Shape memory degradable metal-based large-vessel stent and preparation method thereof

Abstract

The application discloses a shape memory degradable metal-based large blood vessel stent and a preparation method thereof, and belongs to the field of medical devices. The stent comprises a degradable iron-manganese-silicon-copper-niobium series shape memory alloy base, a degradable polymer isolation coating film wrapped on the outer surface of the base, and a composite drug coating loaded on the outer surface of the coating film. The alloy base adopts a wide hinge low strain structure, matches a stress-induced martensite stable-body temperature triggered reverse phase change recovery phase change mechanism, solves the compatibility problem of large-diameter stent conveying stability and self-expansion, releases copper ions during degradation to promote vascular endothelial repair. The composite drug coating comprises two kinds of ultrasonic response drug-loaded microcapsules with different particle sizes, which can be triggered to release drugs in stages through in-vitro low-intensity focused ultrasound. The necking micropores on the surface of the base and the micro-nano concave structures on the surface of the coating film constitute a two-stage embedding system, and the interface bonding reliability is improved. The application takes into account the functions of mechanical support, anatomical adaptive fitting, intraluminal isolation, endothelial repair promotion and controllable drug release.

Description

Technical Field

[0001] This invention belongs to the field of medical device technology, specifically relating to a shape memory biodegradable metal-based large blood vessel stent and its preparation method, which is suitable for endovascular repair treatment of aortic dissection. Background Technology

[0002] Aortic dissection is a dangerous cardiovascular emergency with a high mortality rate. Its pathological feature is a rupture of the aortic intima, allowing high-pressure blood flow to rush into the middle layer of the vessel wall, creating a true and false lumen. Without timely intervention, it can lead to fatal complications such as aortic rupture and organ ischemia. Percutaneous intervention with a endovascular stent graft to seal the dissection and restore blood flow to the true lumen is currently the mainstream minimally invasive treatment.

[0003] Currently, most large vessel stents used in clinical practice use permanent metallic materials such as nickel-titanium alloys as the supporting framework. These stents remain in the patient's body for life after implantation, presenting several long-term challenges: First, long-term foreign body irritation and tissue compression complications. The permanent metallic framework continuously irritates the vessel wall, potentially inducing chronic inflammation, excessive intimal hyperplasia, and in-stent restenosis. Simultaneously, the long-term compression of adjacent tissues such as the trachea, esophagus, and peripheral nerves by large-diameter stents can lead to complications such as chronic cough, dysphagia, and chest pain, significantly impacting the long-term quality of life of younger patients. Furthermore, imaging artifacts produced by the permanent metallic framework can interfere with postoperative CT and MRI follow-up examinations. Second, existing biodegradable stents have insufficient mechanical properties and functionality. Polymer-based biodegradable stents have limited radial support force, making it difficult to withstand the impact of high-pressure blood flow within the aorta and failing to meet the mechanical requirements of large vessels. Iron-based and magnesium-based biodegradable stents offer superior mechanical properties, but pure iron degrades too slowly, and existing iron-based biodegradable stents generally lack shape memory properties, relying heavily on balloon expansion for release. This makes it difficult to conform to the complex anatomical structure of the aorta, resulting in poor sealing of aortic dissection ruptures. Third, existing iron-based shape memory stents suffer from unreasonable phase transition designs and poor delivery stability. If the phase transition temperature is not set properly, spontaneous reverse phase transitions can easily occur within the room-temperature sheath after stent compression, leading to increased delivery resistance and release failure. This is a core engineering bottleneck restricting the clinical application of iron-based self-expanding stents. Fourth, they lack the function of actively promoting endothelial repair. Slow endothelialization after stent implantation is a significant cause of thrombosis and restenosis. Most existing biodegradable stents do not possess the function of actively promoting endothelial repair, making it difficult to synchronize degradation with the vascular repair process. Fifth, drug release cannot match the vascular repair process. Vascular repair is a phased and dynamic process: in the acute postoperative period, inflammation needs to be controlled and smooth muscle layer repair promoted to stabilize the vessel wall; in the mid-to-late postoperative period, excessive intimal proliferation needs to be inhibited to prevent restenosis. Existing stent drug coatings are mostly burst-release or constant-rate release modes, failing to achieve time-sequential drug delivery. Sixth, the safety of manganese ion release from iron-manganese-based biodegradable alloys lacks systematic control. Long-term exposure to high manganese content poses a neurotoxicity risk, a key concern for regulatory review.

[0004] In recent years, numerous studies have been conducted in this field on stents for large blood vessels. Patent CN11331992B discloses a combined aortic endovascular stent graft, which can reduce the risk of endoleak; however, its nickel-titanium alloy framework is a permanent implant, and long-term retention can easily induce chronic inflammation and intimal hyperplasia. Patent CN118000985A discloses a biodegradable metallic vascular stent, which optimizes the overall performance of the mesh structure; however, the material lacks shape memory effect and cannot achieve self-expansion and adhesion. Patent CN112126075A discloses a shape memory polymer biodegradable vascular stent, but it is a stent for small blood vessels in the lower extremities, and its mechanical properties cannot adapt to the aortic environment; furthermore, shape memory requires external stimulation to trigger, and body temperature-triggered self-expansion is not achieved.

[0005] In the field of iron-based shape memory alloys, Fe-Mn-Si alloys have been proven to possess shape memory effects and mechanical properties superior to pure iron. Kajiwara et al. demonstrated that generating NbC precipitates can significantly improve the shape memory effect of Fe-Mn-Si alloys, clarifying the strengthening effect of microalloying elements on phase transformation performance. Prokoshkin et al. conducted research on biodegradable Fe-30Mn-5Si (wt%) shape memory alloys, confirming the regulation of the microstructure and functional fatigue properties of thermomechanically treated alloys, providing experimental basis for the process design of biodegradable iron-based shape memory alloys. A review by Peng H et al. pointed out that stacking fault energy regulation and grain size optimization are core factors in achieving large recoverable strain in polycrystalline Fe-Mn-Si alloys, systematically summarizing the key influencing parameters of the shape memory performance of this system, and further clarifying the regulation logic of the shape memory performance of this system. From the perspective of its mechanism of action, its shape memory mechanism is based on the reversible phase transformation of γ-austenite and ε-martensite at low stacking fault energies. Silicon is the core element for reducing stacking fault energy and realizing reversible phase transformation, and the silicon content in practical alloy systems is usually not less than 4 wt%. However, this alloy system undergoes a semi-thermoelastic martensitic phase transformation, and the driving force of a simple thermally induced phase transformation is weak and the recovery force is limited. It is necessary to achieve the shape memory function by combining stress-induced martensite with heating-induced reverse phase transformation. Existing research has not yet addressed the large deformation requirements of large vascular stents by matching the thin-walled hinge structure with the alloy's phase transformation characteristics, nor has it solved the compatibility problem between room temperature delivery stability and body temperature self-expansion. In recent years, copper-containing Fe-Mn-Si alloys have been preliminarily studied as biodegradable implantable materials. Copper can regulate the corrosion rate and endow the material with certain biological functions. However, the copper content in existing studies is generally low, which has limited the improvement on degradation rate and endothelial-promoting effect. Moreover, this type of material has not yet been systematically applied to aortic stents, and functions such as membrane-covered isolation, ultrasound-guided drug delivery, interface enhancement, and manganese ion safety management have not been integrated.

[0006] In summary, no existing technology has yet produced a biodegradable large vessel stent that can simultaneously meet the requirements of aortic hypertension mechanical support, room temperature stable delivery, body temperature self-expanding fit, effective intraluminal isolation, active promotion of endothelial repair, time-sequential drug release, and controllable degradation safety. Summary of the Invention

[0007] To address the shortcomings of existing technologies, the present invention aims to provide a biodegradable iron-based shape memory alloy large vessel stent. This stent addresses the core bottlenecks of limited recoverable strain and unstable room-temperature delivery in iron-based alloys through a synergistic wide-hinge low-strain structure and phase transition temperature design. It promotes vascular repair and manages manganese ion safety through controlled degradation and ion release; enhances staged drug release through an ultrasound-responsive microcapsule system; and improves the bonding reliability of the overlay and coating through a hierarchical interface interlocking structure, thereby enhancing the long-term efficacy and safety of endovascular treatment for aortic dissection.

[0008] To achieve the above-mentioned objectives, the present invention provides a shape memory biodegradable metal-based large vessel stent and its preparation method, for endovascular repair of aortic dissection, comprising a biodegradable metal stent matrix, a biodegradable coating layer covering the outer surface of the stent matrix, and a composite drug coating loaded on the outer surface of the stent matrix.

