Preparation method and application of anticoagulant coating

Abstract

The invention provides an anticoagulant coating for medical instruments, which can be used in the field of regenerative medical instruments. The prepared anticoagulant coating has short-term and long-term anticoagulant performance, and in addition, the thrombin responsive nanoparticles wrapped in the hydrogel coating can selectively release a factor Xa (FXa) inhibitor apixaban according to needs, so that thrombosis is inhibited in the middle stage of implantation. In-vitro and in-vivo experimental evidences show that after the material is implanted for a long time, an organized endodermis can be seen on the surface of the material, and the material has a special antithrombotic characteristic. The dynamic adjusting system has a multi-layer protection mechanism, and a new thought can be provided for designing the surface of a contact blood material adapting to environmental conditions.

Description

Technical Field

[0001] The present invention belongs to the field of biomedical materials, and specifically relates to a preparation method and application of an anticoagulant coating for use in the field of regenerative medical devices. Technical Background The application of regenerative medicine materials in the medical field is making rapid progress. In particular, functional regenerative medicine materials, along with cutting-edge technologies such as stem cell regenerative medicine, bioregenerative materials, and organ bioprinting, have become research hotspots and garnered widespread attention. As a key branch of regenerative medicine, regenerative medicine materials are increasingly being used in orthopedics, trauma repair, and other fields. These materials, with their excellent physical and chemical properties and exceptional biosafety, demonstrate tremendous potential in inducing the repair and regeneration of human tissues. With the advancement of medical implant materials, there is a shift from traditional non-absorbable materials to new, biodegradable materials that actively induce tissue regeneration. Decellularized extracellular matrix (dECM), derived from organs or tissues, is a promising natural biomaterial well-suited for in vitro and in vivo regenerative medicine. dECM-based bioscaffolds have demonstrated remarkable efficacy in promoting remodeling in tissues such as skin, nerves, heart valves, and blood vessels. However, untreated dECM struggles to overcome immunogenicity and fails to mimic the in vivo microenvironment, limiting its repair capacity. In particular, manufacturing organ tissues such as heart valves poses a huge challenge due to their complex and anisotropic geometries, as well as stringent mechanical and biological requirements in a dynamic microenvironment.

[0002] At the current technological level, commercial bioprosthetic valves and glutaraldehyde-crosslinked bioprosthetic valves have successively provided technical support for regenerative medicine and are widely used in transcatheter valve interventional therapy; however, the durability of commercial bioprosthetic valves composed of porcine or bovine pericardium fixed with glutaraldehyde is limited by structural valve degeneration (SVD), and the residual aldehyde groups on glutaraldehyde-crosslinked bioprosthetic valves can lead to bioprosthetic valve failure. Calcification, thrombosis, cytotoxicity and inflammatory response are the main causes of the degradation of glutaraldehyde-crosslinked bioprosthetic valve materials. Thrombosis remains a major challenge facing bioprosthetic valves.

[0003] Currently, coating strategies are the main approach to enhance the antithrombotic properties of implantable materials. These strategies include drug coatings, such as heparin coatings, which can effectively prevent early thrombosis. However, their long-term efficacy is still affected by issues such as short half-life, limited drug loading, and the risk of bleeding associated with high-dose administration. In addition, hydrophilic / hydrophobic coatings can enhance the anticoagulant effect by preventing the interaction between coagulant substances and implanted materials. Nevertheless, barrier-based antithrombotic approaches are short-lived passive defense mechanisms. As the implantation time increases, the adsorption of proteins on the implant surface will blur the functional coating, rendering it ineffective. In addition, these coatings cannot actively regulate the locally formed microthrombi and may ultimately fail to achieve the intended function.

[0004] An ideal anticoagulant coating should not only mitigate acute thrombosis during the initial implantation phase, actively regulate local microthrombi during use, but also promote endothelialization to ensure sustained anticoagulation. This invention, based on the self-feedback regulation system of a functional hydrogel coating, has developed a dual-regulatory strategy, resulting in a functional bioprosthetic valve material that combines short-term and long-term anticoagulant properties with enhanced endothelialization. Thrombin-responsive nanoparticles are encapsulated in the hydrogel, enabling the selective release of the factor Xa (FXa) inhibitor apixaban on demand, thereby inhibiting thrombosis during the mid-implantation phase. Summary of the Invention

[0005] In a first aspect, the present invention provides an anticoagulant coating for medical devices, characterized in that the coating comprises thrombin-responsive nanoparticles, cleavable polypeptides, a hydrogel, and an active pharmaceutical ingredient.

[0006] Furthermore, the thrombin-responsive nanoparticles are composed of inorganic salt compounds, weak acid compounds, cleavable polypeptides, and active pharmaceutical ingredients.

[0007] Furthermore, the inorganic salt compound is selected from one or more of calcium carbonate, sodium periodate, sodium iodate, potassium iodate, and ammonium periodate; the weak acid compound is selected from one or more of sodium alginate, dextran, chondroitin sulfate, keratan sulfate, and oxidized hyaluronic acid; and the active pharmaceutical ingredient is selected from one or more of hirudin, heparin, rivaroxaban, and apixaban.

[0008] Furthermore, the cleavable polypeptide is selected from NH2-Gly-(D)CHA-Ala-Arg-Ser-Trp-Gly-CONH2.

