alloy bracket is surface-treated with a plasma machine for 4 minutes to achieve surface hydroxylation;

[0073] S2: Copolymer P and copolymer K are dissolved in ethanol to obtain solution I with a concentration of 2.5wt%. Then, the surface hydroxylated magnesium-aluminum alloy support is immersed in solution I. After solution I completely covers the support, it is soaked for 20 min. The support is then removed and placed in an 80℃ oven to dry for 5 h. This is one cycle. Three cycles are repeated. After drying, an anti-corrosion coating is obtained on the surface of the magnesium-aluminum alloy support.

[0074] S3: Dissolve collagen and polylactic acid-glycolic acid copolymer in hexafluoroisopropanol to obtain solution II with a collagen concentration of 4 mg / mL (the mass ratio of collagen and polylactic acid-glycolic acid copolymer in the solution is 1:1). Immerse the magnesium-aluminum alloy scaffold with anti-corrosion coating from step S2 in solution II. After solution II completely covers the scaffold, soak for 30 s. Remove the scaffold and place it in a 35℃ oven to dry for 30 min. The dried product is then obtained.

[0075] Example 5

[0076] A method for preparing a collagen anti-corrosion coating with tissue repair properties, comprising the following steps:

[0077] ①Preparation of copolymers for anti-corrosion coatings

[0078] (1) Synthesis of copolymer P: 50 parts by mass of diaminopropyl-terminated polydimethylsiloxane, 150 parts by mass of 2-formylphenylboronic acid, and 300 parts by mass of sodium cyanoborohydride were placed in a three-necked flask, and 25 mL of anhydrous ethanol was added as a solvent. The mixture was stirred continuously overnight. The organic phase was washed successively with sodium bicarbonate solution, deionized water, and saturated sodium chloride solution. After concentration under reduced pressure, it was redissolved in 100 mL of dichloromethane. The organic phase was washed again using the above method. The resulting organic layer was dried with anhydrous magnesium sulfate, filtered, concentrated under vacuum, and finally dried in an oven at 100 °C for 12 h to obtain a colorless, transparent, viscous solid.

[0079] (2) Synthesis of copolymer K: Take 60 parts by mass of 3-(trimethoxysilyl)propyl methacrylate, 50 parts by mass of butyl methacrylate and 20 parts by mass of azobisisoheptanenitrile in a three-necked flask, add 25 mL of isopropanol as solvent, perform nitrogen purging 5 times, seal the three-necked flask, and then stir the reaction at 70℃ for 12 h to obtain a viscous liquid, which is obtained by vacuum distillation.

[0080] ②Preparation of corrosion-resistant and collagen-functionalized magnesium-based metal coatings

[0081] S1: After cleaning, the magnesium-aluminum alloy bracket is surface treated with a plasma machine for 5 minutes to achieve surface hydroxylation;

[0082] S2: Dissolve copolymer P and copolymer K in ethanol to obtain solution I with a concentration of 2.5wt%. Then, immerse the surface hydroxylated magnesium-aluminum alloy support in solution I. After the solution I completely covers the support, soak for 15 min. Take out the support and place it in an 80℃ oven to dry for 5 h. This is one cycle. Repeat three cycles. After drying, an anti-corrosion coating is obtained on the surface of the magnesium-aluminum alloy support.

[0083] S3: Dissolve collagen and polylactic acid-glycolic acid copolymer in hexafluoroisopropanol to obtain solution II with a collagen concentration of 4 mg / mL (the mass ratio of collagen and polylactic acid-glycolic acid copolymer in the solution is 1:2). Immerse the magnesium-aluminum alloy scaffold with anti-corrosion coating from step S2 in solution II. After solution II completely covers the scaffold, soak for 30 s. Remove the scaffold and place it in a 35℃ oven to dry for 30 min. The dried product is then obtained.

[0084] control group

[0085] Single-component K-coatings and P-coatings were prepared using the same method as in Example 1, and were used as controls.

[0086] Statistical analysis

[0087] In this invention, all experiments were conducted with at least three parallel groups, and experimental data are expressed as mean ± standard deviation. Data between groups were analyzed using one-way ANOVA combined with Tafel post-hoc test, or two-way ANOVA combined with Westdock multiple comparison test; statistical significance was indicated by *P<0.05, **P<0.01, and ***P<0.001.

[0088] The following tests were conducted on the various types of materials prepared in Example 1.

[0089] 1. Coating Characterization

[0090] The prepared P was dissolved in deuterated methanol and subjected to hydrogen nuclear magnetic resonance (NMR). 1 H NMR) tests were performed, and the test results were analyzed using MestReNova 14 software.