[0009] Furthermore, the biodegradable metal scaffold matrix is ​​prepared from an iron-manganese-silicon-copper-niobium biodegradable shape memory alloy. By mass percentage, the composition includes: Mn 28%-33%, Si 4%-6%, Cu 1%-3%, Nb 0.1%-0.5%, C ≤0.02%, with the balance being Fe and unavoidable impurities totaling no more than 0.5 wt%. Manganese is an austenite-stabilizing element used to regulate the martensitic phase transformation temperature of the alloy; silicon is used to reduce the stacking fault energy of the alloy and is a key element for achieving the reversible thermoelastic phase transformation from γ-austenite to ε-martensite; copper is used to refine grains, regulate corrosion potential, and release copper ions during alloy degradation to promote endothelial repair; niobium is used to form microalloyed precipitates to refine grains and improve thermomechanical stability; carbon is a strictly controlled harmful impurity with a maximum content of 0.015%, used to avoid carbides damaging the reversibility of the phase transformation and the uniformity of degradation.

[0010] Furthermore, the martensitic transformation initiation temperature (Ms) of the biodegradable shape memory alloy is 0℃-10℃, the reverse transformation initiation temperature (As) is 28℃-32℃, and the austenitic transformation end temperature (A) is... fThe phase transformation range is 35℃-37℃. This phase transformation range design takes into account both delivery stability and in vivo recovery performance: at room temperature (22℃-25℃), the alloy maintains a stable austenitic state, and the temperature is lower than the initiation temperature of the reverse phase transformation (As). When the stent is compressed by external force, stress induces the transformation of γ austenite to ε martensite, achieving large plastic deformation. The stent can be compressed and stored in delivery sheaths of 20F and below. After compression and storage, the room temperature is lower than As, and the stress-induced martensite phase is in a thermodynamically stable state with no spontaneous reverse phase transformation tendency, ensuring a delivery operation window of more than 4 hours without the risk of sheath jamming or premature expansion. After the stent is delivered into the aortic lumen and released, under human body temperature conditions of 37℃ (higher than As), f The self-expanding stent rapidly undergoes a reverse phase transformation from ε-martensite to γ-austenite (due to temperature), automatically restoring its preset expansion shape. Combined with the elastic rebound of the wide hinge, it achieves self-expansion and release without balloon assistance. This design eliminates mechanical damage to the aortic intima caused by balloon dilation, reducing intraoperative risks such as retrograde dissection and vascular rupture. The self-expanding stent can adaptively match the physiological curvature of the aortic wall and changes in vessel diameter, achieving a more precise closure of intimal tears.

[0011] Furthermore, after undergoing vacuum high-temperature shaping at 350℃-400℃ and 10-15 rounds of "pre-deformation-heat recovery" thermomechanical cycle training, the stent substrate possesses a stable stress-triggered shape memory effect. In in vitro simulated body fluid (PBS, pH 7.4, 37℃), the stent substrate exhibits long-term radial support maintenance capability, enabling it to maintain sufficient luminal support in the high-pressure blood flow environment of the aorta.

[0012] Furthermore, the stent substrate is a hollow tubular structure, integrally cut and formed from several sets of axially spaced corrugated support rings and staggered spirally arranged axial connecting ribs; the crests and troughs of adjacent corrugated support rings are correspondingly arranged axially, and the axial connecting ribs connect the crests and troughs of adjacent support rings; the stent has a diameter of 20mm-46mm after full expansion, and a delivery outer diameter ≤ 20F under compression. The corrugated support rings can provide sufficient radial support force to resist the impact of high-pressure blood flow during aortic systole; the axial connecting ribs are arranged in a staggered spiral manner to improve the axial compliance of the stent, facilitating the stent's passage through tortuous anatomical segments such as the aortic arch. The width of the elastic deformation zone of the wide hinge is 3-5 times the width of the adjacent support rib. The hinge corner adopts a large arc transition, which can fully disperse the local strain concentrated at the corner during the compression of the support along the width direction by increasing the deformation bearing area. When the support is compressed to the minimum conveying outer diameter, the peak equivalent strain in the hinge area is controlled within the recoverable range of the alloy's reversible phase transformation, ensuring the shape memory recovery efficiency and cycle stability under large deformation. The limited martensite conversion rate in the low strain region is one of the core reasons why the overall recovery rate of the support is lower than the material's peak recovery rate.

[0013] Furthermore, the scaffold substrate surface is coated with a 100 nm-300 nm thick anodic passivation layer, which can regulate the initial corrosion rate and prevent premature release of copper ions. Copper ion release exhibits a kinetic characteristic of initial release followed by a peak and then gradually stabilizing, with the long-term dynamic average concentration in the local microenvironment of the scaffold remaining within an effective range. This concentration range has been validated by cell biology for both safety and biological effectiveness: too low a concentration results in insufficient endothelial induction, while too high a concentration easily induces cytotoxicity. Initial release facilitates early endothelial cell recruitment, while subsequent stable release sustainably promotes the endothelialization process. Regarding the safety control of manganese ions: the anodic passivation layer can uniformly regulate the alloy corrosion rate, avoiding sudden release of manganese ions caused by rapid local corrosion, ensuring no risk of systemic neurotoxicity accumulation. Copper ions and growth factors released from the coating can produce a synergistic endothelial-promoting effect: copper ions promote the expression of endogenous growth factors by upregulating the hypoxia-inducible factor pathway, jointly accelerating endothelial cell proliferation and migration with exogenous growth factors.

[0014] Furthermore, the biodegradable polymer insulating membrane is coated on the outer surface of the stent matrix. The material is selected from polylactic acid-caprolactone copolymer (PLCL) or polycaprolactone (PCL), with a thickness of 30 μm-100 μm. The membrane adheres directly to the inner wall of the aorta as the stent expands, completely covering the dissection tear, blocking high-pressure blood flow from the true lumen into the false lumen through the tear, achieving effective intraluminal isolation, and promoting thrombosis of the false lumen.

[0015] Furthermore, the composite drug coating is loaded onto the outer surface of the biodegradable membrane, comprising a biodegradable polymer carrier and two hollow, ultrasound-responsive drug-loaded microcapsules with different particle sizes. The biodegradable polymer carrier is polylactic-co-glycolic acid (PLGA), with a lactic acid:glycolic acid molar ratio of 50:50-85:15 and a molecular weight of 10 kDa-50 kDa. It exhibits good biocompatibility and can serve as a microcapsule dispersion matrix and sustained-release barrier, maintaining low levels of drug leakage in the absence of ultrasound stimulation. The hollow, ultrasound-responsive drug-loaded microcapsules are polymer microcapsule structures containing an inert gas core. Their drug release mechanism is based on the ultrasonic cavitation resonance effect. Under the action of ultrasound at a specific frequency, the gas core inside the microcapsule resonates, expands, and ruptures, causing the capsule wall to break down and release the drug. According to the negative correlation between microcapsule resonance frequency and particle size, a specific response to the ultrasonic frequency can be achieved through particle size differences. Among them, small-particle-size microcapsules (0.5 μm-2 μm) encapsulate anti-inflammatory drugs and smooth muscle cell repair active ingredients, with an inherent resonant frequency corresponding to the 0.8 MHz-2.0 MHz frequency band. Distributed in the shallow layer of the coating, they facilitate rapid drug release after ultrasound triggering. Large-particle-size microcapsules (5 μm-15 μm) encapsulate anti-proliferative drugs, with an inherent resonant frequency corresponding to the 200 kHz-600 kHz frequency band. Distributed in the deep layer of the coating, they have a higher drug loading capacity, meeting the needs of medium- and long-term drug release. The drug coating is entirely attached to the outer surface of a biodegradable polymer insulating membrane. Both types of drug-loaded microcapsules are completely embedded within the polymer carrier of the coating, with a continuous polymer carrier layer remaining between the bottom of the microcapsule and the membrane, eliminating the risk of membrane penetration. The two types of microcapsules have different response thresholds, preventing unexpected premature drug release. Considering the complex acoustic environment of in vivo tissues, in clinical applications, medical imaging localization can be combined with adjustments to parameters such as ultrasound sound pressure and pulse duty cycle to further optimize trigger selectivity and achieve staged drug delivery. The coating preparation process employs a low-temperature, ambient-pressure method throughout to prevent the expansion and rupture of the microcapsule gas cores, ensuring stable ultrasonic response. Regarding ultrasonic drug release testing: The cross-response data for ultrasonic drug release in this application are based on an ideal in vitro testing environment. In actual clinical applications, human tissues will exhibit certain attenuation and scattering effects on ultrasound. During clinical use, the ultrasonic output parameters can be appropriately adjusted according to the target depth and tissue type to achieve precise targeted drug release.