[0009] Furthermore, the hydrogel is selected from one or more of collagen, gelatin, hyaluronic acid, chitosan, fibrin polyacrylic acid, and polyvinyl alcohol.

[0010] Furthermore, the hydrogel is preferably gelatin.

[0011] In a second aspect, the present invention provides a method for preparing an anticoagulant coating for medical devices, characterized in that the preparation method comprises the following steps: S1. Preparation of OX-HA: An inorganic salt compound powder is added to a weak acid compound solution. After stirring in the dark, an alcohol compound is added to interrupt the reaction. OX-HA is obtained by dialysis, freezing, and drying. S2. Preparation of Thrombin-responsive Nanoparticles (Thr NPs): Dissolve the OX-HA and active pharmaceutical ingredient obtained in Step 1, then add the responsive polypeptide chain dropwise to the solution and dialyze overnight to obtain Thr NPs. S3. The medical device to be coated is subjected to simple treatment and then soaked. The side chains, thrombin-responsive nanoparticles, and hydrogel are added to the soaking liquid to form an anticoagulant coating on the surface of the medical device.

[0012] Furthermore, the inorganic salt compound is selected from one or more of calcium carbonate, sodium periodate, sodium iodate, potassium iodate, and ammonium periodate; the weak acid compound is selected from one or more of sodium alginate, dextran, chondroitin sulfate, keratan sulfate, and oxidized hyaluronic acid; and the alcohol compound is selected from one or more of ethanol, isopropanol, and ethylene glycol.

[0013] Furthermore, in step 1, the ratio of the inorganic salt compound, the weak acid compound, and the alcohol compound is 0.6:1:1.

[0014] Furthermore, in step 2, the active pharmaceutical ingredient is selected from one or more of hirudin, heparin, rivaroxaban, and apixaban.

[0015] Furthermore, in step 2, the polypeptide is selected from NH2-Gly-(D)CHA-Ala-Arg-Ser-Trp-Gly-CONH2.

[0016] Furthermore, in step 3, the medical device is selected from a bioartificial heart valve; the side chain is selected from branched polyethyleneimine, and the thrombin-responsive nanoparticles are the product prepared in step 2 of the method for preparing an anticoagulant coating for medical devices described in the second aspect of the present invention.

[0017] Furthermore, in step 3, the hydrogel is selected from one or more of collagen, gelatin, hyaluronic acid, chitosan, fibrin polyacrylic acid, and polyvinyl alcohol.

[0018] Furthermore, the hydrogel is preferably gelatin.

[0019] In a third aspect, the present invention provides a use of the anticoagulant coating as described in the first aspect in the preparation of a regenerative medical device.

[0020] In a fourth aspect, the present invention provides a use of thrombin-responsive nanoparticles in the preparation of auxiliary medical products, characterized in that the preparation method of the thrombin-responsive nanoparticles is as follows: S1. Preparation of OX-HA: An inorganic salt compound powder is added to a weak acid compound solution. After stirring in the dark, an alcohol compound is added to interrupt the reaction. OX-HA is obtained by dialysis, freezing, and drying. S2. Preparation of Thrombin-responsive Nanoparticles (Thr NPs): Dissolve the OX-HA and active pharmaceutical ingredient obtained in step 1, then add the responsive polypeptide chain dropwise to the solution and dialyze overnight to obtain Thr NPs.

[0021] Furthermore, in step 1, the ratio of the inorganic salt compound, the weak acid compound, and the alcohol compound is 0.6:1:1.

[0022] Furthermore, in step 2, the active pharmaceutical ingredient is selected from one or more of hirudin, heparin, rivaroxaban, and apixaban.

[0023] Furthermore, in step 2, the polypeptide is selected from NH2-Gly-(D)CHA-Ala-Arg-Ser-Trp-Gly-CONH2.

[0024] Furthermore, the medical product is selected from one or more of a molecular probe containing thrombin-responsive nanoparticles, a reagent test kit, a tumor drug delivery system, and other auxiliary nanomaterials.

[0025] Beneficial effects of the present invention: In this study, a dual-regulatory strategy based on a functional hydrogel coating and a self-feedback regulation system was developed to develop a functional bioprosthetic valve material with both short-term and long-term anticoagulant properties and enhanced endothelialization. The functional hydrogel shielding layer is designed to provide short-term anticoagulation (early passive protection). Furthermore, these thrombin-responsive nanoparticles encapsulated in the hydrogel coating selectively release the factor Xa (FXa) inhibitor apixaban on demand, thereby inhibiting thrombosis during the mid-implantation period (intermediate active feedback regulation). The cell-affinity interface within the hydrogel attracts endothelial cells to form an endothelial layer, establishing long-term anticoagulation through a synergistic, full-cycle antithrombotic biomimetic control strategy. This integrated approach simultaneously addresses challenges such as early implantation thrombosis, mid-term antithrombotic interface regulation, and delayed late-stage endothelialization, significantly enhancing the multifunctional biological performance of the bioprosthetic valve. In vitro and in vivo experimental evidence demonstrates that after long-term implantation, an organized endothelial layer is visible on the material surface, demonstrating exceptional antithrombotic properties. This dynamic regulation system, with its multi-layered protective mechanism, may provide new insights into the design of environmentally adaptable blood-contacting material surfaces. BRIEF DESCRIPTION OF THE DRAWINGS

[0026] Figure 1 The manufacturing process and biological response of bioprosthetic valves.

