Preparation method of bionic artificial blood vessel based on self-healing hydrogel

Abstract

The invention relates to the technical field of medical devices, in particular to a preparation method of a bionic artificial blood vessel based on self-healing hydrogel. According to the method, plasma treatment and surface modification are carried out on a polytetrafluoroethylene tube, a self-healing hydrogel lining with radial component gradient is prepared, the self-healing hydrogel lining is composed of oxidized sodium alginate, adipic acid dihydrazine modified gelatin, carboxymethyl cellulose grafted phenylboronic acid polymer and the like, and a composite structure is formed through layer-by-layer injection and gelation. The obtained artificial blood vessel simulates the mechanical behavior of a natural blood vessel, has excellent compliance matching, self-healing ability and fatigue resistance, and effectively solves the problems that a traditional artificial blood vessel is easy to damage and poor in long-term patency.

Description

Technical Field

[0001] This invention relates to the field of medical device technology, and in particular to a method for preparing biomimetic artificial blood vessels based on self-healing hydrogels. Background Technology

[0002] Currently, in the field of vascular bypass grafting, synthetic materials such as polytetrafluoroethylene (PTFE) and polyester are widely used to manufacture artificial blood vessels due to their inherent bioinertness and ease of processing. However, these traditional materials have long faced a fundamental challenge after implantation: their mechanical properties, especially radial compliance, differ significantly from those of autologous blood vessels. This mechanical mismatch leads to changes in the hemodynamic environment, generating abnormal shear stress and turbulence at the anastomosis, becoming the initial trigger for intimal hyperplasia and anastomotic stenosis.

[0003] Furthermore, synthetic polymer materials, under long-term cyclic loading of pulsating pressure in vivo, are prone to developing microcracks within the material or at the interface with host tissue due to their inherent limited fatigue resistance. Once formed, these microcracks will gradually expand under continuous physiological load, potentially leading not only to graft leakage but, more importantly, triggering chronic inflammatory responses and possibly eventually developing into graft thrombosis or restenosis.

[0004] The aforementioned problems are particularly pronounced in small-diameter (<6mm) vascular graft applications. Due to the small diameter, even minor geometric changes or deterioration in mechanical properties can significantly impact blood flow. Existing technologies attempt to improve biocompatibility or mechanical properties through surface modification or composite materials, but these improvements are often static and cannot respond to the dynamic physiological environment in vivo. In particular, when the material suffers microscopic damage, there is a lack of effective self-repair mechanisms, and the accumulation of damage ultimately leads to graft failure.

[0005] Furthermore, existing artificial blood vessel designs often focus on optimizing single properties such as initial mechanical strength or anticoagulant properties, failing to systematically address the long-term stability of materials in dynamic physiological environments. An artificial blood vessel structure capable of mimicking the nonlinear mechanical behavior of natural blood vessels and possessing self-repair capabilities after damage is crucial for improving the long-term patency rate of small-diameter vascular grafts. Summary of the Invention

[0006] In view of this, the purpose of this invention is to propose a method for preparing biomimetic artificial blood vessels based on self-healing hydrogels, in order to solve the problem that existing artificial blood vessels made of synthetic materials such as polytetrafluoroethylene have mechanical compliance that does not match that of natural blood vessels, and are prone to micro-damage under long-term pulsating loads and performance degradation due to lack of self-healing ability, ultimately affecting the long-term patency of the implant.

[0007] To achieve the above objectives, this invention provides a method for preparing biomimetic artificial blood vessels based on self-healing hydrogels, comprising the following steps: (1) The polytetrafluoroethylene tube is cleaned and dried, and then subjected to plasma surface treatment in an oxygen-containing atmosphere to obtain a pretreated polytetrafluoroethylene tube. (2) The pretreated polytetrafluoroethylene tube is immersed in a monomer solution containing hydroxyethyl methacrylate and glycidyl methacrylate, and polymerized in the presence of a free radical initiator to form a hydrophilic polymer layer on the tube wall; then, a dopamine solution is injected into the tube cavity, and dopamine is self-polymerized under alkaline and oxidative conditions to form a polydopamine layer on the surface of the polymer layer, thus obtaining a surface-modified polytetrafluoroethylene tube; (3) Sodium alginate was oxidized by sodium periodate, sodium alginate was oxidized, dihydrazine adipic acid was modified gelatin by condensation reaction of dihydrazine adipic acid and pigskin gelatin, and carboxymethyl cellulose grafted phenylboronic acid polymer was obtained by condensation reaction of 4-aminophenylboronic acid and sodium carboxymethyl cellulose. (4) Prepare three types of hydrogel precursor solutions: inner layer, middle layer and outer layer. Each hydrogel precursor solution includes sodium alginate oxide, dihydrazine adipic acid modified gelatin, carboxymethyl cellulose grafted phenylboronic acid polymer, polyvinyl alcohol and calcium chloride. The contents of sodium alginate oxide, carboxymethyl cellulose grafted phenylboronic acid polymer, polyvinyl alcohol and calcium chloride gradually increase from the inner layer to the outer layer, while the contents of dihydrazine adipic acid modified gelatin gradually decrease. (5) The outer layer hydrogel precursor solution is injected into the lumen of the surface-modified polytetrafluoroethylene tube. After partial gelation under the condition of tube rotation, a portion of the precursor solution that is still in a sol state is extracted from the center of the tube to form a hollow channel. Then, the middle layer hydrogel precursor solution is injected into the hollow channel. Partial gelation is continued under the condition of rotation, and a portion of the precursor solution is extracted. Then, the inner layer hydrogel precursor solution is injected into the obtained channel. Gelation is completed under the condition of rotation, so that the inner wall of the polytetrafluoroethylene tube forms a self-healing hydrogel blood vessel wall with a composition gradient from the inside to the outside, and a polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel is obtained. (6) Post-processing of artificial blood vessels: The polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessels were extracted in a buffer salt solution, then disinfected in an ethanol aqueous solution, rinsed with a buffer salt solution and stored under sterile conditions to obtain biomimetic artificial blood vessels based on self-healing hydrogel.

[0008] Preferably, in step (1), the inner diameter of the polytetrafluoroethylene small-diameter tube is 4mm, the outer diameter is 6mm, and the length is 80mm.

[0009] Preferably, in step (1), the plasma surface treatment is performed by treating the surface with 100W power for 10 minutes in a radio frequency oxygen plasma surface treatment machine.

[0010] Preferably, in step (2), the monomer solution consists of 32-48g hydroxyethyl methacrylate, 8-12g glycidyl methacrylate, 0.8-1.2g azobisisobutyronitrile and 40-60mL anhydrous ethanol.

[0011] Preferably, in step (2), the dopamine solution is composed of 1.6-2.4g of dopamine hydrochloride and 160-240mL of Tris-HCl buffer solution; the concentration of the Tris-HCl buffer solution is 40-60mmol / L and the pH is 8.

[0012] Preferably, in step (3), the mass ratio of sodium periodate to sodium alginate is 3.2-4.8:8-12.

[0013] Preferably, in step (3), the mass ratio of dihydrazine adipic acid to porcine gelatin is 2.4-3.6:10.

[0014] Preferably, in step (3), the mass ratio of 4-aminophenylboronic acid to sodium carboxymethyl cellulose is 2.4-3.6:10.

[0015] Preferably, in step (3), the pigskin gelatin is Type A with a Bloom value of 300.

[0016] Preferably, in step (4), the inner layer hydrogel precursor solution is prepared by dissolving 1.6-2.4g of oxidized sodium alginate, 3.2-4.8g of dihydrazine adipic acid modified gelatin, 0.8-1.2g of carboxymethyl cellulose grafted phenylboronic acid polymer and 0.8-1.2g of polyvinyl alcohol in 40mL of phosphate buffer solution with pH 7.4 at 50°C, cooling to 37°C, adding 0.8-1.2g of dihydrazine adipic acid and stirring evenly, and finally adding 0.8-1.2g of calcium chloride and stirring until dissolved.

[0017] Preferably, in step (4), the intermediate hydrogel precursor solution is prepared by dissolving 2.4-3.6g of sodium alginate oxide, 2.4-3.6g of dihydrazine adipic acid modified gelatin, 1.6-2.4g of carboxymethyl cellulose grafted phenylboronic acid polymer and 1.6-2.4g of polyvinyl alcohol in 40mL of phosphate buffer solution with pH 7.4 at 50°C, cooling to 37°C, adding 0.8-1.2g of dihydrazine adipic acid and stirring evenly, and finally adding 1.6-2.4g of calcium chloride and stirring until dissolved.