[0016] Furthermore, the anti-inflammatory drug is selected from at least one of dexamethasone and prednisolone; the smooth muscle repair active ingredient is selected from at least one of vascular endothelial growth factor and fibroblast growth factor; and the anti-proliferative drug is selected from at least one of rapamycin, paclitaxel, and zotamoxetine. The drug release sequence and pharmacological logic are as follows: In the acute postoperative period, 0.8 MHz-2.0 MHz ultrasound is applied to trigger the release of small-diameter microcapsules, glucocorticoids rapidly control acute inflammation, and growth factors promote smooth muscle layer repair; after the peak of inflammation, growth factors and copper ions synergistically and continuously promote endothelialization, with their effects staggered and without significant pharmacological antagonism. In the mid-to-late postoperative period, 200 kHz-600 kHz low-frequency ultrasound is applied to trigger the release of anti-proliferative drugs from large-diameter microcapsules. At this time, the endothelium has initially covered the endothelium, and the drug mainly acts on the smooth muscle cells of the vascular media, inhibiting excessive intimal proliferation, with limited impact on the endothelial barrier.

[0017] Furthermore, the 0.8 MHz-2.0 MHz frequency band ultrasound has a tissue penetration depth of 3-5 cm, effectively reaching the descending thoracic aorta; the 200 kHz-600 kHz low-frequency ultrasound can penetrate to a depth of 5-10 cm, covering the entire aortic segment. Clinically, drug release can be initiated on demand by applying ultrasound of different frequencies non-invasively outside the body, without the need for secondary interventional surgery. The timing and dosage of drug release can be dynamically adjusted based on the patient's imaging follow-up results, improving the individualization and precision of treatment.

[0018] Furthermore, a microporous structured treatment layer is formed on the surface of the scaffold substrate. The micropores are inverted conical structures with necked openings, an opening diameter of 0.5 μm-1 μm, a bottom diameter of 1 μm-5 μm, and a porosity of 10%-30%. The scaffold is prepared using a two-step method combining femtosecond laser direct writing ablation and anodic electrochemical etching for bottom widening: the first step uses a femtosecond laser to ablate the substrate surface to form conical micro-pitted pores (large opening and small bottom); the second step performs anodic electrochemical etching, utilizing the tip current effect of the pit structure, where the bottom curvature radius is small and the current density is high, resulting in a faster metal anodic dissolution rate than the opening, thus selectively widening the bottom of the pore, ultimately forming a necked inverted conical structure with a narrow opening and a wide bottom. During the hot pressing or electrospinning process, the biodegradable polymer insulating coating penetrates into the micropores, and after solidification, forms a mechanically interlocking structure, improving the interfacial bonding strength between the coating and the metal substrate and preventing coating slippage and detachment during scaffold expansion.

[0019] Furthermore, the outer surface of the biodegradable polymer barrier membrane is plasma-etched to form a micro-nano pit structure with an opening size of 0.3 μm-0.8 μm. The polymer carrier of the composite drug coating can be embedded in the pit to form a secondary mechanical interlock, which enhances the bonding strength between the drug coating and the membrane. The drug microcapsules, because their particle size is larger than the pit opening size, are effectively blocked outside the pit, preventing the microcapsules from getting stuck inside the structure and causing the ultrasound response to fail, thus ensuring stable drug loading. At the same time, it prevents the detached microcapsules from migrating with the blood flow and inducing distal vascular embolism, thereby improving the safety of the operation.

[0020] Furthermore, the present invention also provides a method for preparing the above-mentioned support, comprising the following steps: (1) weighing high-purity iron, manganese, silicon, copper and niobium raw materials according to the component ratio, and preparing alloy ingots by vacuum induction melting process under argon protection; (2) homogenizing annealing the alloy ingots at 900℃×10 h-14 h, hot extruding them at 800℃-900℃, and then performing multiple hot drawing at 200℃-400℃ and intermediate annealing at 700℃×1 h to produce alloy tubes with a set wall thickness and outer diameter; (3) processing a preset hollow grid structure on the alloy tubes using femtosecond laser cutting process, wherein the hollow structure includes a wide hinge elastic deformation zone, and obtaining a support blank after deburring and electrochemical polishing; (4) placing the support blank in a vacuum heat treatment furnace and high-temperature shaping at 350℃-400℃ for 15 min-30 min. min, and then 10-15 rounds of "room temperature pre-compression - 37℃ recovery" thermomechanical cycle training to give the stent a stable temperature-triggered shape memory effect; (5) a two-step method of femtosecond laser direct writing ablation combined with anodic electrochemical etching to expand the bottom is used to prepare a microporous structure layer with necked orifice; after cleaning and drying, anodizing passivation treatment is performed, and then a biodegradable polymer isolation coating is prepared on the outer surface of the stent by electrospinning or hot pressing composite process; (6) using S / O / W emulsification-solvent evaporation method or low temperature spray drying method, with ethyl acetate and dichloromethane mixture as organic phase, two kinds of hollow drug-loaded microcapsules with different particle sizes are prepared respectively; among them, the protein drug freeze-dried powder contains 1%-3% trehalose and mannitol composite protectant, and the whole process is operated at low temperature of 0℃-8℃; (7) using a two-step ultrasonic atomization spraying process, first spraying the bottom coating solution containing the second microcapsule, vacuum drying and then spraying the shallow coating solution containing the first microcapsule, and after sufficient vacuum drying to remove residual solvent, the finished stent is obtained.

[0021] Compared with the prior art, the present invention has the following beneficial effects: (1) This invention controls the material strain during compression within the recoverable range of the alloy through the synergistic design of wide hinge low strain structure and phase change temperature regulation, effectively overcoming the problem of limited intrinsic recoverable strain of iron-manganese-silicon alloys, realizing anatomical adaptive fitting of large diameter stents, and reducing dependence on balloon expansion.

[0022] (2) The present invention uses a biodegradable shape memory alloy based on iron, manganese, silicon, copper and niobium, which can gradually degrade within 18-36 months while providing sufficient radial support. The release of copper and manganese ions is controlled synchronously through the anodic oxide layer, so as to achieve the synergistic promotion of endothelial repair by copper ions while ensuring the biosafety of long-term exposure to manganese ions.

[0023] (3) Based on the differences in the ultrasound response characteristics of microcapsules of different particle sizes, this invention combines in vitro low-intensity focused ultrasound to non-invasively enhance the release of corresponding drugs at different stages after surgery, matching the time-sequential needs of vascular repair; the drug is still continuously released at the baseline rate in the absence of ultrasound, ensuring the continuity of treatment.

[0024] (4) The present invention adopts a two-level hierarchical interface interlocking design: the substrate is necked and the micropores are anchored to isolate the coating, and the microstructure on the surface of the coating is anchored to the drug coating. This not only solves the common problem of weak metal-polymer interface bonding in the industry, but also prevents the microcapsules from falling in through the necking structure, avoids the failure of the ultrasonic response function, and improves the reliability of the coating and the coating during the delivery and expansion process. Attached Figure Description

[0025] To more clearly illustrate the technical solutions of the embodiments of the present invention, the accompanying drawings used in the embodiments will be briefly introduced below. It should be understood that the following drawings only show some embodiments of the present invention and should not be regarded as a limitation of the scope. For those skilled in the art, other related drawings can be obtained based on these drawings without creative effort.

[0026] Figure 1 The present invention provides a process flow diagram for the fabrication of a biodegradable large blood vessel stent.

[0027] Figure 2 This is a schematic diagram of the overall structure of the shape memory biodegradable metal-based large blood vessel stent of the present invention, including the coating and composite drug coating.

[0028] Figure 3 This is a schematic diagram showing the distribution of two particle sizes of drug-loaded microcapsules in the composite drug coating of the present invention.

[0029] Figure 4 This is a schematic diagram of the ultrasonic response time-sequential drug release mechanism of the dual-size microcapsules of the present invention.

[0030] Figure 5 This is a schematic diagram showing the positional relationship of the stent implanted in aortic dissection according to the present invention.

[0031] Figure labels: 1- scaffold matrix; 2- biodegradable polymer barrier coating; 3- composite drug coating; 4- small-diameter microcapsule; 5- large-diameter microcapsule; 6- ultrasonically responsive release; 7- interstitial rupture; 8- false cavity. Detailed Implementation

[0032] To make the objectives, technical solutions, and advantages of the embodiments of the present invention clearer, the technical solutions in the embodiments of the present invention are described clearly and completely below. Obviously, the described embodiments are only a part of the embodiments of the present invention, not all of them. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the protection scope of the present invention. Therefore, the following detailed description of the embodiments of the present invention is not intended to limit the scope of protection of the present invention, but merely to represent selected embodiments of the present invention. All embodiments based on the present invention are within the protection scope of the present invention.