[0027] Figure 2 (A) Dynamic light scattering spectra and PDI values ​​of Thr NPs before and after treatment with thrombin solution; (B) Accumulation of API released from thrombin-sensitive nanoparticles in PBS containing different thrombin concentrations; (C) Transmission electron microscopy (TEM) images; (D) Scanning electron microscopy (SEM) images of the surfaces of G-PP, HMM-PP, HPO-PP, and HPT-PP; (F) Evaluation of the hydrophilicity of different PPs; (E) Evaluation of the thermal stability of different PPs; (G) Ultimate tensile strength of different PPs; (H) Ultimate elongation of different PPs.

[0028] Figure 3 Evaluation of cytocompatibility [L929 (A) and HUVECs (B) live / dead cells were incubated with PP extracts for 1 and 2 days and then labeled with FDA and PI, respectively; (C) FDA fluorescence staining of HUVECs on different samples after incubation at 37°C for 1, 2, and 3 days; (D) Confocal images of cell junctions of HUVECs cultured on samples stained with phalloidin (red) and DAPI (blue) for 1, 2, and 3 days; (E) Cell viability of HUVECs cultured on different PPs for 1, 2, and 3 days; (F) Fluorescence quantitative analysis of HUVECs cultured on different PPs for 1, 2, and 3 days; ELISA detection of TNF-α (G) and IL-10 (H) levels of RAW 264.7 after co-incubation with PPs].

[0029] Figure 4 SEM images (A) and lactate dehydrogenase images (B) of quantitative results of platelet adhesion on different PPs, with platelets appearing pink; (C) Microscopic images of thrombus adhesion to the PP surface after incubation with whole blood, with thrombus appearing red; (D) Visual images and coagulation kinetics at 0, 5, 15, and 35 minutes (E), including control, G-PP, HMM-PP, HPO-PP, and HPT-PP; (F) Representative images of hemolysis assessment; (G) Images of the results of the second whole blood recalcification and coagulation test on PP specimens (the first is on the left, the second is on the right); (H) Quantitative results of the second whole blood recalcification and coagulation test on PP specimens.

[0030] Figure 5(A) Schematic diagram of the in vitro arteriovenous shunt model. (B) Visualization of thrombus on the sample surface before and after 1 hour of circulation. (C) SEM image of thrombus adhered to the surface after 1 hour of circulation; the thrombus appears red. (D) Surface and cross-sectional morphology of HPT-PP before and after immersion in PBS for 1 week. Visual image after 1 hour of in vitro arteriovenous shunt (E) and SEM images of HPT-PP samples before and after immersion in PBS for 1 week (F).

[0031] Figure 6 (A) Representative H&E staining of PPs 7 days after subcutaneous implantation, as well as representative immunofluorescence images of macrophages stained for CD68 (all macrophages, red), CD206 (M2, pink), and CD86 (M1, green) in different groups. Scale bar: 200 μm. (B) Alizarin red-stained sections of calcified G-PP, HMM-PP, HPO-PP, and HPT-PP at 30 and 60 days after implantation. (C) Quantification of corresponding fluorescence intensity for CD68 expression; (D) Quantification of corresponding fluorescence intensity for CD86 and CD206 expression. (E) Quantitative analysis of calcification in different PPs 30 and 60 days after implantation using ICP-OES.

[0032] Figure 7 (A) Implantation of PP samples in the rabbit jugular vein or carotid artery; (B) Thrombus and surface micromorphology of G-PP and HPT-PP before and 7 days after implantation; microscopic morphology of the surface of different PPs (C) 30 days and (D) 60 days after carotid artery implantation; immunofluorescence images of CD31 and eNOS expression on the surface of G-PP and HPT-PP (E) 1 month and (F) 2 months after implantation. DETAILED DESCRIPTION

[0033] The G-PP described in the present invention is pig pericardium cross-linked with glutaraldehyde, the HMM-PP is pig pericardium cross-linked with HMM, and HMM is hexamethylene glycol melamine.

[0034] All experimental animals and procedures were approved by the Medical Ethics Committee of the Sichuan Provincial Laboratory Animal Management Committee and were performed in accordance with the Guidelines for the Care and Use of Laboratory Animals of Sichuan University (No. KS2020394). Sprague-Dawley rats and New Zealand white rabbits were provided by Chengdu Dashuo Laboratory Animal Co., Ltd., China.

[0035] Example 1 Synthesis of OX-HA and Thrombin Responsive Nanoparticles (Thr NPs) Preparation of OX-HA: Sodium periodate (NaIO4) powder (0.6 g) was added to 100 mL of HA solution (10 mg / mL). After stirring in the dark for 24 h, ethylene glycol (1 mL) was added to terminate the reaction. OX-HA was obtained by dialysis against deionized water (Mw = 3500 Da) and freeze-drying for 1 week.

[0036] Preparation of thrombin-responsive nanoparticles (Thr NPs): OX-HA (20 mg) and apixaban (Api) (4 mg) were dissolved in DMSO (4 mL). The responsive polypeptide chain (5 mg / mL, 4 mL) was then added dropwise to the stirred solution at 25°C. After overnight dialysis against DI water (Mw = 1000 Da) at room temperature for 24 h to obtain Thr NPs.