[0091] A coating was prepared on the magnesium sheet surface using the method described above for subsequent characterization. The characteristic absorption peaks of the coating were detected using a Nicolet iS50 attenuated total reflectance Fourier transform infrared spectrometer (ATR-FTIR) from Thermo Fisher Scientific, USA, in the wavenumber range of 4000–500 cm⁻¹. -1 The elemental composition and chemical state of the coating were determined using a K-ALPHA X-ray photoelectron spectroscopy (XPS) instrument from Thermo Fisher Scientific (USA); the water contact angle (WCA) was measured using an Attension Theta contact angle meter from Bio-Electronics (Sweden); after the samples were thoroughly dried and sputter-coated with gold, the surface and cross-sectional morphology of the samples were observed using an Apreo scanning electron microscope (SEM) from Thermo Fisher Scientific (USA); the elemental distribution of the coating was observed using an Ultim Max65 energy dispersive spectrometer (EDS) from Oxford Instruments; and the water contact angle (WCA) of the samples was measured again using a Theta Lite static water contact angle analyzer from Bio-Electronics (Sweden). The results are shown in [Figure number missing]. Figure 2 .

[0092] like Figure 2 As shown in Figure A, the P coating is at 749 cm. -1 A characteristic peak appears at 799 cm⁻¹, corresponding to the boron-oxygen six-membered ring (BO)₃. -1 The characteristic peak at 1094 cm⁻¹ is the vibrational peak of the silanine (Si-(CH₂)₂) in the PDMS backbone. The K / P composite coating at 1094 cm⁻¹... -1 and 1019cm -1 The presence of characteristic stretching vibration peaks of silicon-oxygen-silicon (Si-O-Si) at 1723 cm⁻¹ indicates the formation of a cross-linked siloxane network; furthermore, the peak at 1723 cm⁻¹... -1A new absorption peak appeared at 1638 cm⁻¹, corresponding to the stretching vibration of the carbonyl group (C=O), further confirming that the KH570-BMA component had been successfully incorporated into the coating. After the introduction of recombinant human type III collagen (rhCOL III), the coating showed an absorption peak at 1638 cm⁻¹. -1 and 1538 cm -1 Two broad peaks appeared at the point, corresponding to the amide I band (C=O stretching vibration) and the amide II band (NH bending vibration), respectively, indicating that the collagen had been successfully fixed on the coating surface.

[0093] The surface wettability of the coating is evaluated by measuring the water contact angle (WCA). Figure 2 Compared to the bare magnesium alloy surface (46.0°±3.7°), the hydrophobicity of both the K coating and the K / P coating was improved, with water contact angles of 87.9°±2.3° and 108.0°±0.9°, respectively. The improved hydrophobicity of the K / P coating was attributed to the incorporation of phosphorus (P). After collagen modification, the water contact angle of the coating decreased to 60.8°±4.5°, reflecting the alteration of coating wettability by the hydrophilic functional groups introduced by recombinant human type III collagen.

[0094] Further analysis of the surface chemical composition of the coating was performed using XPS. Figure 2 (B) confirmed the successful incorporation of silicon, boron, and collagen-related components into the coating; weak nitrogen (N) signals were detected in both the K / P and K / PC coatings in the high-resolution spectrum of 408–393 eV. Elemental quantitative analysis ( Figure 2 In the K / P / PC coating, no nitrogen was detected, while the nitrogen content of the collagen-modified coating increased significantly, confirming the successful incorporation of recombinant human type III collagen. Simultaneously, a large amount of silicon was detected in both the K / P and K / PC coatings, indicating the stable existence of the siloxane network within the coatings. High-resolution carbon 1s spectroscopy (High-resolution carbon 1s spectroscopy) Figure 2 Further analysis (D) shows that, compared to the K / P coating, the relative strength of oxygen carbonyl groups (OC=O) and carbon-oxygen bonds (CO) in the K / PC coating is increased, which is consistent with the introduction of collagen-related functional groups. These results collectively confirm that the composite coating and subsequent collagen functionalization modifications were successfully achieved.

[0095] The surface morphology of magnesium alloy and coating samples was characterized using scanning electron microscopy (SEM). Figure 2 The results showed that the coating had a uniform and dense morphology and good adhesion to the substrate interface. This was attributed to the silicon-oxygen-magnesium (Si-O-Mg) covalent bonds formed between the silane groups and the surface of the hydroxylated magnesium alloy. The incorporation of P and subsequent collagen modification further improved the continuity and density of the coating.

[0096] Energy dispersive spectrometer (EDS) elemental distribution mapping Figure 2The SEM image (G) shows that silicon and boron are uniformly distributed in the K / P coating, further confirming that phosphorus (P) has been successfully integrated into the coating matrix. The SEM image of the coating cross-section (G) shows... Figure 2 As shown in the figure (H), the coating thickness is approximately 20 μm, which will not have a significant impact on the diameter of the support in practical applications.