[0018] Preferably, in step (4), the outer hydrogel precursor solution is prepared by dissolving 3.2-4.8g of oxidized sodium alginate, 1.6-2.4g of dihydrazine adipic acid modified gelatin, 2.4-3.6g of carboxymethyl cellulose grafted phenylboronic acid polymer and 2.4-3.6g of polyvinyl alcohol in 40mL of phosphate buffer solution with pH 7.4 at 50°C, stirring, cooling to 37°C, adding 0.8-1.2g of dihydrazine adipic acid and stirring evenly, and finally adding 2.4-3.6g of calcium chloride and stirring until dissolved.

[0019] Preferably, in step (4), the weight-average molecular weight of polyvinyl alcohol is 95,000 and the degree of alcoholysis is 99%.

[0020] Preferably, in step (5), the polytetrafluoroethylene tube is rotated at 30 rpm for 5 minutes and then left to stand for 10 minutes during the injection of the outer layer hydrogel precursor solution, the middle layer hydrogel precursor solution and the inner layer hydrogel precursor solution. The volume of the precursor liquid extracted from the center of the tube is one-third of the volume of the injected precursor liquid. Finally, after the inner layer hydrogel precursor solution is injected, the tube is left to stand at 37°C for 2 hours to complete the gelation.

[0021] Preferably, in step (6), the buffer salt solution is a phosphate buffer salt solution with a pH of 7.4.

[0022] The beneficial effects of this invention are: This invention constructs a self-healing hydrogel based on dynamic covalent bonds and multiple physical interactions as a vascular liner, which is then firmly bonded to polytetrafluoroethylene (PTFE) through a specific interface modification layer, forming a composite tubular structure with a radial compositional gradient. This structure enables the obtained artificial blood vessel to mimic the mechanical behavior of natural blood vessels, exhibiting excellent nonlinear elastic response. The inner hydrogel layer near the lumen is rich in a flexible gelatin network, giving it good initial deformation capability and effectively buffering pulsating pressure impacts; while the outer hydrogel layer provides sufficient support strength and creep resistance through an enhanced rigid network. This gradient mechanical property from the inside out significantly improves the compliance matching between the artificial blood vessel and the autologous blood vessel, helping to reduce harmful mechanical stimulation at the anastomosis site.

[0023] Specifically, the hydrogel contains a dual-dynamic cross-linked network comprised of reversible Schiff base bonds between oxidized sodium alginate and modified gelatin, and dynamic borate ester bonds between carboxymethyl cellulose grafted phenylboronic acid and polyvinyl alcohol. These two dynamic chemical bonds exhibit different breakage and reformation kinetics, enabling them to synergistically dissipate energy under stress. When microcracks occur locally in the material, the dynamic bonds in the network can rapidly break to prevent crack propagation and quickly reform after stress relaxation, achieving autonomous repair of microscopic damage. This mechanism significantly enhances the fatigue resistance and structural integrity of artificial blood vessels under long-term cyclic loading.

[0024] A robust interfacial bond is achieved between the PTFE inner wall and the self-healing hydrogel by sequentially introducing a hydrophilic polymer graft layer and a polydopamine coating. The hydrophilic polymer layer improves the wettability of the PTFE surface and provides abundant active grafting sites, while the polydopamine layer, through its strong adhesion properties and various intermolecular interactions with the hydrogel components, forms a tough interfacial transition zone. This structure ensures effective stress transfer between the hydrogel liner and the PTFE body during repeated expansion and contraction of the blood vessel wall, preventing interfacial delamination and thus guaranteeing the long-term stability of the composite structure.

[0025] By optimizing the ratio and spatial distribution of each component, the resulting hydrogel maintains a high water content to mimic the softness of biological soft tissue while exhibiting excellent anti-permeability. The tight cross-linked network effectively limits excessive migration of water molecules, ensuring low permeability of the blood vessel wall and reducing the risk of postoperative complications due to leakage. Furthermore, after thorough purification and post-processing, the material demonstrates good biocompatibility, laying the foundation for its long-term safe service in vivo.

[0026] In summary, the artificial blood vessel provided by this invention has excellent overall performance. Its unique gradient structure design and the introduction of multiple dynamic self-healing mechanisms effectively solve the shortcomings of traditional synthetic blood vessels in terms of compliance matching, long-term fatigue resistance and self-repair of damage, providing a better alternative for small-diameter blood vessel transplantation. Attached Figure Description

[0027] To more clearly illustrate the technical solutions in this invention or the prior art, the accompanying drawings used in the description of the embodiments or the prior art will be briefly introduced below.

[0028] Figure 1 The infrared spectra of sodium alginate oxide, dihydrazine adipic acid modified gelatin, carboxymethyl cellulose grafted phenylboronic acid polymer and self-healing hydrogel ring prepared in Example 2 of the present invention are shown. Detailed Implementation

[0029] To make the objectives, technical solutions, and advantages of this invention clearer, the invention will be further described in detail below with reference to specific embodiments.

[0030] Example 1: Step 1: Pretreatment of PTFE small-diameter pipe supports Take a polytetrafluoroethylene (PTFE) tube with an inner diameter of 4 mm, an outer diameter of 6 mm, and a length of 80 mm. After cutting, grind both ends flat. Add 500 mL of anhydrous ethanol and 500 mL of deionized water to a beaker to obtain an ethanol-water solution. Completely immerse the PTFE tube in the solution and ultrasonically clean it for 30 minutes. After removing it, rinse it with deionized water from both the inside and outside of the tube. Place it in a 60°C forced-air oven to dry for 2 hours. After cooling, place the dried tube into the cavity of a radio frequency oxygen plasma surface treatment machine. Use oxygen as the working gas and perform plasma treatment at a power of 100 W and a treatment time of 10 minutes to obtain a pretreated PTFE tube.

[0031] Step 2: Surface modification of polytetrafluoroethylene (PTFE) tubing In a three-necked flask, 32g of hydroxyethyl methacrylate, 8g of glycidyl methacrylate, 0.8g of azobisisobutyronitrile, and 40mL of anhydrous ethanol were added sequentially. After stirring evenly, the pretreated polytetrafluoroethylene tube was completely immersed in the monomer solution. Nitrogen gas was turned on, and the system was bubbled to deoxygenate for 20 minutes. The reaction was then carried out at a constant temperature of 70℃ for 3 hours to form a poly(hydroxyethyl methacrylate-co-glycidyl methacrylate) layer. After the reaction was completed, the tube was removed from the monomer solution, rinsed inside and outside the tube with anhydrous ethanol, rinsed with deionized water, and dried in a 40℃ oven for 6 hours. Subsequently, one end of the dried tube was sealed, and a prepared dopamine solution was injected into the tube lumen: 1.6 g of dopamine hydrochloride was dissolved in 240 mL of TrisHCl buffer solution (40 mmol / L, pH 8), placed in a brown bottle and stirred in the dark to ensure complete dissolution, and then the solution was poured into the tube lumen. The tube was shaken at room temperature and the reaction was continued under nitrogen protection for 18 h. Then the tube lumen was rinsed with deionized water to obtain a surface-modified polytetrafluoroethylene tube.

[0032] Step 3: Preparation of sodium alginate oxide Add 8g of sodium alginate to 400mL of deionized water and stir at room temperature until completely dissolved to obtain a homogeneous solution. In a separate beaker, add 3.2g of sodium periodate and 80mL of deionized water, stir to dissolve, and wrap both beakers with aluminum foil to protect them from light. Then, under room temperature and light-protected conditions, slowly add the sodium periodate solution to the sodium alginate solution dropwise over 30 minutes. After the addition is complete, continue stirring in the dark for 3 hours. After the reaction is complete, add 4g of glycerol and continue stirring for 30 minutes. Then, transfer the reaction solution to a dialysis bag with a molecular weight cutoff of 7000 Da and dialyze against deionized water for 3 days, changing the external solution 3 times a day to remove inorganic salts and low-molecular-weight byproducts. Finally, freeze-dry the dialysate to obtain oxidized sodium alginate.

[0033] Step 4: Preparation of dihydrazine adipic acid modified gelatin 10g of pigskin gelatin (Type A, Bloom value 300) was weighed and added to 100mL of 2(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5). The solution was stirred at 50℃ until completely dissolved, then cooled to 35℃. While continuously stirring, 2.4g of dihydrazine adipic acid, 2.4g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride, and 1.6g of N-hydroxysuccinimide were added to the solution. The reaction was continued at 35℃ for 18h. After the reaction, the reaction solution was placed in a dialysis bag with a molecular weight cutoff of 8000 Da and dialyzed against deionized water for 3 days. Subsequently, it was freeze-dried to obtain dihydrazine adipic acid modified gelatin.