[0037] Results: Self-feedback nanoparticles were prepared by cross-linking the amino group of the cleavable peptide NH2-Gly-(D)CHA-Ala-Arg-Ser-Trp-Gly-CONH2 with the aldehyde group of OX-HA. The synthesized nanoparticles were uniform in size, with an average diameter of (187.9±0.17) nm, as measured by a nanoparticle size analyzer. The thrombin reaction of the nanoparticles (Thr NPs) was studied in PBS solutions containing 0, 2, 5, 10, and 20 U / ml thrombin. Figure 2 As shown in A, under the stimulation of 20 U / mL thrombin, the particle size distribution changed significantly compared with the pre-response state, indicating that thrombin triggered the disintegration and reorganization of the nanoparticles. The uneven distribution of Thr NPs in PBS after the response can also be observed from the PDI value. TEM image ( Figure 2 C) shows that the structure of Thr NPs was destroyed after thrombin treatment. Figure 2 Figure B shows the release curves of apixaban from Thr NPs treated with different thrombin concentrations (0 U / mL, 2 U / mL, 5 U / mL, and 10 U / mL). Both the amount and rate of apixaban released increased with increasing thrombin concentration. The self-feedback nanoparticles react rapidly with thrombin, releasing the anticoagulant upon contact, blocking approximately 80% of coagulation within 8 hours (at 10 U / mL).

[0038] Example 2 Preparation and Characterization of Bioprosthetic Heart Valve Bioprosthetic heart valve preparation: The bioprosthetic valve was first treated with HMM and then modified. Prior to modification, the leaflets were placed in PBS and shaken for 72 hours to remove excess HMM and DMSO. The leaflets were then immersed in 0.1 M EDC·HCl and 0.1 M NHS for 2 hours to activate the carboxyl groups. 100 mg / mL branched polyethyleneimine (PEI) was then added at 37°C to expand the amine groups (designated H-PEI-PP). After 48 hours, the bioprosthetic valve was rinsed with DI water for further modification. The H-PEI-PP was then immersed in a mixture of 0.5 mg / mL Thr NPs solution and 5 mg / mL gelatin at 37°C for 48 hours to obtain the modified bioprosthetic valve, designated HPT-PP. Additionally, a control (designated HPO-PP) was prepared using OX-HA instead of the NPs solution.

[0039] Preparation of anticoagulant coating: Thr NPs act as a bridge, and the excess aldehyde groups on the particles can react with the amine groups on the gelatin segments and the amine groups on the decellularized matrix material to form a Schiff base. Since gelatin can form a hydrogel after dissolving in water, the gelatin hydrogel coating is connected to the decellularized matrix through this bridge intermediate medium.

[0040] Examination Methods: Scanning electron microscopy (SEM, Hitachi S-3400N, Japan) was used to observe the micromorphology of bioprostheses before and after modification. The effects of modification on the mechanical properties and stability of the bioprostheses were evaluated by mechanical property and thermal shrinkage temperature testing. The mechanical properties of bioprostheses (n = 8, 40 × 10 mm²) were tested under uniaxial tension using a Bio-Tester 5000, Cell scale. Testing was performed at a speed of 12.5 mm / min until fracture. The thermal shrinkage temperature of the bioprostheses under a nitrogen atmosphere was measured using a differential scanning calorimeter (DSC 204 F1, NETZSCH, Germany). The hydrophilicity of the bioprostheses at room temperature was determined using a contact angle meter (Harke - spca1, Harke Laboratory Instrument Factory, Beijing, China).

[0041] HPT-PP (1 cm × 1 cm) was placed in PBS containing thrombin (0 and 10 U / mL) and soaked for 2 h at regular intervals. The amount of released Api was calculated using UV measurement ( Figure 2 B).

[0042] Characterization of HPT-PP: The thermal denaturation temperature can reflect the stability of collagen to a certain extent. The thermal denaturation temperature of HMM-PP and HPT-PP is slightly higher than that of G-PP ( Figure 2F), further indicating that HMM cross-linked collagen has better stability, and the functional hydrogel layer formed is more conducive to maintaining the stability of biological valves. In addition to thermal denaturation temperature, ultimate tensile strength is also an important parameter for determining the maximum load that biological valve materials can withstand. The mechanical properties of G-PP, HMM-PP, HPO-PP, and HPT-PP samples were tested by uniaxial tensile tests. The ultimate tensile strength of HMM-PP and these modified PP materials ( Figure 2 G) and elongation ( Figure 2 H) There is no significant difference compared with G-PP, which indicates that the modification of hydrogel and nanoparticles does not affect the basic mechanical properties of the original material, enabling it to withstand the large amount of mechanical stress imposed by blood flow. The surface morphology of these PP samples was examined using scanning electron microscopy (SEM), as shown in Figure 2. Figure 2 As shown in Figure D. Compared with G-PP, the surface of HMM-PP is smoother. A hydrogel layer is visible on the fiber bundles of HPO-PP and HPT-PP. The hydrophilicity of these PPs is also characterized by water contact angle. Figure 2 As shown in Figure E, the hydrophilicity of these immobilized PPs can be improved by the introduction of a hydrogel coating layer and functional nanoparticles, and the water contact angle of HPT-PP (45.9 ± 1.28°) is significantly lower than that of G-PP (87.8 ± 1.73°).