[0097] To evaluate the stability of the collagen layer, recombinant human type III collagen labeled with fluorescein isothiocyanate (FITC) was introduced into the coating using the same method. Figure 2 (I) Strong and uniform fluorescence signals were observed on the surface of the freshly prepared K / PC coating. After soaking in PBS for 7 days, the coating still retained a large amount of fluorescence intensity, indicating that the collagen component can be stably present on the coating surface, ensuring that it can perform its biological function in the early stage after implantation.

[0098] 2. Electrochemical testing

[0099] The electrochemical workstation of Shanghai Chenhua (CHI660E) was used for testing. A three-electrode system was employed in phosphate-buffered saline (PBS) at 37.0 ± 0.5 °C. The counter electrode was a platinum electrode, the reference electrode was a saturated calomel electrode (CHI150), and the working electrode was the sample. The back of the sample was connected to a copper wire using conductive silver paste, and the back and sides of the sample were sealed with silicone rubber. Before testing, the sample was immersed in PBS buffer for 10 min to stabilize the open circuit potential. Each sample was tested at least three times. The frequency range of electrochemical impedance spectroscopy (EIS) was 10 Hz. 5 ~10 -1 The Hz frequency was measured with a sinusoidal perturbation signal of 20 mV. Electrochemical polarization curves were measured using potentiodynamic polarization (PDP) at a scan rate of 1 mV / s, with a scan range of -2.0 V to -1 V. The corrosion potential (Ecorr) and corrosion current (Icorr) of the sample were calculated using Tafel extrapolation. Electrochemical impedance spectroscopy (EIS) was then performed again at open-circuit potential, with a frequency range of 100–10 Hz. 5 The Hz frequency and the sinusoidal perturbation signal were 20 mV. SEM was used to observe the changes in the surface morphology of the samples after the experiment.

[0100] Potential dynamic polarization (PDP) curves are as follows Figure 3 As shown in Figure A, among all samples, the K / P coating exhibits the best corrosion resistance, with a corrosion potential of -1.51 V and a significantly reduced corrosion current density (≈10). -5.75 A·cm -2 This indicates that magnesium dissolution was effectively inhibited. After collagen modification, the corrosion resistance of the K / PC coating decreased slightly (corrosion potential = -1.53 ​​V, corrosion current density ≈ 10). -5.75 A·cm -2 This may be due to the reduction of hydrophobicity on the coating surface during collagen incorporation.

[0101] Although the single-component K coating exhibited a relatively high corrosion potential, its corrosion current density was only slightly lower than that of bare magnesium, indicating limited protective capability. This may be due to the thinner K coating, making it more susceptible to damage under applied potential. In contrast, the K / P and K / PC coatings showed corrosion current densities reduced by more than an order of magnitude, demonstrating excellent corrosion protection performance. The surface morphology of the samples after electrochemical testing was observed using SEM. Figure 3 In the case of G), the bare magnesium surface showed severe corrosion damage, with a large amount of loose and porous corrosion products accumulating on the surface; while the coated samples (K, K / P, K / PC) still retained a relatively complete surface structure, with only local swelling and partial degradation observed, indicating that the coating has a significant anti-corrosion effect, but cannot completely inhibit corrosion.

[0102] Electrochemical impedance spectroscopy (EIS) results are as follows Figure 3 As shown in Figures B and C, the Nyquist plots reveal that the capacitive arc diameters of the K / P and K / PC coatings are significantly increased compared to bare magnesium and the K coating, indicating improved charge transfer resistance and barrier performance, consistent with the results of potentiodynamic polarization analysis. In summary, the K / P composite coating significantly enhances the corrosion resistance of magnesium alloys. Although subsequent collagen modification slightly reduces the hydrophobicity of the coating, it still maintains effective electrochemical protection.

[0103] In addition to electrochemical tests, this study conducted a long-term immersion experiment in PBS to evaluate the degradation behavior of samples under simulated physiological conditions. To determine the exposed area of ​​the samples, the back and sides of the samples were sealed with silicone rubber before immersion. During immersion, magnesium underwent a corrosion reaction: Mg + H₂O → Mg(OH)₂ + H₂↑, producing hydrogen gas and causing the solution pH to rise. Therefore, this study used the cumulative hydrogen evolution and pH change as evaluation indicators of degradation behavior. Figure 3 (D and E in the middle).