[0034] Step 5: Preparation of carboxymethyl cellulose grafted phenylboronic acid polymer 10g of sodium carboxymethyl cellulose (purchased from Shanghai Aladdin Biotechnology Co., Ltd., catalog number C104987) was weighed and added to 200mL of 2(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5) and stirred until completely dissolved. In another beaker, 2.4g of 4-aminophenylboronic acid was dissolved in 50mL of the same buffer solution. The 4-aminophenylboronic acid solution was slowly poured into the carboxymethyl cellulose solution. Under stirring, 2.4g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride and 1.6g of N-hydroxysuccinimide were added sequentially. The reaction system temperature was maintained at 25℃ and the pH was controlled at around 5. The reaction was carried out for 18h. The reaction solution was then placed in a dialysis bag with a molecular weight cutoff of 7000Da and dialyzed with deionized water for 3 days. Finally, it was freeze-dried to obtain the carboxymethyl cellulose-grafted phenylboronic acid polymer. Step 6: Prepare three different formulations of hydrogel precursor solutions A / B / C In three separate 250mL beakers, the precursor solutions for the inner layer (A), middle layer (B), and outer layer (C) of hydrogel were prepared respectively. The operation of each beaker was as follows: (1) Inner layer precursor solution A: 1.6g of sodium alginate oxide, 3.2g of dihydrazine adipic acid modified gelatin, 0.8g of carboxymethyl cellulose grafted phenylboronic acid polymer and 0.8g of polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%) were added to the first beaker in sequence, and then 40mL of phosphate buffer solution (pH 7.4) was added. The mixture was stirred at 50℃ until the solid was completely dissolved. After cooling to 37℃, 0.8g of adipic acid was added. (2) Intermediate layer precursor solution B: In the second beaker, add 2.4g sodium alginate oxide, 2.4g dihydrazine adipic acid modified gelatin, 1.6g carboxymethyl cellulose grafted phenylboronic acid polymer and 1.6g polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%), then add 40mL phosphate buffer solution (pH 7.4), stir at 50℃ until completely dissolved, cool to 37℃ and add 0.8g dihydrazine adipic acid and stir evenly, then add 1.6g calcium chloride and stir until completely dissolved. (3) Outer layer precursor solution C: In the third beaker, 3.2g of sodium alginate oxide, 1.6g of dihydrazine adipic acid modified gelatin, 2.4g of carboxymethyl cellulose grafted phenylboronic acid polymer and 2.4g of polyvinyl alcohol (weight average molecular weight 95,000, degree of alcoholysis 99%) were added in sequence, followed by 40mL of phosphate buffer solution (pH 7.4). The mixture was stirred at 50℃ to dissolve the solid. After cooling to 37℃, 0.8g of dihydrazine adipic acid was added and stirred evenly. Then, 2.4g of calcium chloride was added and stirred until completely dissolved. The precursor solutions in all three beakers were kept in a 37℃ water bath for standby to ensure that the gelatin remained in a sol state and did not undergo thermal gelation.

[0035] Step 7: Constructing a radially gradient self-healing hydrogel vessel wall A surface-modified polytetrafluoroethylene (PTFE) tube was fixed to a support, with one end sealed with a medical-grade silicone stopper and the other end connected to a syringe via a Luer connector. The entire tube was placed at 37°C. First, the outer precursor solution C was slowly injected into the tube lumen using a 50mL syringe until it was completely filled. During the slow injection, the tube was continuously rotated at 30rpm at 25°C for 5 minutes, and then allowed to stand for 10 minutes. Subsequently, about one-third of the volume of the precursor solution C, which was still in a sol state, was slowly withdrawn from the center of the tube using a syringe, leaving a through hollow channel in the inner lumen. Immediately inject the intermediate layer precursor solution B into the second syringe to fill the inner side of the partially gelled outer shell. Continue to rotate at 30 rpm for 5 minutes at 25°C and let stand for 10 minutes to form the intermediate layer. Repeat the above operation and extract about one-third of the volume of the intermediate layer precursor solution B that has not yet fully gelled from the center. Then inject the inner layer precursor solution A into the third syringe, continue to rotate at 30 rpm for 5 minutes at 25°C and let stand at 37°C for 2 hours to obtain polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel. Step 8: Post-processing of artificial blood vessels After sealing both ends of the polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel obtained in step 7, it was placed in a beaker containing 500 mL of phosphate buffer solution (pH 7.4) and immersed in it at 60 rpm on a shaker at 37°C for 24 h. The phosphate buffer solution (pH 7.4) was replaced every 6 h during this period. After the extraction was completed, the sample was removed and transferred to a beaker containing 500 mL of ethanol aqueous solution (75% by volume) for 24 h at room temperature for sterilization. Then, it was rinsed repeatedly three times with 500 mL of phosphate buffer solution (pH 7.4). Finally, the artificial blood vessel was immersed in phosphate buffer solution (pH 7.4) at 37°C under sterile conditions for later use, thus obtaining a biomimetic artificial blood vessel based on self-healing hydrogel.

[0036] Example 2: Step 1: Pretreatment of PTFE small-diameter pipe supports Take a polytetrafluoroethylene (PTFE) tube with an inner diameter of 4 mm, an outer diameter of 6 mm, and a length of 80 mm. After cutting, grind both ends flat. Add 500 mL of anhydrous ethanol and 500 mL of deionized water to a beaker to obtain an ethanol-water solution. Completely immerse the PTFE tube in the solution and ultrasonically clean it for 30 minutes. After removing it, rinse it with deionized water from both the inside and outside of the tube. Place it in a 60°C forced-air oven to dry for 2 hours. After cooling, place the dried tube into the cavity of a radio frequency oxygen plasma surface treatment machine. Use oxygen as the working gas and perform plasma treatment at a power of 100 W and a treatment time of 10 minutes to obtain a pretreated PTFE tube.

[0037] Step 2: Surface modification of polytetrafluoroethylene (PTFE) tubing In a three-necked flask, 40g of hydroxyethyl methacrylate, 10g of glycidyl methacrylate, 1g of azobisisobutyronitrile, and 50mL of anhydrous ethanol were added sequentially. After stirring evenly, the pretreated polytetrafluoroethylene tube was completely immersed in the monomer solution. Nitrogen gas was turned on, and the system was bubbled to deoxygenate for 30 minutes. The reaction was then carried out at a constant temperature of 70℃ for 4 hours to form a poly(hydroxyethyl methacrylate-co-glycidyl methacrylate) layer. After the reaction was completed, the tube was removed from the monomer solution, rinsed inside and outside the tube with anhydrous ethanol, rinsed with deionized water, and dried in a 40℃ oven for 6 hours. Subsequently, one end of the dried tube was sealed, and a prepared dopamine solution was injected into the tube lumen: 2g of dopamine hydrochloride was dissolved in 200mL of Tris-HCl buffer solution (50mmol / L, pH 8), placed in a brown bottle in the dark and stirred until fully dissolved, and then the tube lumen was filled. The solution was shaken at room temperature and the reaction was continued under nitrogen protection for 24h. Then the tube lumen was rinsed with deionized water to obtain a surface-modified polytetrafluoroethylene tube.

[0038] Step 3: Preparation of sodium alginate oxide 10g of sodium alginate was added to 500mL of deionized water and stirred at room temperature until completely dissolved to obtain a homogeneous solution. In a separate beaker, 4g of sodium periodate and 100mL of deionized water were added and stirred until dissolved. Both beakers were then wrapped with aluminum foil to protect them from light. Subsequently, under room temperature and light-protected conditions, the sodium periodate solution was slowly added dropwise to the sodium alginate solution, controlling the addition time at 30 minutes. After the addition was complete, the reaction was continued in the dark with stirring for 4 hours. After the reaction was complete, 5g of glycerol was added, and stirring was continued for 30 minutes. The reaction solution was then placed in a dialysis bag with a molecular weight cutoff of 7000 Da, and dialyzed against deionized water for 3 days, changing the external solution 3 times a day to remove inorganic salts and low-molecular-weight byproducts. The dialysate was then freeze-dried to obtain oxidized sodium alginate.

[0039] Step 4: Preparation of dihydrazine adipic acid modified gelatin 10g of pigskin gelatin (Type A, Bloom value 300) was weighed and added to 100mL of 2-(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5). The solution was stirred at 50℃ until completely dissolved, and then cooled to 35℃. 3g of dihydrazine adipic acid, 3g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride, and 2g of N-hydroxysuccinimide were added to the solution while continuously stirring. The reaction was continued at 35℃ for 24h. After the reaction was completed, the reaction solution was placed in a dialysis bag with a molecular weight cutoff of 8000 Da and dialyzed against deionized water for 3 days. The solution was then freeze-dried to obtain dihydrazine adipic acid modified gelatin.