[0043] Example 3 In vitro cell compatibility verification and evaluation 1. In vitro cytocompatibility evaluation The cytotoxicity of G-PP, HMM-PP, HPO-PP, and HPT-PP against mouse fibroblasts (L929) and human umbilical vein endothelial cells (HUVECs) was analyzed according to ISO 10993-5.

[0044] Methods: The modified HPT-PP bioprosthesis was placed in a 24-well plate and sterilized. The leaflets were incubated with DMEM for 72 hours to prepare an extract. The cells were cultured for 24 hours at a density of 5,000 cells per well. The culture medium was then replaced with 150 μL of the extract, and the cells were incubated for 1-2 days. Cell viability was assessed using the CCK-8 assay. L929 cells and HUVECs were stained with FDA / PI, and images were observed using a fluorescence microscope (Nexcope NIB900, USA).

[0045] 2. In vitro cell proliferation and morphology detection 0.5 ml contains approximately 2 × 10 4A suspension of HUVECs containing 100 cells was inoculated onto the bioprosthetic valve. After 1, 2, and 3 days of culture, cell proliferation in each group was assessed using CCK-8 assay. Adherent cells were fluorescently stained with FDA, rhodamine-labeled phalloidin (TRITC-Phalloidin), and 4',6-diamidino-2-phenylindole (DAPI). Adherence of cells to the bioprosthetic valve was observed using fluorescence microscopy and laser confocal microscopy (Zeiss LSM 800, DE).

[0046] 3. Quantification of inflammatory potential After sterile preparation of the bioprosthetic valves in each group, RAW 264.7 macrophages were seeded onto the surface of the valves and plated in 48-well plates (25,000 cells / well). After culturing for 1 and 2 days, the supernatants were centrifuged and the release of inflammatory cytokines IL-10 and TNF-α was assayed using ELISA kits.

[0047] Results: The interaction, adhesion, and proliferation abilities of L929 and HUVECs were tested by co-culture extracts to evaluate the in vitro biocompatibility of L929 and HUVECs. The effects of different PP extracts on the viability of L929 cells and HUVECs were observed, and fluorescence imaging was performed. FDA / PI staining was performed on cells co-cultured with extracts and L929 or HUVECs ( Figure 3 A and B) show that the incidence of dead cells (red) on the G-PP surface was higher, while the incidence of live cells (green) on the surfaces of the other three sample groups was comparable to that of the control group, indicating that these modified HMM-PP samples have good biocompatibility.

[0048] The natural endothelium of the cardiovascular system acts as an anti-thrombotic surface. HUVECs adhesion and recruitment is a key step in initiating endothelialization. Endothelial formation involves cell adhesion, migration, and proliferation. To determine the enhanced endothelialization potential of these modified PP materials, we seeded HUVECs onto the surface of these PP materials. Figure 3 As shown in C and 3D, after co-incubation for 1d, 2d or 3d, a small number of HUVECs were present on G-PP. In addition, cytoskeleton and nuclear staining showed that HUVECs on G-PP exhibited an abnormal state with a round and incomplete cytoskeleton, which may be due to the presence of toxic aldehyde groups on its surface. After HPT-PP was co-incubated with HUVECs for 1 d or 2 d, a large number of well-proliferated HUVECs were observed on the HPT-PP surface. Cytoskeleton and nuclear staining showed that HUVECs on HPT-PP were in a fully spread and healthy growth state. The fluorescence intensity of living cells on the surface of the material was semi-quantitatively analyzed using Image J, as shown in Figure 3D. Figure 3As shown in F, the fluorescence intensity of the HPT-PP group was significantly higher than that of the G-PP group. The number of live HUVECs in the sample was then quantified using the CCK-8 kit ( Figure 3 E) The absorbance of G-PP remained consistently low, at approximately 0.2 OD, while that of HPT-PP increased significantly over time, reaching approximately 0.8 OD after 3 days. This indicates that these modified HMM-PP materials favor HUVEC proliferation due to the functional surface coating. These results demonstrate that HPT-PP has low cytotoxicity and the potential to improve endothelialization, showing promising clinical applications for HPT-PP.

[0049] The PP samples were co-incubated with macrophages (Raw 264.7), and the concentrations of anti-inflammatory factors (IL-10) and pro-inflammatory factors (TNF-α) in the supernatant secreted by the cells were detected at specific time points to evaluate the inflammatory response of these PP materials. Figure 3 As shown in G and 3H, after 2 days of incubation, RAW 264.7 cells secreted less IL-10 on the G-PP surface (48.85 ± 5.97 pg / mL) than on the HPT-PP surface (69.28 ± 3.23 pg / mL). Conversely, TNF-α secretion on the G-PP surface (110.2 ± 17.46 pg / mL) was significantly higher than that on the HPT-PP surface (79.08 ± 7.30 pg / mL). These in vitro inflammatory assay results support the conclusion that HPT-PP is less likely to induce a more severe inflammatory response.