[0104] like Figure 3 As shown in Figure D, the amount of hydrogen evolved by bare magnesium increases rapidly, with a cumulative hydrogen evolution volume of approximately 23 mL·cm³ within 240 h. -2 In contrast, the hydrogen evolution rate of all coated samples was significantly reduced, with the K / P and K / PC coatings showing extremely low degradation within the initial 120 h, and the cumulative hydrogen evolution within 240 h being less than 1 / 3 of that of bare magnesium. The trend of solution pH change was consistent with the hydrogen evolution rate. Figure 3 (E), further confirming the inhibitory effect of the coating on the corrosion of magnesium alloys.

[0105] 3. In vitro immersion experiment

[0106] Similarly, the samples were immersed in PBS solution in vitro, with each group taking approximately 2 cm² of total surface area. 2Four samples were placed in 200 mL of PBS solution with pH=7.4 at an experimental temperature of 37.0±0.5℃. Before immersion, the back and sides of the samples were sealed with silicone rubber, and three parallel samples were set up for each group. The hydrogen evolution volume of each group of samples was recorded every 12 h, and the pH value of the immersion solution was measured with a pH meter. The total immersion time was 240 h. After the immersion experiment, the surface corrosion of the samples was observed using SEM.

[0107] Optical images of the sample after immersion ( Figure 3 (F) showed that the bare magnesium surface was covered with a large amount of white corrosion products, the K coating showed localized corrosion, while the K / P and K / PC coatings showed no obvious large-area degradation. SEM observation results ( Figure 3 The above conclusions were further verified by the H-test: the bare magnesium surface was severely corroded, with a large number of porous degradation products; while the K / P coating and K / PC coating maintained their structural integrity, with no obvious accumulation of corrosion products. In summary, the immersion test results show that the K / P coating and K / PC coating can effectively inhibit the degradation of magnesium alloys and provide long-term protection for the substrate under physiological conditions.

[0108] 4. Blood test

[0109] (1) Blood compatibility test

[0110] Fresh rabbit whole blood used in the experiment was obtained from the Experimental Animal Center of Sichuan University, and all blood compatibility tests were conducted in accordance with the "Ethical Guidelines for Experimental Animals of Sichuan University". Blood was collected in blood collection tubes containing 10% sodium citrate, centrifuged at 1500 r / min for 15 min, and the supernatant was taken as platelet-rich plasma (PRP); the lower layer of red blood cells was repeatedly washed with physiological saline and used for hemolysis experiments.

[0111] First, the sample was incubated with fresh rabbit whole blood and platelet-rich plasma (PRP) for 1 hour to evaluate its blood compatibility. Figure 4 (A and B). The bare magnesium surface showed obvious corrosion and surface cracking after incubation, while the coated sample maintained a relatively intact surface morphology, indicating that the coating can still provide effective corrosion protection under blood contact conditions.

[0112] (2) Platelet adhesion test

[0113] To simulate platelet adhesion and activation on the stent surface after in vivo implantation of vascular stent materials, an in vitro static platelet adhesion experiment was conducted. Samples were placed in 24-well plates, washed twice with physiological saline, and 1 mL of platelet-rich plasma (PRP) was added to each well. The plates were incubated at 37°C for 60 min. The sample surface was then gently rinsed with physiological saline, fixed overnight at 4°C with 2.5% glutaraldehyde solution, and finally dehydrated. The number and morphology of platelets adhering to the sample surface were observed using SEM.

[0114] Quantitative analysis of platelet adhesion ( Figure 4 A large number of platelets were observed adhering to the surfaces of the D), K coating and K / P coating, indicating that their antithrombotic properties are limited; while the amount of platelets adhering to the surface of the K / PC coating is extremely small, and the few adhering platelets still maintain a circular disc shape, indicating that the platelets are in an inactive state, and the blood compatibility of the coating is improved.

[0115] (3) Hemolysis test

[0116] Hemolysis rate (HR) is used to detect the aggregation, rupture, and lysis of erythrocytes after in vitro interaction between biomaterials and human erythrocytes. According to international standards (ISO) and Chinese national standards, biomaterials with a hemolysis rate of less than 5% meet the requirements for clinical medical biomaterials. The specific procedure is as follows: Centrifuged erythrocytes were prepared into a 2% suspension using ultrapure water as the positive control; a 2% erythrocyte suspension was prepared using physiological saline as the negative control. Samples were placed in 24-well plates, each sample was pre-washed twice with 1 mL of physiological saline, followed by the addition of 1 mL of the 2% erythrocyte suspension prepared with physiological saline, and incubated at 37°C for 60 min. The liquid in the wells was collected, centrifuged at 3000 r / min for 5 min, and 120 μL of the supernatant was taken and its absorbance was measured at 540 nm.