[0033] Performance Indicators Explanation: The performance values ​​listed in this article, such as radial support force, steady-state ion concentration, interfacial peel strength, and shape memory recovery rate, are test results of the corresponding embodiments under specific parameter conditions, and represent the performance under preferred parameters; the performance of the stent with different parameter configurations may fluctuate within a reasonable range, all of which fall within the protection scope of this invention.

[0034] Example 1 The alloy composition is as follows by mass percentage: Mn 30%, Si 5%, Cu 2%, Nb 0.2%, C ≤0.015%, with the balance being Fe and unavoidable impurities totaling no more than 0.5 wt%. The alloy's martensitic transformation initiation temperature (Ms) is 8℃, the reverse transformation initiation temperature (As) is 30℃, and the austenitic transformation completion temperature (A) is... f The temperature was 36℃. The structural parameters of the scaffold substrate 1 were: wall thickness 0.15 mm, fully expanded diameter 32 mm, compression transport outer diameter 18 F; the width of the wide hinge elastic deformation zone was 4 times that of the support rib, and the maximum equivalent strain of the hinge zone after compression was ≤3.5%. After 20 min of vacuum shaping at 380℃ and 12 rounds of thermomechanical cycling training (room temperature pre-compression - 37℃ recovery), the diameter recovery rate was 87% after 10 shape memory cycles. A 150 nm anodic oxide passivation layer was prepared on the substrate. A microporous structure layer was formed on the surface of the scaffold substrate 1 by femtosecond laser direct writing combined with anodic electrochemical etching: femtosecond laser wavelength 1030 nm, pulse width 300 fs, frequency 100 kHz, scanning ablation to form positive conical micro-pit pores; 12 wt% sodium phosphate electrolyte was used, with a current density of 8 mA / cm². 2Anodic etching for 90 s was performed, utilizing the current concentration effect at the bottom of the pores to expand the pore size and form a necking structure. The resulting micropores had an opening diameter of 0.8 μm-1.0 μm, a bottom diameter of 2 μm-3 μm, and a porosity of approximately 20%. A PLCL biodegradable polymer barrier film 2, approximately 50 μm thick, was electrospun onto the outer surface of the scaffold substrate 1. Electrospinning parameters were: 10% PLCL hexafluoroisopropanol solution, voltage 15 kV, receiving distance 12 cm, and spinning rate 0.8 mL / h. Plasma etching was then performed on the surface of the biodegradable polymer barrier film 2 to form 0.5-0.7 μm micro-nano pits, creating a two-level interlocking structure. The composite drug coating 3 used PLGA (molecular weight 50kDa-80 kDa, lactic acid:glycolic acid = 75:25) as the carrier; small-particle microcapsules 4, with a particle size of 0.5 μm-2 μm, encapsulated dexamethasone and vascular endothelial growth factor (VEGF), primarily responding to ultrasound at 0.8 MHz-2 MHz; large-particle microcapsules 5, with a particle size of 5 μm-15 μm, encapsulated rapamycin, primarily responding to ultrasound at 200 kHz-600 kHz. VEGF encapsulation employed a S / O / W method combined with low-temperature spray drying: VEGF lyophilized powder (containing 2% trehalose and 1% mannitol as a protectant) was dispersed in an organic phase containing PLGA in ethyl acetate / dichloromethane (volume ratio 1:1), emulsified at 4°C, and then spray-dried to prepare microcapsules. After encapsulation, the microcapsules were reconstituted, and the encapsulation efficiency was determined to be 72% using the BCA protein quantification method. The MTT assay was used to detect the proliferative activity of human umbilical vein endothelial cells, verifying the integrity of VEGF's biological function.

[0035] The preparation steps are as follows: (1) Weigh high-purity raw materials according to the composition ratio, and prepare alloy ingots by vacuum induction melting under argon protection; (2) The ingots are homogenized and annealed at 900℃ for 12 h, hot extruded at 850℃, and drawn at 300℃ in multiple passes (single pass deformation amount 12%, total deformation amount 70%, and intermediate annealing at 700℃ for 1 h every 3 passes) to make alloy tubes with specifications of φ18 mm × 0.25 mm and grain size of 7; (3) The tubes are cut by femtosecond laser to form a hollow grid support structure. Laser parameters: wavelength 1030 nm, pulse width 300 fs, frequency 200 kHz, scanning speed 250 mm / s, argon gas assistance; the cutting pattern includes a wide hinge elastic deformation area (the hinge width is 4 times the support rib, and the corner radius is 0.5 mm). The support blank is obtained by deburring by electrochemical polishing; (4) Vacuum shaping at 380℃ for 20 h min, followed by 12 rounds of "room temperature pre-compression - 37℃ recovery" thermomechanical cycle training to give the stent a stable shape memory effect; (5) a microporous structure layer with necked orifice was prepared by femtosecond laser direct writing combined with anodic electrochemical etching, and after cleaning and drying, it was subjected to anodic oxidation passivation treatment, and then electrospinning was used to prepare PLCL biodegradable polymer isolation coating 2. The outer surface of biodegradable polymer isolation coating 2 was subjected to plasma etching treatment; (6) S / O / W emulsification-solvent evaporation method combined with low temperature spray drying was used to prepare hollow drug-loaded microcapsules of two particle sizes respectively. Protein drugs were operated at 4℃ throughout the process; (7) two-step ultrasonic atomization spraying method: first, a PLGA solution containing large-particle-size microcapsules 5 was sprayed to form the bottom layer, and after drying, a PLGA solution containing small-particle-size microcapsules 4 was sprayed to form the shallow layer. The spraying temperature was 6℃, and the residual solvent was removed by vacuum drying for 24 h to obtain the finished stent.

[0036] No spontaneous expansion was observed after standing at room temperature for 4 hours, and the rate of increase in delivery resistance was ≤8%. After 6 months of immersion in simulated body fluid at 37℃, the radial support force was ≥52 kPa. Copper ions reached their peak release rate at 7 days, and the release rate dropped to less than 30% of the initial value after 14 days, with a steady-state concentration of 8-15 μmol / L. The steady-state concentration of manganese ions was ≤170 μmol / L after 30 days, with no cytotoxicity. The targeted ultrasound-triggered drug release rate was excellent, and the cross-response of non-target frequencies was ≤12%. In the axial peel test of the finished stent, the average peak peel load of the coating corresponding to a single support rib was 24.8 N. The failure mode was mainly pull-out of the polymer anchoring structure in the micropores accompanied by localized coating tearing, with no large-area interface debonding. The coating peeling rate during simulated sheath push was ≤8%.

[0037] Example 2 The alloy composition is as follows by mass percentage: Mn 28%, Si 4%, Cu 1%, Nb 0.1%, C ≤0.015%, with the balance being Fe and unavoidable impurities totaling no more than 0.5 wt%. The alloy temperature is set at Ms 10℃, As at 32℃, and A... fThe temperature was 37℃. The material shape recovery rate at 2.8% strain was 93%. The support substrate 1 has a wall thickness of 0.25 mm, a fully expanded diameter of 20 mm, a compression delivery outer diameter of 16 F, a wide hinge width three times that of the support rib, and a maximum equivalent strain of ≤3.2% under compression. After vacuum shaping at 350℃ for 15 min and 10 rounds of thermomechanical training, the diameter recovery rate after 10 cycles was 86%. The substrate passivation layer is 100 nm, the microporous structure layer of the support has pore sizes of 0.5 μm-0.8 μm at the orifice and 1 μm-2 μm at the bottom, with a porosity of 10%. A PCL biodegradable polymer barrier film 2, approximately 30 μm thick, is coated on the outer surface. Micro-nano pits with opening sizes of 0.3-0.5 μm are formed on the outer surface of the biodegradable polymer barrier film 2 by plasma etching. The composite drug coating 3 contains small-diameter microcapsules 4 encapsulating prednisolone and fibroblast growth factor; large-diameter microcapsules 5 encapsulate zotamoxetine; the lyophilized fibroblast growth factor powder contains 2% trehalose and 1% mannitol as a protectant. The remaining material parameters and preparation steps are consistent with those in Example 1.