[0050] Example 4 In vitro blood compatibility evaluation In vitro hemocompatibility assessment was performed using porcine pericardium as a bare substrate using platelet-rich plasma (PRP), red blood cells (RBCs), or whole blood. Fresh rabbit blood was centrifuged at 1500 rpm for 10 minutes to obtain a PRP and RBC pellet. The bioprosthesis was incubated with PRP at 37°C for 1 hour. The samples were then gently washed to remove non-adherent platelets, fixed with 2.5% glutaraldehyde, and subjected to gradient dehydration for observation under a scanning electron microscope (SEM, Phenom Pro, Thermo Fisher Scientific). The number of platelets adhered to the samples was determined by quantifying lactate dehydrogenase (LDH) released from platelet lysates. The whole blood adhesion assay followed the same procedures as the platelet adhesion assay, except that PRP was replaced with CaCl2-activated whole blood, which was then incubated with the bioprosthesis for 15 minutes.

[0051] In addition, the hemocompatibility of the samples was further evaluated by a hemolysis test. Fresh rabbit blood was centrifuged at 1500 rpm for 10 minutes, and the lower precipitate was removed and incubated with the samples at 37°C. PBS and deionized water served as the negative (-) and positive (+) groups, respectively.

[0052] The dynamic coagulation time test was used to detect the coagulation reaction that occurred when the sample interacted with blood. The coagulation kinetics of blood on the surface was studied after 5, 15 and 35 minutes of interaction. The plate was sized and placed in a 12-well plate. Activated whole blood was prepared by reacting 5 mL of anticoagulated blood with 500 μL of 0.1 M CaCl2. 100 μL of activated blood was then added to the plate and incubated at room temperature for 5, 15, and 35 minutes. At the end of each time point, 2 mL of deionized water was added and incubated for 5 minutes. Red blood cells not trapped in the clot were dissolved with deionized water to release hemoglobin into the solution. 200 μL of solution from each sample was transferred to a 96-well plate, and the free hemoglobin concentration in the solution was measured at 540 nm using a microscope reader. An absorbance curve over time was obtained. Blank blood was added as a positive control.

[0053] The whole blood recalcification test evaluates the anticoagulant properties of the blood-contact material. Fresh rabbit blood neutralized with 0.1M anticoagulant was incubated with the material at 37°C for 15 minutes. The sample surface was gently rinsed with PBS, and surface thrombus formation was photographed. A Triton X-100 solution (0.5 wt%) was added to the plate and shaken at 37°C until the thrombus was completely dissolved. Thrombus formation was quantitatively characterized by measuring the absorbance of the solution at 405 nm.

[0054] In vitro arteriovenous shunt experiments were performed according to previous reports. The bioprosthesis was tightly attached to the inner wall of a PVC tube, blood was allowed to clot on the smooth surface, and then treated with heparin. New Zealand rabbits (weighing 2.5 to 3 kg) were anesthetized with sodium pentobarbital, and the left carotid artery and right jugular vein were carefully isolated. The assembled catheters were connected to the carotid artery and jugular vein, respectively, to form a pathway. Blood circulation was continued for 1 hour. After the experiment, samples from each group were rinsed, photographed, and fixed with 2.5% glutaraldehyde for dehydration and observation under a scanning electron microscope.

[0055] Example 5 In vitro anticoagulant performance Implantable cardiovascular devices require anticoagulant therapy because they can cause blood cell aggregation and activation, leading to thrombosis. Most foreign bodies that enter the human cardiovascular system are thrombogenic. Good blood compatibility is a key criterion for cardiovascular implant materials, such as heart valves and vascular stents. Thrombosis is considered a potential threat to early thickening and calcification of bioprosthetic valves. Therefore, the present invention evaluated the blood compatibility of HPT-PP from several aspects, including protein adhesion, hemolysis, platelet adhesion, and thrombosis.

[0056] Results: During the initial implantation phase, protein adsorption initiated a cascade of adverse reactions, including thrombosis. Platelet aggregation and activation were associated with increased thrombin generation. The number of platelet adhesions was determined by scanning electron microscopy (SEM) and lactate dehydrogenase (LDH); Figure 4 As shown in A, the G-PP membrane was extensively covered with a large number of irregularly shaped platelets, whereas the HPT-PP surface showed only a very small number of platelets. The number of adhered platelets was quantified using the LDH assay ( Figure 4 B) The trend of platelet adhesion is consistent with the results of SEM images.

[0057] Depend on Figure 4 As can be seen from the experimental results, the hemolysis rate of the PP samples remained significantly below the 5% threshold specified by medical standards, indicating that these materials are safe and reliable for biomedical applications. The anti-thrombotic properties of these PP materials were evaluated through whole blood adhesion tests, kinetic clotting time tests, and whole blood recalcification tests. Whole blood coagulation kinetics were studied by measuring the amount of free hemoglobin in the absorbance of different PP samples. Higher absorbance values ​​of hemoglobin solutions indicate a lower likelihood of clot formation on the sample surface. Figure 4 Figures D and 4E show qualitative and semiquantitative analyses of coagulation kinetics for different samples. The absorbance values ​​of G-PP, HMM-PP, and HPO-PP were very similar during the first 5 minutes, but HPT-PP exhibited a higher absorbance. At 15 minutes, while the level of free hemoglobin in the blood, which interacted with all PPs, decreased rapidly, HPT-PP exhibited delayed coagulation kinetics. At 35 minutes, the absorbance values ​​of HMM-PP and G-PP were lower than those of the control group, indicating that the coagulation they induced was very severe. In contrast, HPT-PP maintained a high absorbance, indicating its good anticoagulant properties.