[0040] Step 5: Preparation of carboxymethyl cellulose grafted phenylboronic acid polymer 10g of sodium carboxymethyl cellulose (purchased from Shanghai Aladdin Biotechnology Co., Ltd., catalog number C104987) was weighed and added to 200mL of 2-(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5) and stirred until completely dissolved. In another beaker, 3g of 4-aminophenylboronic acid was dissolved in 50mL of the same buffer solution. The solution containing 4-aminophenylboronic acid was slowly poured into the carboxymethyl cellulose solution. Under stirring, 3g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride and 2g of N-hydroxysuccinimide were added sequentially. The reaction system temperature was maintained at 25℃ and the pH was controlled at around 5. The reaction was carried out for 24h. Then, the reaction solution was placed in a dialysis bag with a molecular weight cutoff of 7000Da and dialyzed with deionized water for 3 days. Finally, it was freeze-dried to obtain the carboxymethyl cellulose-grafted phenylboronic acid polymer. Step 6: Prepare three different formulations of hydrogel precursor solutions A / B / C In three separate 250mL beakers, the precursor solutions for the inner layer (A), middle layer (B), and outer layer (C) of hydrogel were prepared respectively. The operation for each beaker was as follows: (1) Inner layer precursor solution A: 2g of sodium alginate oxidized, 4g of dihydrazine adipic acid modified gelatin, 1g of carboxymethyl cellulose grafted phenylboronic acid polymer and 1g of polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%) were added to the first beaker in sequence, followed by 40mL of phosphate buffer solution (pH 7.4), and stirred at 50℃ until the solid was completely dissolved. After cooling to 37℃, 1g of dihydrazine adipic acid was added. (2) Intermediate layer precursor solution B: In the second beaker, add 3g of sodium alginate oxide, 3g of gelatin modified with dihydrazine adipic acid, 2g of carboxymethyl cellulose grafted with phenylboronic acid polymer and 2g of polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%), then add 40mL of phosphate buffer solution (pH 7.4), stir at 50℃ until completely dissolved, cool to 37℃ and add 1g of dihydrazine adipic acid and stir evenly, then add 2g of calcium chloride and stir until completely dissolved. (3) Outer layer precursor solution C: In the third beaker, add 4g of sodium alginate oxide, 2g of dihydrazine adipic acid modified gelatin, 3g of carboxymethyl cellulose grafted phenylboronic acid polymer and 3g of polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%), and add 40mL of phosphate buffer solution (pH 7.4). Stir at 50℃ to dissolve the solid. After cooling to 37℃, add 1g of dihydrazine adipic acid, stir evenly, and then add 3g of calcium chloride. Stir until completely dissolved. The precursor solutions in all three beakers are kept in a 37℃ water bath for later use to ensure that the gelatin is in a sol state and does not undergo thermal gelation.

[0041] Step 7: Constructing a radially gradient self-healing hydrogel vessel wall A surface-modified polytetrafluoroethylene (PTFE) tube was fixed to a support, with one end sealed with a medical-grade silicone stopper and the other end connected to a syringe via a Luer connector. The entire tube was placed at 37°C. First, the outer precursor solution C was slowly injected into the tube lumen using a 50mL syringe until it was completely filled. During the slow injection, the tube was continuously rotated at 30rpm for 5 minutes at 25°C, and then allowed to stand for 10 minutes. Subsequently, about one-third of the volume of the precursor solution C, which was still in a sol state, was slowly withdrawn from the center of the tube using a syringe, leaving a through hollow channel in the inner lumen. Immediately inject the intermediate layer precursor solution B into the second syringe to fill the inner side of the partially gelled outer shell. Continue to rotate at 30 rpm for 5 minutes at 25°C and let stand for 10 minutes to form the intermediate layer. Repeat the above operation and extract about one-third of the volume of the intermediate layer precursor solution B that has not yet fully gelled from the center. Then inject the inner layer precursor solution A into the third syringe, continue to rotate at 30 rpm for 5 minutes at 25°C, and let stand at 37°C for 2 hours to obtain polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel. Step 8: Post-processing of artificial blood vessels After sealing both ends of the polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel obtained in step 7, it was placed in a beaker containing 500 mL of phosphate buffer solution (pH 7.4) and immersed in it at 60 rpm on a shaker at 37°C for 24 h. The phosphate buffer solution (pH 7.4) was replaced every 6 h during this period. After the extraction was completed, the sample was removed and transferred to a beaker containing 500 mL of ethanol aqueous solution (75% by volume) for 24 h at room temperature for sterilization. Then, it was rinsed repeatedly three times with 500 mL of phosphate buffer solution (pH 7.4). Finally, the artificial blood vessel was immersed in phosphate buffer solution (pH 7.4) at 37°C under sterile conditions for later use, thus obtaining a biomimetic artificial blood vessel based on self-healing hydrogel.

[0042] Example 3: Step 1: Pretreatment of PTFE small-diameter pipe supports Take a polytetrafluoroethylene (PTFE) tube with an inner diameter of 4 mm, an outer diameter of 6 mm, and a length of 80 mm. After cutting, grind both ends flat. Add 500 mL of anhydrous ethanol and 500 mL of deionized water to a beaker to obtain an ethanol-water solution. Completely immerse the PTFE tube in the solution and ultrasonically clean it for 30 minutes. After removing it, rinse it with deionized water from both the inside and outside of the tube. Place it in a 60°C forced-air oven to dry for 2 hours. After cooling, place the dried tube into the cavity of a radio frequency oxygen plasma surface treatment machine. Use oxygen as the working gas and perform plasma treatment at a power of 100 W and a treatment time of 10 minutes to obtain a pretreated PTFE tube.

[0043] Step 2: Surface modification of polytetrafluoroethylene (PTFE) tubing In a three-necked flask, 48g of hydroxyethyl methacrylate, 12g of glycidyl methacrylate, 1.2g of azobisisobutyronitrile, and 60mL of anhydrous ethanol were added sequentially. After stirring evenly, the pretreated polytetrafluoroethylene tube was completely immersed in the monomer solution. Nitrogen gas was turned on, and the system was bubbled to deoxygenate for 40 minutes. The reaction was then carried out at a constant temperature of 70℃ for 5 hours to form a poly(hydroxyethyl methacrylate-co-glycidyl methacrylate) layer. After the reaction was completed, the tube was removed from the monomer solution, rinsed inside and outside the tube with anhydrous ethanol, rinsed with deionized water, and dried in a 40℃ oven for 6 hours. Subsequently, one end of the dried tube was sealed, and a prepared dopamine solution was injected into the tube lumen: 2.4 g of dopamine hydrochloride was dissolved in 160 mL of TrisHCl buffer solution (60 mmol / L, pH 8), placed in a brown bottle in the dark and stirred until fully dissolved, and then the tube lumen was filled. The solution was shaken at room temperature and the reaction was continued for 30 h under nitrogen protection. Then the tube lumen was rinsed with deionized water to obtain a surface-modified polytetrafluoroethylene tube.

[0044] Step 3: Preparation of sodium alginate oxide 12g of sodium alginate was added to 600mL of deionized water and stirred at room temperature until completely dissolved to obtain a homogeneous solution. In a separate beaker, 4.8g of sodium periodate and 120mL of deionized water were added and stirred until dissolved. Both beakers were then wrapped with aluminum foil to protect them from light. Subsequently, under room temperature and light-protected conditions, the sodium periodate solution was slowly added dropwise to the sodium alginate solution, controlling the addition time at 30min. After the addition was complete, the reaction was continued in the dark with stirring for 5h. After the reaction was complete, 6g of glycerol was added, and stirring was continued for 30min. The reaction solution was then placed in a dialysis bag with a molecular weight cutoff of 7000Da and dialyzed against deionized water for 3 days, changing the external solution 3 times a day to remove inorganic salts and low-molecular-weight byproducts. The dialysate was then freeze-dried to obtain oxidized sodium alginate.

[0045] Step 4: Preparation of dihydrazine adipic acid modified gelatin 10g of pigskin gelatin (Type A, Bloom value 300) was weighed and added to 100mL of 2(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5). The solution was stirred at 50℃ until completely dissolved, then cooled to 35℃. While continuously stirring, 3.6g of dihydrazine adipic acid, 3.6g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride, and 2.4g of N-hydroxysuccinimide were added to the solution. The reaction was continued at 35℃ for 30h. After the reaction, the reaction solution was placed in a dialysis bag with a molecular weight cutoff of 8000Da and dialyzed against deionized water for 3 days. Subsequently, it was freeze-dried to obtain dihydrazine adipic acid modified gelatin.