[0038] Stable delivery at room temperature for 4 hours with a resistance increase rate ≤7%; radial support force ≥50 kPa after 6 months; steady-state copper ion concentration 5-10 μmol / L; manganese ion concentration ≤145 μmol / L after 30 days; average peak axial peel load of the finished stent is 21.2 N, with the failure mode mainly being microporous anchorage pull-out at the interface, accompanied by localized tearing of the covering body; coating push-off rate ≤10%, adaptable to the tortuous anatomical environment of the aortic arch.

[0039] Example 3 The alloy composition is as follows by mass percentage: Mn 33%, Si 6%, Cu 3.0%, Nb 0.5%, C ≤0.015%, with the balance being Fe and unavoidable impurities totaling no more than 0.5 wt%. The alloy temperature is Ms = 5℃, As = 28℃, and A = ... f The temperature was 35℃. The material shape recovery rate at 2.8% strain was 95%. The scaffold substrate 1 has a wall thickness of 0.35 mm, a fully expanded diameter of 46 mm, a compression delivery outer diameter of 20 F, a wide hinge width 5 times that of the support rib, and a maximum equivalent strain of ≤3.8% under compression. After vacuum shaping at 400℃ for 30 min and 15 rounds of thermomechanical training, the diameter recovery rate was 88% after 10 cycles. The substrate passivation layer is 300 nm, the scaffold microporous structure layer has pore sizes of 0.8 μm-1.0 μm at the orifice and 3 μm-5 μm at the bottom, and a porosity of 30%. The outer surface is coated with a PLCL biodegradable polymer barrier film 2, approximately 100 μm thick. The micro-nano pit opening size on the outer surface of the biodegradable polymer barrier film 2 is 0.6-0.8 μm. The composite drug coating 3 contains small-particle microcapsules 4 encapsulating dexamethasone and fibroblast growth factor; and large-particle microcapsules 5 encapsulating paclitaxel. The remaining parameters and preparation process are the same as in Example 1.

[0040] The stent was delivered without interruption at room temperature for 4 hours, with a resistance increase rate of ≤9%; the radial support force was ≥55 kPa after 6 months; the steady-state copper ion concentration was 15-20 μmol / L, which had the best effect on promoting endothelial repair; the manganese ion concentration was ≤190 μmol / L after 30 days, which met the safety standards; the average peak load of axial peeling of the finished stent was 27.5 N, and the failure mode was mainly the fracture of the cover body, while the interface anchoring structure remained intact, indicating that the interface bonding strength was higher than the strength of the cover body; the coating push-off rate was ≤7%, the interface bonding strength was optimal, and it could withstand the impact of high-pressure blood flow in the abdominal aorta.

[0041] Example 4 The alloy composition, structural parameters, and preparation process were the same as in Example 1, with the only difference being the drug combination: small-diameter microcapsules 4 encapsulated prednisolone and fibroblast growth factor (bFGF), and large-diameter microcapsules 5 encapsulated rapamycin. The bFGF lyophilized powder contained 2% trehalose and 1% mannitol as protective agents, and the same microcapsule preparation process as in Example 1 was used.

[0042] Tests showed that the diameter recovery rate after 10 shape memory cycles was 87%, the increase rate of delivery resistance at room temperature for 4 hours was ≤8%, and the radial support force was ≥52 kPa after 6 months. The steady-state concentration of copper ions was 8-15 μmol / L, and the concentration of manganese ions after 30 days was ≤170 μmol / L. The interfacial peel strength was 24.5 N, and the coating peeling rate was ≤8%. Prednisolone can effectively inhibit the acute inflammatory response, and bFGF can promote the repair of the smooth muscle layer. The timing of its anti-proliferative effect matches that of rapamycin, and its pharmacological function is complete, achieving a therapeutic effect comparable to that of Example 1.

[0043] Example 5 The alloy composition, structural parameters, and preparation process are the same as in Example 1, with the only difference being the drug combination: small-particle-size microcapsules 4 encapsulate dexamethasone and vascular endothelial growth factor (VEGF), and large-particle-size microcapsules 5 encapsulate zotamoxetine.

[0044] Tests showed that: the diameter recovery rate after 10 shape memory cycles was 86%; the increase rate of delivery resistance at room temperature for 4 hours was ≤9%; the radial support force after 6 months was ≥51 kPa; the steady-state concentration of copper ions was 8-15 μmol / L; the concentration of manganese ions after 30 days was ≤172 μmol / L; the interfacial peel strength was 24.2 N; and the coating peeling rate was ≤9%. Zotamoxetine can effectively inhibit the excessive proliferation of smooth muscle cells, and its effects are synergistic with those of dexamethasone and VEGF in the acute phase, without pharmacological antagonism. The therapeutic effect is comparable to that in Example 1.

[0045] Example 6 The alloy composition, structural parameters, and preparation process are the same as in Example 1. The only difference is the drug combination: small-particle-size microcapsules 4 are loaded with dexamethasone, prednisolone, VEGF, and bFGF (two anti-inflammatory drugs and two repair active ingredients used in combination), and large-particle-size microcapsules 5 are loaded with rapamycin and paclitaxel (two anti-proliferative drugs used in combination).

[0046] Tests showed that the diameter recovery rate after 10 shape memory cycles was 87%, the increase rate of delivery resistance at room temperature for 4 hours was ≤8%, and the radial support force was ≥52 kPa after 6 months; the steady-state concentration of copper ions was 8-15 μmol / L, and the concentration of manganese ions after 30 days was ≤170 μmol / L; the interfacial peel strength was 24.7 N, and the coating peeling rate was ≤8%; combined use of drugs can enhance anti-inflammatory, repair, and anti-proliferative effects, with no adverse interactions between drugs, which can meet the intensive treatment needs of critically ill patients.

[0047] Example 7 The alloy composition is the same as in Example 1, and the preparation process adopts the lower limits of the key process parameters in claim 10: 10% deformation per pass during warm drawing, 60% total deformation; energy density of 2 J / cm² for femtosecond laser micropore fabrication; sodium phosphate electrolyte concentration of 10 wt%, anodic etching current density of 5 mA / cm², and processing time of 60 s; emulsification rotation speed of 8000 r / min for microcapsule fabrication. The remaining structural parameters, biodegradable polymer barrier coating 2, and composite drug coating 3 are the same as in Example 1.

[0048] Testing revealed the following: 85% diameter recovery rate after 10 shape memory cycles; ≤9% increase in transport resistance at room temperature over 4 hours; ≥50 kPa radial support force after 6 months; micropore opening diameter of 0.5 μm-0.7 μm, bottom diameter of 1 μm-2 μm, and porosity of approximately 15%; interfacial peel strength of 22.3 N, approximately 48% higher than that of a smooth substrate; steady-state copper ion concentration of 7-13 μmol / L; ≤165 μmol / L manganese ion concentration after 30 days; and coating peeling rate of ≤9%. All performance characteristics meet clinical application requirements, demonstrating that the lower limit of process parameters can stably produce qualified products.

[0049] Example 8 The alloy composition is the same as in Example 1, and the preparation process adopts the upper limits of the key process parameters in claim 10: 15% deformation per pass during warm drawing, 75% total deformation; 5 J / cm² energy density for femtosecond laser micropore fabrication; 15 wt% sodium phosphate electrolyte concentration; 10 mA / cm² anolyte etching current density; 120 s processing time; and 12000 r / min emulsification rotation speed for microcapsule fabrication. The remaining structural parameters, biodegradable polymer barrier coating 2, and composite drug coating 3 are the same as in Example 1.

[0050] Testing revealed the following: 88% diameter recovery rate after 10 shape memory cycles; ≤8% increase in transport resistance at room temperature over 4 hours; ≥53 kPa radial support force after 6 months; micropore opening diameter of 0.9 μm-1.0 μm, bottom diameter of 4 μm-5 μm, and porosity of approximately 28%; interfacial peel strength of 26.1 N; steady-state copper ion concentration of 9-17 μmol / L; ≤175 μmol / L manganese ion concentration after 30 days; and ≤7% coating peeling rate. All performance characteristics meet clinical application requirements, demonstrating that the upper limit of the process parameters can stably produce qualified products.

[0051] Example 9 The alloy composition, structure, and preparation process are the same as in Example 1, except that the molar ratio of lactic acid to glycolic acid in the composite drug coating 3 carrier PLGA is 50:50, and the molecular weight is 40 kDa-60 kDa.