[0058] The anticoagulant activity of the specimen can be evaluated by the recalcification whole blood coagulation test. Figure 4 The surfaces of these PP samples were examined using SEM after incubation with recalcified whole blood, as shown in Figures C, 4G, and 4H. Notably, fewer red blood cells (red) and platelets (orange) were observed on the HPT-PP surface compared to the other samples. Figure 4 G provides a more intuitive visualization of this observation. To comprehensively evaluate the long-term antithrombotic capabilities of the materials, the samples were subjected to a second recalcification experiment. During the second recalcification experiment, the absorbance values ​​of G-PP, HMM-PP, and HPO-PP showed a significant increase in thrombi, while the absorbance of HPT-PP remained at a lower level, highlighting its enhanced antithrombotic properties.

[0059] Example 6 Stability Test The functional integrity of implanted heart valves is closely related to their service life. After HPT-PP was soaked in PBS at 37°C for 7 days, the anti-thrombotic properties of the soaked samples were evaluated by scanning electron microscopy, in vitro arteriovenous shunt experiments, and whole blood recalcification experiments to determine the stability of HPT-PP blood compatibility. Figure 5 As shown in Figure 2, SEM images of HPT-PP showed no significant changes in morphology before and after immersion in PBS for one week. Furthermore, in vitro venous shunt experiments revealed that only a small number of red blood cells were observed on the HPT-PP sample, similar to that observed on HPT-PP immersed in PBS, demonstrating its superior long-term antithrombotic properties.

[0060] Example 7 Evaluation of the immunomodulatory ability of PPs and the anti-calcification ability of HPT-PP 1. Evaluation of the immunomodulatory ability of PPs The biocompatibility of PP samples was preliminarily investigated by subcutaneous implantation in rats. Acute inflammation usually occurs near implanted medical devices. Overall, all PP samples showed a certain immune response, and the following analysis was performed: Figure 6 Representative histological images (H&E staining) are shown in A. On day 7 after subcutaneous implantation, significant cell infiltration was observed in the G-PP group, while no significant cell infiltration was observed in the other three groups. CD68 is a macrophage-specific marker, while CD86 and CD206 are markers of M1 and M2 macrophages, respectively. CD68, CD86, and CD206 staining were used to semi-quantitatively evaluate inflammatory, pro-inflammatory, and anti-inflammatory responses. Qualitative ( Figure 6 A) and semiquantitative immunofluorescence analysis ( Figure 6 C) shows that HPT-PP showed the least positive cells around the sample, while G-PP showed the most obvious inflammatory response. This may be due to the large number of unreacted aldehyde groups on the surface of G-PP, which caused cell damage and triggered a severe immune response. Immunofluorescence detection of the expression of inflammatory-related factors CD206 and CD86. Figure 6 As shown in A, immunofluorescence analysis showed that M1 macrophages (CD86 positive) were mainly distributed around G-PP, while M2 macrophages (CD206 positive) were mainly distributed in the tissue around HPT-PP. Semi-quantitative analysis of fluorescence intensity obtained by Image J software ( Figure 6 D) Confirmed the above findings, indicating that HPT-PP has no tendency to induce a significant acute inflammatory response.

[0061] 2. Evaluation of HPT-PP anti-calcification ability Calcification is a key factor affecting the lifespan of commercial bioprosthetic valves, particularly G-PP. Residual aldehyde groups in glutaraldehyde damage blood cells and readily bind to intranuclear calcium ions. Calcification after subcutaneous implantation was assessed using Alizarin Red staining and inductively coupled plasma optical emission spectroscopy (ICP-OES). Figure 6 B shows that G-PP has obvious calcification, which is manifested as a large number of red spots on the slice. In contrast, the non-glutaraldehyde cross-linked PP group still maintained a complete fiber structure at 60 days, and no red calcification area was observed. Figure 6 As shown in Figure E, the calcium content of G-PP was significantly higher (30 days: 177.25 ± 6.84 μg / mg, 60 days: 4.71 ± 1.06 μg / mg). The calcium content of HMM-PP, HPO-PP, and HPT-PP was only (4.71 ± 1.06) μg / mg after 60 days of subcutaneous implantation, indicating that these non-glutaraldehyde cross-linked PPs have good anti-calcification properties.

[0062] Example 8 In vivo antithrombotic ability The rabbit jugular vein implantation model was used to observe its short-term antithrombotic ability. The flowchart of the rabbit model is shown in Figure 7 A. The material was cut into strips and implanted into the jugular vein of a rabbit. After 7 days, the samples were recovered, fixed, and dehydrated to examine the surface micromorphology. Figure 7 Figure B shows that thrombus formation on the G-PP sample was significantly more pronounced than that observed on the HPT-PP surface. SEM imaging further revealed that a large number of red blood cells (RBCs) bound to fibrin formed thrombi on the G-PP surface. In contrast, only a very small number of RBCs were observed on the HPT-PP surface, with no thrombi forming. These results demonstrate that HPT-PP materials have significant resistance to acute thrombosis in vivo.