[0046] Step 5: Preparation of carboxymethyl cellulose grafted phenylboronic acid polymer 10g of sodium carboxymethyl cellulose (purchased from Shanghai Aladdin Biotechnology Co., Ltd., catalog number C104987) was weighed and added to 200mL of 2(N-morpholine) ethanesulfonic acid buffer solution (0.1mol / L, pH 5) and stirred until completely dissolved. In another beaker, 3.6g of 4-aminophenylboronic acid was dissolved in 50mL of the same buffer solution. The 4-aminophenylboronic acid solution was slowly poured into the carboxymethyl cellulose solution. Under stirring, 3.6g of 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride and 2.4g of N-hydroxysuccinimide were added sequentially. The reaction system temperature was maintained at 25℃ and the pH was controlled at around 5. The reaction was carried out for 30h. Then, the reaction solution was placed in a dialysis bag with a molecular weight cutoff of 7000Da and dialyzed with deionized water for 3 days. Finally, it was freeze-dried to obtain the carboxymethyl cellulose-grafted phenylboronic acid polymer. Step 6: Prepare three different formulations of hydrogel precursor solutions A / B / C In three separate 250mL beakers, the precursor solutions for the inner layer (A), middle layer (B), and outer layer (C) of hydrogel were prepared respectively. The operation of each beaker was as follows: (1) Inner layer precursor solution A: 2.4g of sodium alginate oxide, 4.8g of dihydrazine adipic acid modified gelatin, 1.2g of carboxymethyl cellulose grafted phenylboronic acid polymer and 1.2g of polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%) were added to the first beaker in sequence, and then 40mL of phosphate buffer solution (pH 7.4) was added. The mixture was stirred at 50℃ until the solid was completely dissolved. After cooling to 37℃, 1.2g of adipic acid was added. (2) Intermediate layer precursor solution B: In the second beaker, add 3.6g sodium alginate oxide, 3.6g dihydrazine adipic acid modified gelatin, 2.4g carboxymethyl cellulose grafted phenylboronic acid polymer and 2.4g polyvinyl alcohol (weight average molecular weight 95000, degree of alcoholysis 99%), then add 40mL phosphate buffer solution (pH 7.4), stir at 50℃ until completely dissolved, cool to 37℃ and add 1.2g dihydrazine adipic acid and stir evenly, then add 2.4g calcium chloride and stir until completely dissolved. (3) Outer layer precursor solution C: In the third beaker, 4.8g of sodium alginate oxide, 2.4g of dihydrazine adipic acid modified gelatin, 3.6g of carboxymethyl cellulose grafted phenylboronic acid polymer and 3.6g of polyvinyl alcohol (weight average molecular weight 95,000, degree of alcoholysis 99%) were added in sequence, followed by 40mL of phosphate buffer solution (pH 7.4). The mixture was stirred at 50℃ to dissolve the solid. After cooling to 37℃, 1.2g of dihydrazine adipic acid was added and stirred evenly. Then, 3.6g of calcium chloride was added and stirred until completely dissolved. The precursor solutions in all three beakers were kept in a 37℃ water bath for standby to ensure that the gelatin remained in a sol state and did not undergo thermal gelation.

[0047] Step 7: Constructing a radially gradient self-healing hydrogel vessel wall A surface-modified polytetrafluoroethylene (PTFE) tube was fixed to a support, with one end sealed with a medical-grade silicone stopper and the other end connected to a syringe via a Luer connector. The entire tube was placed at 37°C. First, the outer precursor solution C was slowly injected into the tube lumen using a 50mL syringe until it was completely filled. During the slow injection, the tube was continuously rotated at 30rpm for 5 minutes at 25°C, and then allowed to stand for 10 minutes. Subsequently, about one-third of the volume of the precursor solution C, which was still in a sol state, was slowly withdrawn from the center of the tube using a syringe, leaving a through hollow channel in the inner lumen. Immediately inject the intermediate layer precursor solution B into the second syringe to fill the inner side of the partially gelled outer shell. Continue to rotate at 30 rpm for 5 minutes at 25°C and let stand for 10 minutes to form the intermediate layer. Repeat the above operation and extract about one-third of the volume of the intermediate layer precursor solution B that has not yet fully gelled from the center. Then inject the inner layer precursor solution A into the third syringe, continue to rotate at 30 rpm for 5 minutes at 25°C, and let stand at 37°C for 2 hours to obtain polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel. Step 8: Post-processing of artificial blood vessels After sealing both ends of the polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel obtained in step 7, it was placed in a beaker containing 500 mL of phosphate buffer solution (pH 7.4) and immersed in it at 60 rpm on a shaker at 37°C for 24 h. The phosphate buffer solution (pH 7.4) was replaced every 6 h during this period. After the extraction was completed, the sample was removed and transferred to a beaker containing 500 mL of ethanol aqueous solution (75% by volume) for 24 h at room temperature for sterilization. Then, it was rinsed repeatedly three times with 500 mL of phosphate buffer solution (pH 7.4). Finally, the artificial blood vessel was immersed in phosphate buffer solution (pH 7.4) at 37°C under sterile conditions for later use, thus obtaining a biomimetic artificial blood vessel based on self-healing hydrogel.

[0048] Comparative Example 1: The dopamine layer was removed, and only the hydrophilic copolymer layer was retained. The difference between Comparative Example 1 and Example 2 is that in step 2, the pretreated polytetrafluoroethylene tube was immersed in a monomer solution containing hydroxyethyl methacrylate, glycidyl methacrylate, and azobisisobutyronitrile, and reacted at 70°C for 4 hours to form a poly(hydroxyethyl methacrylate-co-glycidyl methacrylate) layer. After rinsing with anhydrous ethanol and deionized water and drying at 40°C for 6 hours, one end was no longer sealed and a Tris-HCl buffer solution of dopamine hydrochloride was injected into the tube. Instead of reacting under nitrogen protection for 24 hours, the polytetrafluoroethylene / hydrophilic copolymer structure obtained by rinsing with deionized water was used as the surface-modified polytetrafluoroethylene tube and entered steps 3-8. The remaining steps and conditions were the same as in Example 2.

[0049] Comparative Example 2: The hydrophilic copolymer layer was omitted, and only the dopamine layer was constructed. The difference between Comparative Example 2 and Example 2 is that in step 2, the process of immersing the pretreated polytetrafluoroethylene tube in a monomer solution containing hydroxyethyl methacrylate, glycidyl methacrylate, and azobisisobutyronitrile and reacting it at 70°C for 4 hours to form a poly(hydroxyethyl methacrylate-co-glycidyl methacrylate) layer is omitted. Instead, one end of the pretreated polytetrafluoroethylene tube obtained in step 1 is directly sealed, and the dopamine solution prepared in Example 2 (2g of dopamine hydrochloride dissolved in 200mL Tris-HCl buffer solution, 50mmol / L, pH 8) is injected into the tube. After shaking and reacting for 24 hours at room temperature under nitrogen protection, it is rinsed with deionized water to obtain a surface-modified polytetrafluoroethylene tube containing only a dopamine layer, which is then proceeded to subsequent steps 3-8. The remaining steps and conditions are the same as in Example 2.

[0050] Comparative Example 3: Carboxymethyl cellulose grafted with phenylboronic acid polymer was removed, and only the Schiff base network was retained. The difference between Comparative Example 3 and Example 2 is that the preparation process of carboxymethyl cellulose grafted phenylboronic acid polymer in step 5 of Example 2 is not performed. Instead, in step 6 when preparing the three hydrogel precursor solutions A / B / C, 1g, 2g, and 3g of sodium carboxymethyl cellulose are directly added to replace the same mass of carboxymethyl cellulose grafted phenylboronic acid polymer in Example 2. The remaining steps and conditions are the same as in Example 2.

[0051] Comparative Example 4: Adihydrazide-modified gelatin was removed, and only the borate ester network was retained. The difference between Comparative Example 4 and Example 2 is as follows: In step 4, dihydrazine adipic acid, 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide hydrochloride, and N-hydroxysuccinimide were not added to the pigskin gelatin solution. Instead, the pigskin gelatin was completely dissolved at 50°C, cooled to 35°C, and then directly placed into a dialysis bag with a molecular weight cutoff of 8000 Da. The solution was dialyzed against deionized water for 3 days and then freeze-dried to obtain unmodified pigskin gelatin. In step 6, the original dihydrazine adipic acid modified gelatin in the three hydrogel precursor solutions A / B / C was completely replaced with an equal mass of the above-mentioned pigskin gelatin, and no additional 1g of dihydrazine adipic acid was added. The remaining steps and conditions were the same as in Example 2.

[0052] Comparative Example 5: The radial component gradient was eliminated, and the three-layer formulation was unified into a mid-layer formulation. The difference between Comparative Example 5 and Example 2 is that, in step 6, the inner layer precursor solution A, the middle layer precursor solution B, and the outer layer precursor solution C are no longer prepared separately. Instead, the hydrogel precursor solution is prepared in three 250mL beakers according to the formula and steps of the middle layer precursor solution B in Example 2. That is, 3g of sodium alginate oxide, 3g of dihydrazine adipic acid modified gelatin, 2g of carboxymethyl cellulose grafted phenylboronic acid polymer, and 2g of polyvinyl alcohol are added to each beaker in sequence, and 40mL of phosphate buffer solution (pH 7.4) is added and stirred at 50°C. Dissolve the solution by stirring, and after cooling to 37°C, add 1g of dihydrazine adipic acid and 2g of calcium chloride and stir until completely dissolved. In step 7, follow the same suction and re-injection process as in Example 2: first inject the first cup of precursor solution, then extract about one-third of the sol, inject the second cup of precursor solution and repeat the operation, then inject the third cup of precursor solution. The rotation speed, rotation time, standing time, and standing at 37°C for 2 hours remain unchanged, thereby forming a hydrogel blood vessel wall with basically uniform composition and crosslinking density inside the polytetrafluoroethylene tube. The remaining steps and conditions are the same as in Example 2.