[0117] Formula for calculating hemolysis rate:

[0118]

[0119] Among them, A sample A represents the absorbance of the solution on the surface of the experimental sample. negative A represents the absorbance of the negative control sample. positive The absorbance is the value of the positive control sample.

[0120] Results of hemolysis test ( Figure 4 The results (C) show that, except for the bare magnesium group which showed obvious hemolysis, the hemolysis rate of all coated samples was less than 5%, which meets the standards for blood contact materials; among them, the hemolysis rate of K / P coating and K / PC coating was lower than that of K coating, further indicating that they have better blood compatibility.

[0121] (4) Evaluation of in vitro thrombotic activity

[0122] To simulate the interaction between the material and flowing blood after implantation, an in vitro arteriovenous shunting experiment (ex vivo antithrombotic experiment) was conducted. A coating was prepared on the surface of magnesium foil using the method described above. The sample was rolled into a cylindrical shape and placed in a polyvinyl chloride (PVC) tube, with three parallel samples per group. Rabbits were anesthetized, and the left common carotid artery and right internal jugular vein were separated. The PVC tube containing the sample was connected to the rabbit's common carotid artery and internal jugular vein using a puncture needle to form a circulatory loop. After 1 hour of circulation, the PVC tube containing the sample was cut, and the sample surface was gently rinsed with physiological saline to remove thrombi. The remaining mass of the material was then weighed.

[0123] An in vitro arteriovenous shunt model was used to evaluate the antithrombotic properties of the samples under dynamic conditions. Figure 4 (E). After 1 hour of blood circulation, significant thrombus formation was observed on the surface of the K / P coated sample, leading to partial luminal obstruction. Figure 4 (F); while no obvious thrombus formation was observed on the surface of the bare magnesium and K / PC coated samples. The remaining mass of the samples after gentle rinsing ( Figure 4 The results were further verified by the G test: the bare magnesium sample suffered a large amount of mass loss due to rapid degradation and shedding under flowing conditions, while the K / P coated and K / PC coated samples maintained structural integrity and relatively stable mass retention.

[0124] In summary, the results indicate that collagen-functionalized coatings (K / PC) can significantly improve the blood compatibility and antithrombotic properties of materials while maintaining corrosion resistance under dynamic blood flow conditions.

[0125] 5. Study on the growth behavior of human umbilical vein endothelial cells (HUVECs) and smooth muscle cells (SMCs)

[0126] Using 316L stainless steel (SS), commonly used in commercially available cardiovascular stents, as a control, the specific experimental procedure is as follows: The back and sides of the freshly prepared sample were protected with silicone rubber. After pre-sterilization, endothelial cells (ECs) were introduced at a concentration of 5 × 10⁻⁶ mm. 4 Cells were directly seeded at a density of cells / mL on bare 316L stainless steel, bare magnesium, K / P coated, and K / PC coated surfaces and cultured at 37°C in a 5% CO2 incubator for 1 day and 3 days. Cell proliferation was then evaluated using the CCK-8 assay, and cells were stained with fluorescein diacetate (FDA). Cell morphology and distribution on the sample surface were observed using an inverted fluorescence microscope.

[0127] The experimental procedures for smooth muscle cells (SMCs) and mouse macrophages (RAW264.7) were consistent with those for human umbilical vein endothelial cells (ECs).

[0128] Human umbilical vein endothelial cells co-cultured with magnesium alloy extract showed significantly lower adhesion density than those in other groups. Figure 5(A), CCK-8 assay also showed reduced cell viability ( Figure 5 (D) This is because the magnesium hydroxide and other products generated by the corrosion of magnesium alloys increase the pH of the culture medium, inhibiting the growth of endothelial cells. In contrast, the endothelial cell viability of K / P coating, K / PC coating and control group showed no significant difference, indicating that the cell compatibility of the material was improved after coating modification.

[0129] To further investigate cell-material interactions, human umbilical vein endothelial cells were directly seeded onto the sample surface, with medical-grade 316L stainless steel (SS) used as a control. Figure 5 (B) After 1 day of culture, there was no significant difference in cell proliferation among the groups ( Figure 5 The low degradation rate of magnesium alloy in the early stages (E) is due to the relatively low degradation rate of magnesium alloy in this early stage; endothelial cells can adhere uniformly to the surface of all samples. However, after 3 days of culture, the degradation of bare magnesium leads to the accumulation of corrosion products and the release of hydrogen, which significantly inhibits the proliferation and adhesion of endothelial cells; the endothelial cells on the bare magnesium surface exhibit a round morphology rather than a typical spindle shape, indicating poor cell spreadability. In contrast, the adhesion and proliferation of endothelial cells on the K / P coating and K / PC coating surfaces are significantly improved, with the K / PC coating performing better than the K / P coating, indicating that collagen functionalization modification can further enhance the responsiveness of endothelial cells. Although the 316L stainless steel group showed the highest level of endothelial cell proliferation due to its surface stability and inertness, the collagen functionalized coating (K / PC) showed comparable cell compatibility while also possessing biodegradability and biofunctionality. These results indicate that the composite coating can not only inhibit the adverse effects of magnesium alloy degradation but also promote endothelial cell adhesion and proliferation, thereby accelerating the process of vascular reendothelialization.