[0063] Example 9 Reendothelialization of HPT-PP in vivo After heart valve implantation, it is necessary not only to ensure anti-thrombosis and anti-calcification, but also to achieve surface endothelialization to obtain long-term function. Two rabbit artery models, including a carotid artery implantation model and a covered stent model, were established to evaluate the long-term endothelialization ability of PPs. First, the PP material was cut into strips and implanted into the rabbit carotid artery to evaluate its in situ re-endothelialization ability. We obtained arterial tissue containing PP material 30 and 60 days after implantation for histological evaluation. Scanning electron microscope image of the implant surface ( Figure 7 (C and D) show that at 30 days, endothelial cells (ECs) on the G-PP surface did not grow in an orderly manner. After 60 days, the G-PP surface became disorganized and covered with a protein network. In contrast, ECs on the HPT-PP surface gradually grew in the direction of blood flow at 30 and 60 days after implantation.

[0064] Immunofluorescence was used to detect the expression of CD31 and endothelial nitric oxide synthase (eNOS) on the surface of G-PP and HPT-PP to evaluate the degree of re-endothelialization. CD31 is localized to the tight junctions between endothelial cells and is associated with the mature phenotype of endothelial cells. The expression of eNOS, a key function of endothelial cells, reflects the overall degree of endothelialization of different materials. High expression of CD31 and eNOS indicates the integrity and health of the endothelial layer. Figure 7 As shown in Figures E and 7F, 30 days after implantation, the G-PP surface showed only minimal, disorganized expression of CD31 and eNOS. While CD31 and eNOS expression levels were significantly higher in the HPT-PP group than in the G-PP group, CD31 expression remained relatively disorganized relative to eNOS expression in the native vessels. Sixty days after transplantation, CD31 and eNOS expression in the HPT-PP group approached that in the native vessel group, transitioning to a more organized arrangement, with fluorescence intensity showing no significant difference from that in the native vessel group. These findings are consistent with the endothelial cell growth behavior observed in vitro. The functionalized hydrogel layer on the HPT-PP surface promoted rapid and comprehensive endothelial cell coverage, thereby accelerating the reendothelialization process in vivo.

Claims

1. An anticoagulant coating for medical devices, characterized in that: The coating comprises thrombin-responsive nanoparticles, cleavable polypeptides, hydrogels, and active pharmaceutical ingredients.

2. The anticoagulation coating for medical devices according to claim 1, wherein The thrombin-responsive nanoparticles are composed of inorganic salt compounds, weak acid compounds, cleavable polypeptides and active pharmaceutical ingredients.

3. The anticoagulation coating for medical devices according to claim 2, wherein: The inorganic salt compound is selected from one or more of calcium carbonate, sodium periodate, sodium iodate, potassium iodate, and ammonium periodate; the weak acid compound is selected from one or more of sodium alginate, dextran, chondroitin sulfate, keratan sulfate, and oxidized hyaluronic acid; and the active pharmaceutical ingredient is selected from one or more of hirudin, heparin, rivaroxaban, and apixaban.

4. The anticoagulation coating for medical devices according to claim 2, wherein: The cleavable polypeptide is selected from NH2-Gly-(D)CHA-Ala-Arg-Ser-Trp-Gly-CONH2.

5. A method for preparing an anticoagulant coating for medical devices, characterized in that: The preparation method comprises the following steps: S1. Preparation of OX-HA: An inorganic salt compound powder is added to a weak acid compound solution. After stirring in the dark, an alcohol compound is added to interrupt the reaction. OX-HA is obtained by dialysis, freezing, and drying. S2. Preparation of Thrombin-responsive Nanoparticles (Thr NPs): Dissolve the OX-HA and active pharmaceutical ingredient obtained in Step 1, then add the responsive polypeptide chain dropwise to the solution and dialyze overnight to obtain Thr NPs. S3. The medical device to be coated is subjected to simple treatment and then soaked. The side chains, thrombin-responsive nanoparticles, and hydrogel are added to the soaking liquid to form an anticoagulant coating on the surface of the medical device.

6. The method for preparing an anticoagulant coating for medical devices according to claim 5, wherein: In step 1, the ratio of the inorganic salt compound, the weak acid compound, and the alcohol compound is 0.6:1:

1.

7. The method for preparing an anticoagulant coating for medical devices according to claim 5, wherein: In step 3, the medical device is selected from a bioprosthetic heart valve; the side chain is selected from branched polyethyleneimine, and the thrombin-responsive nanoparticles are the product prepared in step 2 of claim 5.

8. Use of the anticoagulant coating according to claim 1 in the preparation of regenerative medical devices.

9. Use of thrombin-responsive nanoparticles in the preparation of auxiliary medical products, characterized in that: The preparation method of the thrombin-responsive nanoparticles is as follows: S1. Preparation of OX-HA: An inorganic salt compound powder is added to a weak acid compound solution. After stirring in the dark, an alcohol compound is added to interrupt the reaction. OX-HA is obtained by dialysis, freezing, and drying. S2. Preparation of Thrombin-responsive Nanoparticles (Thr NPs): Dissolve the OX-HA and active pharmaceutical ingredient obtained in step 1, then add the responsive polypeptide chain dropwise to the solution and dialyze overnight to obtain Thr NPs.

10. Use of the thrombin-responsive nanoparticles according to claim 9 in preparing auxiliary medical products, characterized in that: The medical product is selected from one or more of a molecular probe containing thrombin-responsive nanoparticles, a reagent test kit, a tumor drug delivery system, and other auxiliary nanomaterials.

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