[0053] Comparative Example 6: Reversing the radial component gradient direction (hard inside, soft outside) The difference between Comparative Example 6 and Example 2 is that the formulations and preparation methods of the three precursor solutions A / B / C in step 6 are exactly the same as in Example 2, while in step 7, only the injection order and corresponding layers are adjusted: First, the inner layer precursor solution A is slowly injected into the tube using a 50mL syringe until it is completely filled. After rotating at 30rpm for 5 minutes and standing at 37°C for 10 minutes, about one-third of the volume of sol is withdrawn from the center of the tube; then, the middle layer precursor solution B is injected and the rotation and standing are repeated, and about one-third of the volume of the precursor solution that has not yet fully gelled is withdrawn again; finally, the middle layer precursor solution B is injected. The outer precursor solution C was rotated at 30 rpm for 5 min and then allowed to stand at 37°C for 2 h. The remaining steps, such as the sealing method, extraction volume ratio, and post-processing step 8, were the same as in Example 2. The side near the inner wall of the polytetrafluoroethylene was a hydrogel layer with a low proportion of sodium alginate and carboxymethyl cellulose grafted phenylboronic acid polymer, a high content of dihydrazine adipic acid modified gelatin, and a low content of calcium chloride. The inner cavity side was a hydrogel layer with a high content of sodium alginate, carboxymethyl cellulose grafted phenylboronic acid polymer, polyvinyl alcohol, and calcium chloride, thus constructing an anti-gradient structure.

[0054] Performance testing: Water absorption test: The water absorption test is performed according to GB / T 1034-2008. Each sample is cut into a 10mm long annular tube along the axial direction. The outer polytetrafluoroethylene body is gently peeled off, leaving only the self-healing hydrogel ring. The width and thickness of the annular sample are measured and the dry weight m0 is recorded. The sample is completely immersed in a phosphate buffer solution with pH 7.4 and left to stand at 37℃ for 24 hours. After that, it is taken out and the surface moisture is quickly blotted with filter paper. The wet weight m is then measured. t According to the water absorption rate W = (mt The calculation is performed using (-m0) / m0×100%, and the results are shown in Table 1.

[0055] Infrared characterization: Fourier transform infrared spectroscopy was used to characterize the sodium alginate oxide, dihydrazine adipic acid modified gelatin, carboxymethyl cellulose grafted phenylboronic acid polymer, and the self-healing hydrogel ring prepared according to the water absorption test prepared in Example 2. The results are as follows: Figure 1 As shown.

[0056] Interfacial shear adhesion test: Referring to GB / T 7124-2008, the composite artificial blood vessel was cut along the axial direction and unfolded into a flat piece of 25mm×100mm. The pieces were folded inward with the hydrogel layer facing inward, so that the hydrogel layers of the two polytetrafluoroethylene bodies were bonded together. Light pressure was applied at 37°C to ensure full contact of the hydrogel layers, and the pieces were left to stand in a moist state for 30 minutes to form a stable interface. Then, the pieces were loaded on a universal testing machine according to the single lap tensile shear method specified in GB / T 7124-2008, with an lap length of 12.5mm and a loading speed of 10mm / min. The maximum breaking load was recorded, and the interfacial shear strength was calculated. The results are shown in Table 1.

[0057] Radial compliance test: Referring to the requirements for dynamic radial compliance test of tubular vascular grafts in YY / T 0500-2021 and in conjunction with the relevant test instrument instructions, the compliance was measured under simulated arterial pulsation pressure conditions at 37℃. The sample was cut into a tube with a length of 80mm and placed in a constant temperature water bath. The inner cavity was filled with 37℃ deionized water and cyclically loaded at a frequency of 60 times / min in three pressure ranges: (1) 7-12kPa (50-90mmHg); (2) 10.7-16.0kPa (80-120mmHg); (3) 14.7-20kPa (110-150mmHg). The outer diameters D1 and D2 under different pressures were directly measured using a non-contact laser displacement sensor, and the peripheral compliance C was calculated according to the formula C=[(D2-D1) / D1] / (P2-P1)×100mmHg×100%. The results are shown in Table 1.

[0058] Burst pressure and overall water permeability test: Referring to the relevant clauses of YY / T 0500-2021 on pressure burst strength and overall permeability, an artificial blood vessel compliance, burst strength and overall permeability tester conforming to this standard was used. The sample was cut into a tube with a length of 80 mm, and both ends were sealed to the Luer connector of the tester. The inner cavity was filled with distilled water at 37℃, and the pressure was continuously increased at a pressure increase rate of 10 kPa / s until the sample ruptured. The maximum internal pressure was recorded as the burst pressure. Another sample of the same type was taken and injected with water at a constant pressure of 16 kPa (120 mmHg) for 10 min. The volume of water permeating through the blood vessel wall per unit time was collected. The overall water permeability was expressed as Q = V / (A·t). The results are shown in Table 1.

[0059] Pulsating fatigue and self-healing performance test: Referring to YY / T 0500-2021 regarding the durability requirements of tubular vascular grafts under simulated blood flow circulation, 4×10 7 The cycle was designed to represent the equivalent number of heartbeats used in vivo over one year. The sample was connected to a closed circulating water circuit, using a 37°C phosphate buffer solution as the medium, and continuously pulsed for 4.0 × 10⁻⁶ times at a frequency of 120 beats / min and a pressure of 80-160 mmHg. 7 Before each cycle, a shallow scratch with a length of 2 mm and a depth of about half the thickness of the hydrogel was lightly made along the axial direction on the inner wall of each sample with a scalpel. The initial leakage pressure and radial compliance 80-120 were recorded after the scratch. After the cycle, the leakage pressure recovery rate and compliance decrease rate were calculated. The results are shown in Table 1.

[0060] In vitro cytotoxicity test: Following GB / T 16886.5-2017, the artificial blood vessel was cut into small pieces, with a sample surface area to extraction medium volume ratio of 3 cm². 2 Add / mL of minimum essential culture medium (containing 10% fetal bovine serum) and extract at 37℃ and 5% CO2 for 24h. The resulting extract is sterilely filtered through a 0.22μm filter membrane and used for cell treatment. Logarithmic growth phase L929 cells are seeded into 96-well plates at 1×10⁶ cells per well. 4 Cells were cultured for 24 hours, and then extraction solution, negative control (medical grade polypropylene), and positive control (0.64% phenol solution) were added. After another 24 hours of culture, the relative cell viability of each group was determined by the MTT assay. The results are shown in Table 1.

[0061] Table 1 Performance Test Results Data Analysis: As can be seen from the data in Examples 1-3 of Table 1, the biomimetic artificial blood vessels based on self-healing hydrogel prepared in this invention comprehensively meet multiple requirements, including high water content, interfacial bonding strength, radial compliance, burst pressure, anti-leakage ability, and cell compatibility. The water absorption rate of the three examples is at a relatively high but controllable level, ensuring that the hydrogel maintains its softness without excessive expansion leading to luminal narrowing. Radial compliance gradually decreases with increasing pressure in different pressure ranges, reflecting a nonlinear elastic response of the vessel wall similar to that of natural blood vessels, which is beneficial for buffering pulsating pressure. The burst pressure is generally within a safe range significantly higher than physiological blood pressure, and maintains a high pressure margin without significantly sacrificing compliance. The overall water permeability remains at an extremely low level, indicating that the polytetrafluoroethylene body and the hydrogel liner form a continuous and dense composite wall layer, inhibiting leakage without relying on preoperative precoagulation. After long-term pulsating load, the leakage pressure recovery rate remains high. The relatively high level of compliance with a small decrease indicates that the reversible covalent bonds between sodium alginate and dihydrazine adipic acid-modified gelatin, the borate ester bonds between carboxymethyl cellulose grafted phenylboronic acid and polyvinyl alcohol, and the multi-point hydrogen bonds synergistically form a multi-network that can continuously dissipate energy and achieve structural reconstruction at the crack tip, thereby inhibiting fatigue-accumulated damage. At the same time, the relative cell survival rate of all three sets of examples was significantly higher than the evaluation threshold, indicating that after introducing active groups such as dopamine, dihydrazine adipic acid, and phenylboronic acid, good cell compatibility can still be maintained while ensuring self-healing and mechanical properties through sufficient dialysis and post-processing. This provides a basis for its long-term implantation as a liner material for small-diameter artificial blood vessels.