[0130] Excessive proliferation of smooth muscle cells is a major cause of in-stent restenosis; therefore, evaluating the effect of modified surfaces on smooth muscle cells is crucial for assessing their anti-restenosis potential. Smooth muscle cells were seeded onto the sample surface using the same method as for endothelial cells, with medical-grade 316L stainless steel (SS) serving as a control.

[0131] like Figure 5 As shown in C and F, the 316L stainless steel group exhibited the highest smooth muscle cell proliferation, with cell numbers approximately twice that of the other groups after 1 and 3 days of culture. This is attributed to the fact that its inert surface did not inhibit cell growth. Among the remaining groups, the bare magnesium surface showed the lowest smooth muscle cell density. This is because the rapid degradation of the magnesium alloy, local alkalization, and hydrogen release created a microenvironment unfavorable to cell survival. However, this inhibitory effect is non-specific and can also negatively impact endothelial cells.

[0132] Persistent infiltration of inflammatory cells at the implantation site and the release of large amounts of pro-inflammatory factors can trigger an excessive inflammatory response, which can hinder the reendothelialization process on the scaffold material surface and lead to serious postoperative complications. This study evaluated the inflammatory response of the scaffold material by investigating the in vitro culture of mouse macrophages (RAW264.7) on the material surface. Figure 5 As shown in G and H, after 1 day of culture, compared with the 316L stainless steel group, the number of macrophages adhering to the surface of the Mg and K / PC group samples was lower, and the macrophage morphology did not become activated. This indicates that the K / PC coating can inhibit macrophage polarization and reduce the expression of pro-inflammatory factors, thereby alleviating the inflammatory response.

[0133] In contrast, both K / P and K / PC coatings significantly reduced smooth muscle cell adhesion and proliferation while maintaining good compatibility with endothelial cells. The K / PC coating showed a particularly strong inhibitory effect on smooth muscle cell growth, indicating that the collagen-functionalized interface can better balance the relationship between vascular reendothelialization and excessive smooth muscle proliferation. These results suggest that this composite coating not only alleviates the degradation problem of magnesium alloys but also regulates vascular cell behavior, which is beneficial in reducing the risk of in-stent restenosis.

[0134] 6. In vivo subcutaneous implantation experiment

[0135] The in vivo degradation performance and immune response of the samples were initially evaluated through subcutaneous implantation experiments. All procedures were performed in accordance with the "Ethical Guidelines for Laboratory Animals of Sichuan University". Healthy SD rats weighing approximately 200 g were selected. Naked magnesium, K, K / P, and K / PC sheets (8 mm in diameter and 1 mm in thickness) were wrapped with epoxy resin on the dorsal and lateral sides and implanted subcutaneously on the back of the rats under anesthesia. Four weeks after implantation, the implanted samples and surrounding tissues were collected, peeled off, rinsed with PBS, and fixed with 2.5% paraformaldehyde. The fixed tissues were stained with hematoxylin and eosin (H&E) for histological analysis, and the average thickness of the fibrous capsule was calculated. Immunohistochemical labeling was performed on macrophage surface marker CD68, anti-inflammatory cytokine interleukin-10 (IL-10), and pro-inflammatory factor tumor necrosis factor-α (TNF-α). The fluorescence intensity was quantitatively analyzed using ImageJ software to evaluate the degree of inflammatory response after material implantation.

[0136] To further evaluate in vivo inflammatory behavior, samples were subcutaneously implanted into SD rats. Figure 6 (A) The sides and back of the sample were sealed with epoxy resin to control the exposed area. Four weeks after implantation, the implant was removed and histologically analyzed, including hematoxylin-eosin (H&E) staining. Figure 6 (B) shows that a fibrous capsule forms around the implant, and the thickness of the fibrous capsule is often used as an indicator of the degree of inflammation. Quantitative analysis results ( Figure 6The results (C) show that the magnesium alloy group had the thickest fiber capsule (152.5±11.4 μm), followed by the K coating group (115.3±17.6 μm), the K / P coating group (83.7±10.3 μm), and the K / PC coating group (67.5±4.5 μm), indicating that the inflammatory response induced by the material gradually decreased after surface modification.