[0052] Testing revealed the following: 87% diameter recovery rate after 10 shape memory cycles; ≤8% increase in delivery resistance at room temperature over 4 hours; ≥52 kPa radial support force after 6 months; slightly faster PLGA degradation rate; approximately 15% higher basal drug release rate compared to the 75:25 group; approximately 12% higher cumulative drug release over 14 days without ultrasound stimulation; ≤13% ultrasound cross-response; 24.3 N interfacial peel strength; ≤8% coating peeling rate; and no significant difference in copper and manganese ion release concentrations compared to Example 1. This meets the clinical requirements for early and rapid drug release and its performance is satisfactory.

[0053] Example 10 The alloy composition, structure, and preparation process are the same as in Example 1, except that the molar ratio of lactic acid to glycolic acid in the composite drug coating 3 carrier PLGA is 85:15, and the molecular weight is 60 kDa-80 kDa.

[0054] Testing revealed the following: 87% diameter recovery rate after 10 shape memory cycles; ≤8% increase in delivery resistance at room temperature over 4 hours; and ≥52 kPa radial support force after 6 months. PLGA degradation was slightly slower, with the basic drug release rate decreasing by approximately 12% compared to the 75:25 group. The cumulative drug release over 14 days without ultrasound stimulation decreased by approximately 10%. Ultrasonic cross-response was ≤12%. Interface peel strength was 25.1 N, and coating peeling rate was ≤7%. Copper and manganese ion release concentrations showed no significant difference compared to Example 1. This meets the requirements for clinical scenarios requiring long-term stable drug release, and its performance is satisfactory.

[0055] Comparative Example 1 The procedure utilizes a commercially available permanent nickel-titanium alloy aortic stent graft (a commonly used clinical control). The stent is non-degradable, relies on superelastic self-expansion for release, and has a permanent polyethylene terephthalate (PET) coating without a drug-loaded coating. This stent's metallic framework remains permanently within the blood vessel, continuously stimulating the vessel wall over a long period; it lacks copper ion sustained-release function to promote endothelial repair and has no ultrasound-guided time-series drug release capability.

[0056] Comparative Example 2 This stent utilizes a conventional iron-based biodegradable large vessel stent with an alloy composition of Fe-30Mn-5Si-0.2Nb (copper-free). It lacks thermomechanical shape memory training and requires balloon dilation upon implantation. The stent surface lacks a microporous structure layer; the outer surface is coated with a PLCL membrane, and only a single PLGA coating loaded with rapamycin is directly sprayed on. It lacks dual-particle-size, ultrasound-responsive drug-loaded microcapsules. This stent cannot achieve temperature-triggered self-expansion, and balloon dilation easily causes intimal damage. The alloy cannot release copper ions, lacking the function of actively inducing endothelial repair. The coating has low bonding strength with the smooth substrate, making it prone to peeling during delivery. Drug release is a passive diffusion mode, unable to achieve time-sequential drug delivery enhancement.

[0057] Comparative Example 3 The stent matrix 1 and the biodegradable polymer insulating membrane 2 are prepared using the same process as in Example 1, except that the outer surface of the stent is only coated with a pure PLGA coating and does not have drug-loaded microcapsules. This stent can achieve temperature-triggered adaptive adhesion, and the biodegradable polymer insulating membrane 2 can effectively isolate false lumen blood flow. The degradation process slowly releases copper ions to promote endothelial repair. However, it cannot release anti-inflammatory, repair, and anti-proliferative drugs in stages, and cannot use in vitro ultrasound to regulate drug release, making it difficult to specifically intervene in acute postoperative inflammation and mid-to-late-stage endometrial hyperplasia.

[0058] Comparative Example 4 The stent material, drug type, coating formulation, and preparation process are basically the same as in Example 1. The difference is that the stent substrate surface is not treated with microporous structuring, and the substrate surface remains smooth without a microporous structure layer. Testing showed that the peel strength at the coating-substrate interface was approximately 12.9 N, indicating a risk of coating slippage during stent expansion. Simultaneously, the outer surface of the coating was not treated with micro-nano pitting, resulting in decreased drug coating adhesion strength. The coating loss rate reached 22% after sheath advancement, posing a potential risk of coating detachment migrating with blood flow and inducing distal vascular embolism.

[0059] Comparative Example 5 The alloy composition, phase transformation temperature, wall thickness, diameter, passivation layer, coating, and preparation process of the support are all the same as in Example 1. The only difference is that the hollow structure does not have a wide hinge elastic deformation zone, and a conventional hinge with the same width as the support rib is used (the hinge width is equal to the support rib width). Testing showed that when the support is compressed to the outer diameter of the conveying structure, the peak equivalent strain in the hinge zone reaches 6.2%, exceeding the recoverable strain range of the alloy's reversible phase transformation; the diameter recovery rate after 10 shape memory cycles is 72%, a decrease of 15 percentage points compared to Example 1; after repeated compression, significant residual plastic deformation appears in the hinge zone, and the long-term cyclic support force decays faster; the conveying resistance increases by 19% after standing at room temperature for 4 hours, posing a risk of conveying jamming. These results directly prove that a wide hinge elastic deformation zone can effectively reduce the peak compression strain, improve shape memory recovery stability and conveying reliability, and is an essential structural feature of this invention.

[0060] Comparative Example 6 The scaffold alloy composition, structure, phase transition, coating, and preparation process were consistent with those in Example 1, the only difference being that the scaffold substrate was not subjected to anodic passivation treatment. Testing revealed that copper ion release had no significant regulatory effect, with a burst release peak occurring within one day, reaching a concentration of 32 μmol / L, exceeding the upper limit of safe concentration for cells; manganese ion release rate was significantly accelerated, with a steady-state concentration of 265 μmol / L after 30 days, exceeding the safety threshold and posing a neurotoxicity risk; simultaneously, the initial corrosion rate of the alloy was too rapid, causing a premature decrease in the radial support force of the scaffold, with a radial support force of only 38 kPa after 6 months, lower than clinical requirements. These results directly demonstrate the necessity of the anodic passivation layer in regulating ion release kinetics, ensuring degradation safety, and maintaining the support cycle.

[0061] Comparative Example 7 The alloy composition exceeds the limits defined in claim 2: Mn 25%, Si 3%, Cu 4%, Nb 1%, C ≤0.015%, with the balance being Fe and unavoidable impurities. The remaining structure and preparation process are consistent with those in Example 1.

[0062] Tests showed that the alloy Ms was at 15℃, As at 25℃, and A... f At 30°C, the delivery resistance increased by 29% after standing at room temperature for 2 hours, posing a risk of delivery blockage; the diameter recovery rate after 10 shape memory cycles was only 74%, indicating poor recovery stability; the radial support force after 6 months was only 42 kPa, lower than clinical requirements; the steady-state concentration of copper ions reached 28 μmol / L after 30 days, exceeding the upper limit of the safe concentration, posing a risk of cytotoxicity; the steady-state concentration of manganese ions reached 210 μmol / L after 30 days, exceeding the safety threshold. These results demonstrate that when the components exceed the scope defined in the claims, the phase change characteristics, shape memory performance, mechanical support, and biosafety of the scaffold cannot meet the standards, validating the rationality of the component scope of the claims.

[0063] Comparative Example 8 The alloy composition and structural parameters are the same as in Example 1. The difference is that the two-step method of "femtosecond laser direct writing and anodic electrochemical etching" is not used to prepare micropores. Instead, the micropores are prepared by one-step ablation with femtosecond laser (without bottom expansion step and without necking structure). Furthermore, thermomechanical cycle training is not performed. The micropores are used directly after high-temperature shaping.

[0064] Testing revealed that the micropores exhibited a conical structure with no necking effect, resulting in a coating interface peel strength of only 14.2 N and a coating peeling rate of 18%. Without thermomechanical training, the diameter recovery rate after 10 shape memory cycles was only 70%, and the transport resistance increase rate at room temperature for 4 hours reached 20%, indicating poor transport stability. These results demonstrate that the two-step necking micropore process and thermomechanical cycle training are crucial for ensuring interfacial bonding strength and shape memory stability; their absence prevents the achievement of the technical effects of this invention.

[0065] Testing methods and evaluation standards 1. Alloy phase transformation temperature testing Differential scanning calorimetry (DSC) was used for testing in a high-purity nitrogen atmosphere. The heating and cooling rates were 10℃ / min, and the temperature range was -50℃ to 100℃. The martensitic transformation initiation temperature Ms was obtained from the cooling curve, and the reverse transformation initiation temperature As and austenitic transformation completion temperature A were obtained from the heating curve. f .

[0066] 2. Shape memory recovery performance test Compress the support to the corresponding outer diameter of the conveyor and fix it. Place it in a 37℃ constant temperature water bath for 5 minutes and then remove it. Use a high-precision vernier caliper to measure the diameter of the support after it has fully recovered. Calculate the diameter recovery rate = recovered diameter / nominal fully expanded diameter × 100%. Repeat the "compression-recovery" cycle 10 times and record the diameter recovery rate of the final cycle.