[0062] As can be seen from the data in Table 1 for Example 2 and Comparative Example 1, when the dopamine coating is removed and the poly(hydroxyethyl methacrylate-glycidyl methacrylate) layer is directly bonded to the self-healing hydrogel, the overall water absorption rate and static radial compliance do not change significantly. However, the interfacial shear strength is significantly reduced, the leakage pressure recovery rate after pulsating fatigue decreases significantly, and the degree of compliance decay is significantly aggravated. The main reason is that the hydrophilic hydroxyl groups provided by hydroxyethyl methacrylate and the epoxy groups of glycidyl methacrylate can form a certain number of covalent bonds with the hydroxyl and amino groups in the hydrogel. However, the lack of catechol-metal coordination, catechol-aromatic π-π stacking, and hydrogen bond network provided by dopamine results in a relatively simple interfacial energy dissipation mechanism, making it difficult to passivate interfacial microcracks in a timely manner. Under long-term cyclic loading, irreversible debonding bands easily form between the polytetrafluoroethylene body and the hydrogel, leading to the gradual accumulation of local leakage channels and compliance mutation zones, which manifests as insufficient leakage pressure recovery and a high compliance decay rate. It is evident that the dopamine coating and the hydrophilic copolymer graft layer have a significant synergistic effect on interfacial adhesion and fatigue self-healing, and it is difficult to achieve the stable interfacial performance required by this invention by retaining either layer alone.

[0063] As can be seen from the data in Table 1 for Example 2 and Comparative Example 2, when the poly(hydroxyethyl methacrylate-glycidyl methacrylate) graft layer is omitted and a dopamine coating is constructed only on the inner wall of polytetrafluoroethylene, the interfacial shear strength and post-fatigue leakage pressure recovery are better than those with only a hydrophilic copolymer layer, but still significantly worse than the example with both modified layers. This is because the polydopamine coating, generated by the self-polymerization of dopamine under oxidative conditions, can form an adhesive underlayer rich in catechols and amines on a low surface energy substrate, establishing strong adhesion with the hydrogel through hydrogen bonds and covalent crosslinking. However, lacking the flexible hydrophilic segments provided by hydroxyethyl methacrylate and the high-density epoxy grafting sites provided by glycidyl methacrylate, the mechanical interlocking and chemical anchoring between the polydopamine layer and the polytetrafluoroethylene substrate remain limited. During repeated radial expansion and contraction, whole-piece peeling or shear slip easily occurs, causing interfacial stress concentration, which in turn affects the long-term stability of compliance and pressure resistance. In the three-layer structure, the hydrophilic copolymer layer enhances interfacial wetting and free radical grafting, while providing a rough skeleton and intercalation points for polydopamine. This allows the covalent linkages of catechol-metal coordination and hydrogel-epoxy to be synergistically distributed in deeper interfacial regions, resulting in significant adhesion and anti-fatigue effects.

[0064] As can be seen from the data in Example 2 and Comparative Example 3 in Table 1, when the carboxymethyl cellulose grafted phenylboronic acid polymer was replaced by ordinary sodium carboxymethyl cellulose and the borate ester dynamic network was removed, the water absorption rate of the self-healing hydrogel increased significantly, the overall water permeability increased, the radial compliance was slightly higher in the low-pressure range, showed an abnormal further increase in the medium-pressure range, and decreased again in the high-pressure range, while the leakage pressure recovery rate and compliance retention both deteriorated significantly. This phenomenon indicates that relying solely on the Schiff base network between sodium alginate oxidized and dihydrazine adipic acid modified gelatin is insufficient to stably constrain hydrogels containing a large number of hydrophilic segments under multiaxial stress. Local water-rich areas are more prone to softening and expansion in the medium-pressure range, resulting in an abnormal increase in compliance. However, in the high-pressure stage, irreversible plastic deformation occurs due to the fatigue accumulation of repeated Schiff base bond breakage and recombination, and the compliance decreases. Without the second set of dynamic crosslinks between phenylboronic acid and polyvinyl alcohol segments, the crack tip lacks sufficient reversible bonds to participate in the coordinated opening and closing. The crack tends to propagate along the locally water-rich softening zone, resulting in insufficient leakage pressure recovery and severe compliance degradation after fatigue cycles. This result demonstrates that the borate ester network not only provides additional stiffness and creep resistance, but more importantly, it complements the Schiff base network in terms of time scale and response mode, thus exhibiting a synergistic effect in overall fatigue resistance and self-healing behavior.

[0065] As can be seen from the data in Example 2 and Comparative Example 4 in Table 1, when all the adipic acid dihydrazine modified gelatin was replaced with unmodified pigskin gelatin, and only the borate ester network composed of carboxymethyl cellulose grafted with phenylboronic acid and polyvinyl alcohol was retained, the water absorption rate of the hydrogel decreased slightly, and the burst pressure and initial compliance even increased slightly in some ranges. However, the overall water permeability increased significantly, and the leakage pressure recovery rate and compliance retention also deteriorated significantly. The mechanism is that the adipic acid dihydrazine modified gelatin, by introducing multiple hydrazine groups on the gelatin backbone, can form a dense Schiff base skeleton with the aldehyde groups of oxidized sodium alginate, providing a flexible support framework for the borate ester network, so that the stress is distributed between the two networks. After removing this modification, the gelatin is only combined with the network in the form of physical entanglement and hydrogen bonding. The borate ester crosslinking provides high stiffness in a short time, but under repeated combined action of damp heat and shear, the local crosslinking points are more likely to relax and undergo dehydration-rehydration cycles, forming penetrating micropores, resulting in an increase in overall water permeability and a decrease in compliance after fatigue. Compared with Comparative Example 3, which only retains the Schiff base network, Comparative Example 4 seems to have an advantage in terms of initial pressure resistance. However, under long-term pulsating loads, neither can maintain a balance between low permeability and high compliance at the same time. This, in turn, confirms that the combination of Schiff base network and borate ester network in multi-scale time response is the key to achieving long-lasting self-healing and fatigue resistance. It is a typical example of the synergistic effect where a single network is insufficient on its own, and the superposition of two networks shows advantages.

[0066] As can be seen from the data in Table 1 for Example 2 and Comparative Example 5, under the premise of keeping the polytetrafluoroethylene surface modification method and the dual dynamic network composition unchanged, after unifying the three hydrogel precursor solutions of the inner, middle and outer layers into a middle layer formulation, the static burst pressure, initial radial compliance and cell compatibility are not significantly different from those of the examples. However, the leakage pressure recovery rate after long-term pulsating cycling is significantly reduced, and the compliance decline rate is increased, indicating that the uniform hydrogel wall layer is significantly less fatigue-resistant than the structure with a radial gradient. The reason is that in the examples, by increasing the proportion of dihydrazine adipic acid modified gelatin and reducing the content of carboxymethyl cellulose grafted phenylboronic acid and calcium chloride in the inner layer, the Schiff base gelatin network dominates the elasticity and biocompatibility of the inner layer. Meanwhile, by increasing the content of carboxymethyl cellulose grafted phenylboronic acid and calcium chloride and relatively reducing the content of dihydrazine adipic acid modified gelatin in the outer layer, the borate ester-calcium alginate-polyvinyl alcohol network dominates the strength and creep resistance of the outer layer, thus forming a gradient distribution of inner softness and outer rigidity. After removing the gradient, the strain experienced by the inner and outer layers under the same pressure tends to be consistent. The blood-near side of the inner cavity cannot dissipate energy through preferential deformation of the softer network, while the outer side lacks a higher degree of cross-linking to inhibit creep. This results in cracks propagating straight through the wall thickness, manifested as compliance decaying with accelerated cycling and insufficient recovery from leakage pressure. These results demonstrate that superimposing a radial composition gradient on top of an already introduced dual dynamic network can achieve a fatigue resistance enhancement effect far exceeding that of a simple superposition effect.

[0067] As can be seen from the data in Example 2 and Comparative Example 6 in Table 1, when the radial gradient direction in the examples is reversed, and the high carboxymethyl cellulose grafted phenylboronic acid and high calcium chloride formulation originally used for the outer layer is moved to the inner layer, while the soft formulation originally biased towards dihydrazine adipic acid modified gelatin is moved to the outer layer, the radial compliance of the low and medium pressure zones decreases overall, exhibiting a significantly harder characteristic on the inner lumen side. After long-term pulsating circulation, the rate of compliance decline increases and the leakage pressure recovery is significantly reduced. The main reason is that once the inner layer is dominated by a highly cross-linked borate ester-calcium alginate-polyvinyl alcohol network, the radial expansion caused by blood flow pulsation is more borne by the softer network of the outer layer, leading to the shift of the strain peak to the outside and an increase in shear stress at the interface between the inner layer and polytetrafluoroethylene body. At the same time, excessively high stiffness of the inner layer will significantly amplify the compliance mismatch with natural blood vessels, causing stress concentration near the anastomosis and increasing the risk of intimal hyperplasia and microcrack initiation. Because the outer layer is relatively soft, it is difficult to provide timely backing constraints for the inner layer cracks. Cracks tend to propagate rapidly along the interface, resulting in insufficient recovery of leakage pressure after fatigue, increased overall water permeability, and rapid decline in compliance. Compared to Comparative Example 5 with the gradient removed, the reverse gradient amplifies the adverse effects, further highlighting the synergistic effect of the combination of the inner-soft, outer-rigid gradient design and the dual dynamic network in compliance matching and crack self-healing.