[0137] Further evaluation of inflammatory markers, including CD68 (macrophage marker), interleukin-10 (IL-10, an anti-inflammatory factor), and tumor necrosis factor-α (TNF-α, a pro-inflammatory factor), was performed using immunofluorescence staining. Figure 6 As shown in D and F, compared with bare magnesium, the CD68 fluorescence intensity of the K / P coating group and the K / PC coating group was significantly reduced, indicating a decrease in macrophage infiltration; in addition, the K / PC coating group had higher IL-10 expression and lower TNF-α expression. Figure 6 The presence of G and H indicates the formation of an anti-inflammatory microenvironment around the material. In summary, the results demonstrate that the collagen-functionalized coating can effectively regulate the host's immune response by reducing macrophage aggregation and promoting anti-inflammatory signal transduction, thereby alleviating post-implantation inflammatory responses.

[0138] In general, bare magnesium degrades rapidly in physiological environments, releasing large amounts of magnesium ions, which triggers local alkalization and hydrogen release. These effects adversely impact cell viability and trigger inflammatory responses. In contrast, a dense K / P anti-corrosion coating acts as an effective barrier, reducing direct contact between the magnesium alloy substrate and body fluids, thereby inhibiting oxidation reactions and reducing the release of corrosion products and hydrogen.

[0139] Furthermore, recombinant human type III collagen (rhCOL III), a key component of the extracellular matrix, can provide bioactive signals to regulate cellular behavior. Specifically, the collagen-functionalized interface can promote macrophage polarization to the anti-inflammatory M2 phenotype while inhibiting the activation of the pro-inflammatory M1 phenotype, accompanied by a decrease in the expression of pro-inflammatory cytokines such as tumor necrosis factor-α and an increase in the secretion of anti-inflammatory factors such as interleukin-10. In summary, the synergistic effect of antiseptic action and bioactive regulation enables this coating to reduce inflammatory responses and create a favorable microenvironment for vascular repair.

[0140] 7. In vivo vascular stent implantation experiment

[0141] This experiment strictly followed the "Sichuan University Guidelines for Experimental Animal Ethics." Healthy New Zealand white rabbits weighing approximately 2.5 kg were selected. After complete anesthesia, a metal balloon was attached to a metal guidewire, and the prepared bare magnesium, K / P, and K / PC stents were implanted into the abdominal aorta of the rabbits. Four weeks after implantation, the artery containing the stent was removed, cleaned to remove residual blood, and fixed with 2.5% formaldehyde solution. Five days later, three-dimensional tomographic imaging of the stent was performed using PerkinElmer's Quantum GX II small animal in vivo three-dimensional tomographic scanning system (Micro-CT). The inner surface of the stent was observed using a FEI Quanta 450 scanning electron microscope (SEM). The cross-section of the blood vessel was stained with hematoxylin and eosin (H&E) for histological analysis. In addition, the inner wall of the blood vessel was immunofluorescently stained with anti-CD31 antibody (ab9498, Abcam, UK) and anti-endothelial nitric oxide synthase (eNOS) antibody (ab5589, Abcam, UK) to observe the growth of endothelial cells on the stent surface. ImageJ software was used to calculate relevant statistical data.

[0142] Bare magnesium, K / P, and K / PC stents were implanted into rabbits, and the vascular segment containing the stent was removed 4 weeks later. Figure 7 E and F). Micro-computed tomography (Micro-CT) reconstruction and histological analysis ( Figure 7 The results (G and H) show that the bare magnesium stent undergoes partial degradation, while the K / P stent and K / PC stent maintain their structural integrity and provide effective radial support for blood vessels, which is attributed to the improved corrosion resistance of the coating.

[0143] Hematoxylin and eosin (H&E) staining further revealed that the bare magnesium scaffold was only partially covered by a thin layer of newly formed endometrium, while both the K / P and K / PC scaffolds were completely covered by continuous, mature newly formed endometrium; SEM observation results ( Figure 7 In line with this, the magnesium alloy scaffold structure was incomplete and degraded, while the coated scaffold maintained its complete morphology and was covered with a uniform inner skin layer.

[0144] The degree of endothelialization of the scaffold was evaluated by analyzing the expression of CD31 and endothelial nitric oxide synthase (eNOS) using immunofluorescence staining. CD31 is a marker of endothelial cell junction and maturation, while eNOS is crucial for nitric oxide production and the maintenance of vascular function. Figure 7As shown in J and K, the expression level of CD31 in the magnesium alloy group was low (13.75±1.78%), while the endothelial coverage of the K / P group (41.30±5.40%) and the K / PC group (36.81±3.20%) was significantly improved. Notably, the eNOS expression level in the K / PC group was the highest (84.77±5.60%), which was much higher than that in the magnesium alloy group (24.68±6.00%) and the K / P group (37.46±5.85%), indicating that the function of endothelial cells was improved.