[0067] 3. Radial support force test Using a radial support force tester, the stent was immersed in simulated body fluid at 37℃ for 6 months and then removed. The stent was radially compressed at a constant rate, and the radial load when the stent diameter was compressed by 10% was recorded, which is the radial support force.

[0068] 4. Conveying stability test Using a simulated vascular sheath delivery device, a compressed stent was inserted into a delivery sheath of the corresponding specification, and the initial delivery resistance was recorded. After standing at room temperature for 4 hours, the delivery resistance was tested again, and the resistance increase rate was calculated as (resistance after standing - initial resistance) / initial resistance × 100%. At the same time, it was observed whether the stent would jam or spontaneously expand.

[0069] 5. Ion release performance test The stent was immersed in simulated body fluid at 37°C and pH 7.4, and samples were taken at preset time points. The concentrations of Mn and Cu ions in the extract were determined using inductively coupled plasma optical emission spectrometry (ICP-OES).

[0070] 6. Interface integration performance testing Axial tensile peel test: After the specimen was conditioned to standard environmental conditions for 24 hours, the metal substrate end and the free end of the coating were clamped in the upper and lower clamps of the testing machine, respectively. The centering was adjusted so that the loading direction was parallel to the axis of the support rib. The loading rate was set to (50±5) mm / min, and the specimen was loaded uniformly until the coating and substrate were completely separated or the specimen broke. The load-displacement curve and failure mode were recorded, and the peak load was extracted as the characterization value of the interface peel strength. Coating peel rate: The support was inserted into the delivery sheath and pushed repeatedly 3 times. After removal, the coating mass loss rate was calculated by weighing or image analysis, which is the coating peel rate pushed by the sheath.

[0071] 7. Ultrasonic response drug release performance test The drug-loaded stent was placed in the release medium and triggered by ultrasound at different frequencies. The drug release was determined by high performance liquid chromatography (HPLC), and the drug release ratio at the target frequency and the proportion of drug release at non-target frequencies (cross-response) were calculated.

[0072] 8. Characterization of drug-loaded microcapsule properties The particle size distribution of microcapsules was determined using a laser particle size analyzer; the encapsulation efficiency of protein drugs was determined using a BCA protein quantitative kit; and the bioactivity of protein drugs was verified by the MTT assay and the proliferation rate of human umbilical vein endothelial cells.

[0073] Comprehensive Analysis of the Advantages of the Solution Table 1 Summary of key performance test results for each embodiment and comparative example

[0074] 1. Shape memory effect and transport stability Examples 1-10 showed a conveying resistance increase rate of ≤ 9% after standing at room temperature for 4 hours, with no spontaneous expansion or jamming. Comparative Example 5, without the wide hinge structure, had a peak strain of 6.2% in the hinge region, exceeding the alloy's recoverable range; the recovery rate after 10 cycles was only 72%, and the conveying resistance increased by 18% after 4 hours. Comparative Example 7, with its excessive component limits, and Comparative Example 8, lacking a key process, both exhibited decreased recovery rate and deteriorated conveying stability. The above data indicate that... f Maintaining a temperature of 35℃-37℃ and Ms of 0℃-10℃, along with a wide-hinge, low-strain structure, are essential conditions for ensuring stable room temperature delivery and body temperature self-expansion.

[0075] 2. The Influence of Wall Thickness Design on Mechanical and Conveying Properties The wall thicknesses of the three embodiments are 0.15 mm, 0.25 mm, and 0.35 mm, corresponding to fully expanded diameters of 32 mm, 20 mm, and 46 mm, respectively. Embodiment 2, with a wall thickness of 0.25 mm and a diameter of 20 mm, has a compressed delivery outer diameter of only 16F, suitable for tortuous segments such as the aortic arch, and exhibits good compliance. Embodiment 1, with a wall thickness of 0.15 mm and a diameter of 32 mm, balances compliance and support, representing a general specification for the descending thoracic aorta. Embodiment 3, with a wall thickness of 0.35 mm and a diameter of 46 mm, provides radial support ≥ 55 kPa after 6 months, suitable for high-pressure blood flow environments such as the abdominal aorta. Increasing wall thickness directly improves radial support, but a trade-off must be made between delivery outer diameter and compliance. The wall thickness range of 0.15 mm to 0.35 mm, combined with the diameter range of 20 mm to 46 mm, can cover different clinical scenarios from the thoracic aorta to the abdominal aorta.

[0076] 3. Enhanced effect of the combination of two-level interlocking interface structures Example 1 exhibits an interfacial peel strength of 24.8 N, representing an improvement of approximately 52% compared to Comparative Example 4 (approximately 12.9 N) on a smooth substrate. Comparative Example 4 shows a coating peel rate of 22%, while Example 1's is only ≤8%. The substrate necking micropores allow the coating to embed and form a mechanical anchor, while the micro-nano pits on the coating surface allow the coating to embed and prevent microcapsules from falling in. These two structural levels directly determine the interfacial reliability. Comparative Example 8 lacks the two-step necking process, resulting in a significant decrease in interfacial peel strength. Comparative Example 4, lacking both micropores and pits, significantly increases the risk of coating slippage and peeling.

[0077] 4. Timing control of dual-particle-size ultrasonically responsive drug loading Examples 1, 9, and 10 exhibit an ultrasound cross-response ≤ 13%, enabling non-invasive, staged drug release. Comparative Example 2 utilizes passive diffusion drug release, while Comparative Example 3 lacks drug-loaded microcapsules. Example 1 releases anti-inflammatory drugs and growth factors via 0.8-2 MHz ultrasound during the acute phase, and releases anti-proliferative drugs via 200-600 kHz ultrasound in the mid-to-late stages, with controllable timing. Examples 4-6 validated that different drug combinations can achieve time-sequential drug release, with pharmacological effects matching the vascular repair process. Comparative Example 3, while possessing shape memory and membrane-based isolation functions, lacks a time-sequential drug release mechanism.

[0078] 5. Ion release safety and the necessity of anodic oxide layers In Examples 1-10, the 30-day manganese ion concentration was ≤ 190 μmol / L, and the steady-state copper ion concentration was 5-20 μmol / L. In Comparative Example 6, which lacked an anodic passivation layer, the copper ion burst release reached 32 μmol / L on day 1, and the manganese ion concentration reached 265 μmol / L on day 30, exceeding the safety threshold. The radial support force of Comparative Example 6 at 6 months was only 38 kPa, lower than clinical requirements. In Comparative Example 7, the components exceeded limits, with both copper and manganese ion concentrations exceeding the safe range. The anodic oxide layer is a key structure for preventing ion burst release, maintaining support period, and ensuring biocompatibility.

[0079] 6. Comparison with permanent stents and existing biodegradable stents Comparative Example 1 is a nickel-titanium alloy permanent stent, which is non-degradable. Long-term foreign body stimulation can induce chronic inflammation, and it lacks endothelial repair promotion and ultrasound drug delivery functions. Comparative Example 2 is a conventional iron-based biodegradable stent, which lacks shape memory effect, requires balloon dilation, lacks copper ion endothelial promotion function, and has a coating peeling rate ≥ 18%. Examples 1-10 degrade gradually over 18-36 months, avoiding the long-term complications of permanent stents. Simultaneously, they achieve functional integration not found in existing biodegradable stents through shape memory self-expansion, ion sustained release, and ultrasound drug delivery.

[0080] 7. Verification of the range of process parameters and material proportions Examples 7-8 verified that the upper and lower limits of the preparation process parameters could be used to prepare stents with satisfactory performance. Examples 9-10 verified that the upper and lower limits of the PLGA molar ratio could be used to meet the requirements, proving that the parameter range defined in the claims is feasible and that the product performance within the range is stable and controllable.

[0081] Limitations of the experiment: All performance characterizations in this invention were performed in an in vitro simulated environment, which can fully verify the basic properties of the materials and structures. The long-term degradation behavior, shape memory recovery effect, ultrasonic drug release efficiency, and tissue compatibility of the stent in a living aorta can be further verified in vivo using a large animal aortic dissection model. The pulsating blood flow fatigue life of the stent can be verified by conducting specialized tests in accordance with medical device industry standards.