[0068] from Figure 1 It can be seen that oxidized sodium alginate at 1605 cm⁻¹ -1 and 1415cm -1 The typical carboxylate COO is exhibited at this location. - The asymmetric and symmetric stretching vibration peaks of hydrazine adipic acid modified gelatin are at 1650 cm⁻¹. -1 and 1550cm -1 The amide I and II characteristic absorptions are obvious at 1602 cm⁻¹; the carboxymethyl cellulose-grafted phenylboronic acid polymer exhibits significant absorption at 1602 cm⁻¹. -1 and 1510cm -1 Aromatic ring skeletal vibrations were observed at 1335 and 1210 cm⁻¹. -1 The spectral composition of the self-healing hydrogel biomimetic artificial blood vessel hydrogel ring combines the above characteristic peaks, with the 1650 cm⁻¹ peak being particularly prominent. -1 Amide I / C=N composite band and 1605, 1415cm -1 The absorption of carboxylate salts occurs simultaneously, at 1090-1030 cm⁻¹. -1 The significant broadening of the COC / CO band indicates the successful synthesis of oxidized sodium alginate, hydrazine adipic acid-modified gelatin, and carboxymethyl cellulose grafted with phenylboronic acid, as well as the successful construction of a self-healing hydrogel structure containing a Schiff base and a borate ester dual dynamic crosslinking network.

[0069] Those skilled in the art should understand that the discussion of any of the above embodiments is merely exemplary and is not intended to imply that the scope of the invention is limited to these examples; within the framework of the invention, the technical features of the above embodiments or different embodiments can also be combined, the steps can be implemented in any order, and there are many other variations of the different aspects of the invention as described above, which are not provided in detail for the sake of brevity.

Claims

1. A method for preparing biomimetic artificial blood vessels based on self-healing hydrogels, characterized in that, Includes the following steps: (1) The polytetrafluoroethylene tube is cleaned and dried, and then subjected to plasma surface treatment in an oxygen atmosphere to obtain a pretreated polytetrafluoroethylene tube. (2) The pretreated polytetrafluoroethylene tube is immersed in a monomer solution containing hydroxyethyl methacrylate and glycidyl methacrylate, and polymerized in the presence of a free radical initiator to form a hydrophilic polymer layer on the tube wall; then, a dopamine solution is injected into the tube cavity, and dopamine is self-polymerized under alkaline and oxidative conditions to form a polydopamine layer on the surface of the polymer layer, thus obtaining a surface-modified polytetrafluoroethylene tube; (3) Sodium alginate was oxidized by sodium periodate, sodium alginate was oxidized, dihydrazine adipic acid was modified gelatin by condensation reaction of dihydrazine adipic acid and pigskin gelatin, and carboxymethyl cellulose grafted phenylboronic acid polymer was obtained by condensation reaction of 4-aminophenylboronic acid and sodium carboxymethyl cellulose. (4) Prepare three types of hydrogel precursor solutions: inner layer, middle layer and outer layer. Each hydrogel precursor solution includes sodium alginate oxide, dihydrazine adipic acid modified gelatin, carboxymethyl cellulose grafted phenylboronic acid polymer, polyvinyl alcohol and calcium chloride. The contents of sodium alginate oxide, carboxymethyl cellulose grafted phenylboronic acid polymer, polyvinyl alcohol and calcium chloride gradually increase from the inner layer to the outer layer, while the contents of dihydrazine adipic acid modified gelatin gradually decrease. (5) The outer layer hydrogel precursor solution is injected into the lumen of the surface-modified polytetrafluoroethylene tube. After partial gelation under the condition of tube rotation, a portion of the precursor solution that is still in a sol state is extracted from the center of the tube to form a hollow channel. Then, the middle layer hydrogel precursor solution is injected into the hollow channel. Partial gelation is continued under the condition of rotation, and a portion of the precursor solution is extracted. Then, the inner layer hydrogel precursor solution is injected into the obtained channel. Gelation is completed under the condition of rotation, so that the inner wall of the polytetrafluoroethylene tube forms a self-healing hydrogel blood vessel wall with a composition gradient from the inside to the outside, and a polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessel is obtained. (6) Post-processing of artificial blood vessels: The polytetrafluoroethylene / self-healing hydrogel composite artificial blood vessels were extracted in a buffer salt solution, then disinfected in an ethanol aqueous solution, rinsed with a buffer salt solution and stored under sterile conditions to obtain biomimetic artificial blood vessels based on self-healing hydrogel.

2. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (1), the inner diameter of the polytetrafluoroethylene small-diameter tube is 4mm, the outer diameter is 6mm, and the length is 80mm.

3. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (1), the plasma surface treatment is performed by treating the surface with 100W power for 10 minutes in a radio frequency oxygen plasma surface treatment machine.

4. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (2), the monomer solution consists of 32-48g hydroxyethyl methacrylate, 8-12g glycidyl methacrylate, 0.8-1.2g azobisisobutyronitrile and 40-60mL anhydrous ethanol.

5. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (2), the dopamine solution is composed of 1.6-2.4g of dopamine hydrochloride and 160-240mL of Tris-HCl buffer solution; the concentration of the Tris-HCl buffer solution is 40-60mmol / L and the pH is 8.

6. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (3), the mass ratio of sodium periodate to sodium alginate is 3.2-4.8:8-12; the mass ratio of dihydrazine adipic acid to porcine gelatin is 2.4-3.6:10; and the mass ratio of 4-aminophenylboronic acid to sodium carboxymethyl cellulose is 2.4-3.6:

10.

7. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (4), the inner hydrogel precursor solution is prepared by dissolving 1.6-2.4g sodium alginate oxide, 3.2-4.8g dihydrazine adipic acid modified gelatin, 0.8-1.2g carboxymethyl cellulose grafted phenylboronic acid polymer and 0.8-1.2g polyvinyl alcohol in 40mL of phosphate buffer solution with pH 7.4 at 50℃. After cooling to 37℃, 0.8-1.2g dihydrazine adipic acid is added and stirred evenly. Finally, 0.8-1.2g calcium chloride is added and stirred until dissolved. The intermediate hydrogel precursor solution was prepared by dissolving 2.4-3.6 g sodium alginate oxide, 2.4-3.6 g dihydrazine adipic acid modified gelatin, 1.6-2.4 g carboxymethyl cellulose grafted phenylboronic acid polymer, and 1.6-2.4 g polyvinyl alcohol in 40 mL of phosphate buffer solution with pH 7.4 at 50 °C with stirring. After cooling to 37 °C, 0.8-1.2 g dihydrazine adipic acid was added and stirred until homogeneous. Finally, 1.6-2.4 g calcium chloride was added and stirred until dissolved. The precursor solution for the outer hydrogel was prepared by dissolving 3.2-4.8 g of sodium alginate oxide, 1.6-2.4 g of dihydrazine adipic acid modified gelatin, 2.4-3.6 g of carboxymethyl cellulose grafted with phenylboronic acid polymer, and 2.4-3.6 g of polyvinyl alcohol in 40 mL of phosphate buffer solution with pH 7.4 at 50 °C with stirring. After cooling to 37 °C, 0.8-1.2 g of dihydrazine adipic acid was added and stirred until homogeneous. Finally, 2.4-3.6 g of calcium chloride was added and stirred until dissolved.

8. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (4), the weight-average molecular weight of polyvinyl alcohol is 95,000 and the degree of alcoholysis is 99%.

9. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (5), the polytetrafluoroethylene tube is rotated at 30 rpm at 25°C for 5 minutes and then left to stand for 10 minutes during the injection of the outer layer hydrogel precursor solution, the middle layer hydrogel precursor solution and the inner layer hydrogel precursor solution. The volume of the precursor liquid extracted from the center of the tube is one-third of the volume of the injected precursor liquid. Finally, after the inner layer hydrogel precursor solution is injected, the tube is left to stand at 37°C for 2 hours to complete the gelation.

10. The method for preparing biomimetic artificial blood vessels based on self-healing hydrogels according to claim 1, characterized in that, In step (6), the buffer salt solution is a phosphate buffer salt solution with a pH of 7.4.

Images