[0145] Overall, the rapid degradation of bare magnesium leads to structural instability and the release of corrosion products such as magnesium hydroxide and hydrogen, which are detrimental to vascular repair. In contrast, the K / P coating can effectively stabilize the magnesium alloy substrate in vivo, while collagen functionalization further promotes vascular endothelialization and functional recovery. These synergistic effects enhance the performance of the stent, maintaining structural integrity while accelerating vascular regeneration.

[0146] Finally, it should be noted that the above specific embodiments are only used to illustrate the technical solutions of the present invention and not to limit it. Although the present invention has been described in detail with reference to examples, those skilled in the art should understand that modifications or equivalent substitutions can be made to the technical solutions of the present invention without departing from the spirit and scope of the technical solutions of the present invention, and all such modifications and substitutions should be covered within the scope of the claims of the present invention.

Claims

1. A method for preparing a collagen anti-corrosion coating with tissue repair properties, characterized in that, Includes the following steps: (1) The magnesium alloy substrate material is pretreated to obtain a magnesium alloy with surface hydroxylation; (2) The surface-hydroxylated magnesium alloy is immersed in solution I and dried. This is one cycle. After three cycles, it is dried under vacuum and high temperature to form an anti-corrosion coating. The solute in solution I includes copolymer P and copolymer K with a molar ratio of 1:1 to 1:

2. Copolymer P is obtained by polymerization reaction of siloxane, phenylboronic acid and reducing agent. Copolymer K is obtained by polymerization reaction of siloxane coupling agent, long-chain alkyl and initiator. (3) Immerse the magnesium alloy with anti-corrosion coating obtained in step (2) into solution II and dry it. This is one cycle. Repeat three cycles to obtain the collagen anti-corrosion coating. The solutes in solution II are excipients and collagen.

2. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In step (1), the hydroxylation process involves immersing the magnesium alloy substrate in a 0.5-2 mol / L sodium hydroxide ethanol solution at 50-70℃ for 0.5-2 h, or reacting the magnesium alloy substrate in a plasma cleaner for 3-6 min.

3. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In step (2), the reaction temperature for preparing copolymer P is 15~25℃ and the reaction time is 2-4 days; the reaction temperature for preparing copolymer K is 60~80℃ and the reaction time is 10-12 h.

4. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In step (2), the molar ratio of siloxane to phenylboronic acid is 1:2~5; wherein, the siloxane is at least one of bis(3-aminopropyl)-terminated polydimethylsiloxane and 3-mercaptopropyltriethoxysilane; the phenylboronic acid is at least one of 3-aminophenylboronic acid and 2-formylphenylboronic acid; and the reducing agent is at least one of sodium cyanoborohydride and sodium borohydride.

5. The method for preparing a collagen anti-corrosion coating with tissue repair properties according to claim 1 or 4, characterized in that, The silane coupling agent, long-chain alkyl group and initiator used in step (2) are all liquid phases with a volume ratio of 2~3:18~30:1; The silane coupling agent is at least one of 3-(trimethoxysilyl)propyl methacrylate, methacryloyloxypropyltriethoxysilane, and acryloyloxypropyltrimethoxysilane; the long-chain alkyl group is at least one of laurate methacrylate, tridecyl methacrylate, and isodecanyl methacrylate; and the initiator is at least one of dimethyl azobisisobutyrate, azobisisoheptanenitrile, and azobisisobutyronitrile.

6. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In step (2), the magnesium alloy with surface hydroxylation is immersed in solution I for 5-25 min, and the drying conditions after immersion are 70-100℃ under vacuum for 2-24 h.

7. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In solution II, the mass ratio of collagen to excipient is 0.5 to 1:1, and the final concentration of collagen is 1 to 5 mg / mL; the excipient is at least one of polylactic acid, polycaprolactone, polylactic-glycolic acid copolymer, silk fibroin, and poly(benzyl-L-glutamic acid).

8. The method for preparing the collagen anti-corrosion coating with tissue repair properties according to claim 1, characterized in that, In step (3), the magnesium alloy with anti-corrosion coating is immersed in solution II for 10-120 s, and the drying conditions are 30-40℃ for 15-60 min.

9. A collagen anti-corrosion coating with tissue repair properties, characterized in that, It is prepared by any one of the methods of claims 1 to 8.

10. The use of the collagen anti-corrosion coating with tissue repair properties as described in claim 9 in the preparation of medical device materials.

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