Claims

1. A shape-memory biodegradable metal-based large vessel stent for endovascular repair of aortic dissection, characterized in that, The system comprises a biodegradable metal scaffold matrix, a biodegradable polymer barrier coating covering the outer surface of the scaffold matrix, and a composite drug coating loaded on the outer surface of the biodegradable polymer barrier coating. The biodegradable metal scaffold is a hollow tubular structure made of an iron-manganese-silicon-copper biodegradable shape memory alloy. The alloy has a martensitic transformation initiation temperature of 0℃-10℃, a reverse transformation initiation temperature of 28℃-32℃, and an austenitic transformation completion temperature of 35℃-37℃. The hollow structure has a wide hinge elastic deformation zone with a width 3-5 times that of the supporting ribs. The scaffold wall thickness is 0.15 mm-0.35 mm, and the scaffold matrix surface has a microporous structure. The biodegradable polymer barrier coating is made of a biodegradable polymer material. The composite drug coating includes a biodegradable polymer carrier and at least two types of ultrasound-responsive drug-loaded microcapsules with different particle sizes. The smaller microcapsules have a particle size of 0.5 μm-2 μm and encapsulate anti-inflammatory drugs and vascular repair active ingredients, with a main response frequency of 0.

8. In vitro low-intensity focused ultrasound (HIFU) with a frequency of MHz-2MHz; large-particle microcapsules with a particle size of 5 μm-15 μm, loaded with anti-proliferative drugs, and with a main response frequency of 200 kHz-600 kHz.

2. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The biodegradable metal scaffold matrix is ​​made of iron-manganese-silicon-copper-niobium biodegradable shape memory alloy. By mass percentage, the composition includes: Mn 28%-33%, Si 4%-6%, Cu 1%-3%, Nb 0.1%-0.5%, C ≤ 0.02%, and the balance is Fe and unavoidable impurities with a total content not exceeding 0.5 wt%.

3. The shape-memory biodegradable metal-based large vessel stent according to claim 1, characterized in that, The support substrate undergoes vacuum high-temperature shaping at 350℃-400℃ and 10-15 rounds of "pre-deformation-heat recovery" thermomechanical cycle training.

4. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The stent substrate is a hollow tubular structure, integrally cut and formed from several sets of axially spaced corrugated support rings and staggered spirally arranged axial connecting ribs; the crests and troughs of adjacent corrugated support rings are arranged axially correspondingly, and the axial connecting ribs are connected between the crests and troughs of adjacent support rings; the stent has a diameter of 20 mm-46 mm after full expansion, which can be adapted to different segments of the aorta, and the outer diameter delivered after compression is ≤ 20 F.

5. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The surface of the support substrate is provided with an anodic passivation layer with a thickness of 100 nm-300 nm.

6. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The biodegradable coating layer is selected from polylactic acid-caprolactone copolymer or polycaprolactone, with a thickness of 30 μm-100 μm. The mechanical integrity maintenance time of the coating layer is matched with the radial support force maintenance time of the scaffold matrix.

7. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The biodegradable polymer carrier of the composite drug coating is a polylactic acid-glycolic acid copolymer; the ultrasound-responsive drug-loaded microcapsules are polymer microcapsule structures containing an inert gas core; the small-diameter microcapsules are distributed in the shallow layer of the coating, 0-10 μm from the coating surface, and the large-diameter microcapsules are distributed in the deep layer of the coating, 10 μm-30 μm from the coating surface, corresponding to drug release in the mid-to-late stages of vascular repair; the drug coating is entirely attached to the outer surface of the biodegradable polymer barrier membrane, and both types of drug-loaded microcapsules are completely embedded inside the polymer carrier of the coating, with a continuous polymer carrier layer remaining between the bottom of the microcapsule and the membrane, eliminating the risk of membrane penetration; the coating preparation process is controlled at ambient pressure and temperature of 4-10℃, with a microcapsule gas core retention rate ≥ 75% and a capsule wall breakage rate ≤ 12%; drug release from small-diameter microcapsules can be selectively enhanced by in vitro 0.8 MHz-2 MHz low-intensity focused ultrasound; drug release from large-diameter microcapsules can be selectively enhanced by in vitro 200 kHz-600 kHz low-intensity focused ultrasound.

8. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The anti-inflammatory drug is selected from at least one of dexamethasone and prednisolone; the vascular repair active ingredient is selected from at least one of vascular endothelial growth factor and fibroblast growth factor; and the anti-proliferative drug is selected from at least one of rapamycin, paclitaxel, and zotamoxetine.

9. The shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, The scaffold substrate has a microporous structure on its surface. The micropores have a necked-out structure at the pore opening, with an opening diameter of 0.5 μm-1 μm, a bottom diameter of 1 μm-5 μm, and a porosity of 10%-30%. The micropores are prepared by a two-step method combining femtosecond laser direct writing ablation and anodic electrochemical etching for bottom expansion. The biodegradable polymer barrier coating is embedded inside the micropores during the composite process to form a mechanical interlock. The outer surface of the biodegradable polymer barrier coating has a micro-nano pit structure formed by plasma etching. The pit opening size is 0.3 μm-0.8 μm. The polymer carrier of the composite drug coating is embedded in the pit to form a secondary interlock, and the drug-loaded microcapsule particle size is larger than the pit opening size and is blocked outside the pit.

10. The method for preparing a shape memory biodegradable metal-based large blood vessel stent according to claim 1, characterized in that, Includes the following steps: (1) Weigh the raw materials according to the composition ratio and prepare iron-manganese-silicon-copper-niobium alloy ingots by vacuum induction melting under argon protection; (2) After the alloy ingot is homogenized and annealed at 900℃ for 10 h-14 h, it is hot extruded at 800℃-900℃, and then hot drawn at 200℃-400℃ in multiple passes. The deformation amount per pass is 10%-15%, and the total deformation amount is 60%-75%. An intermediate annealing treatment at 700℃ for 1 h is performed every 3 passes to produce alloy pipes with set wall thickness and outer diameter. Finally, the grain size of the pipe is controlled at level 5-10. (3) A hollowed-out mesh structure is processed on the alloy tube using femtosecond laser cutting technology. The laser parameters are wavelength 1030nm, pulse width 300 fs, frequency 200 kHz, and scanning speed 200-300 mm / s. The auxiliary gas is high-purity argon. The hollowed-out mesh structure includes a wide hinge elastic deformation zone. The support blank is obtained after deburring and electrochemical polishing. (4) The blank of the support is shaped at high temperature in a vacuum environment of 350℃-400℃ for 15 min-30 min, and then subjected to 10-15 rounds of room temperature "pre-compression-37℃ recovery" thermomechanical cycle training to give the support a stable stress-triggered shape memory effect. (5) A two-step method combining femtosecond laser direct writing ablation and anodic electrochemical etching to expand the bottom was used to prepare a microporous structure layer with narrowed apertures: the femtosecond laser parameters were wavelength 1030 nm, pulse width 300 fs, frequency 100 kHz, and energy density 2-5 J / cm², which scanned and ablated to form a positive conical micro-pit base hole; then, anodic electrochemical etching was performed using 10 wt%-15 wt% sodium phosphate electrolyte, current density 5-10 mA / cm², and processing time 60-120 s, which utilized the current concentration effect at the bottom of the pit to selectively dissolve and widen the bottom of the hole, forming a narrowed aperture structure; after ultrasonic cleaning with deionized water and vacuum drying, the substrate was subjected to anodic oxidation passivation treatment, and then a biodegradable polymer barrier film was prepared by electrospinning or hot pressing composite process; the electrospinning process parameters were: solution concentration 8%-12%, voltage 12-18 kV, receiving distance 10-15 cm, and spinning rate 0.5-1.0 mL / h; (6) Small-particle-size hollow microcapsules loaded with anti-inflammatory drugs and repair active ingredients and large-particle-size hollow microcapsules loaded with anti-proliferative drugs were prepared by S / O / W emulsification-solvent evaporation method or low-temperature spray drying method, respectively. Among them, protein drugs were prepared into freeze-dried powders by using a composite freeze-drying protectant containing 1%-3% trehalose and mannitol for encapsulation. The organic phase used a mixed solvent of ethyl acetate and dichloromethane with a volume ratio of 1:

1. The aqueous phase contained 1%-2% polyvinyl alcohol emulsifier. The emulsification speed was 8000-12000 r / min. The emulsification and drying process was carried out in a low-temperature environment of 0℃-8℃ throughout. (7) A two-step ultrasonic atomization spraying process is adopted. First, a bottom coating solution containing the second microcapsule is sprayed. After vacuum drying, a shallow coating solution containing the first microcapsule is sprayed. The spraying process controls the ambient pressure to normal pressure and the temperature to 4-10℃. After sufficient vacuum drying to remove residual solvent, the finished bracket is obtained.

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