Abstract:
The invention discloses an absorbable iron alloy stent, comprising an iron alloy substrate and a degradable polymer in contact with the surface of the substrate, in which the degradable polymer has a weight average molecular weight of more than or equal to 20,000 and less than or equal to 1,000,000 and a polydispersity index of more than 1.0 and less than or equal to 50. After the iron alloy stent is implanted into the body, the degradable polymer degrades to produce a carboxyl group. After the degradable stent is implanted into the human body, the effects of oxygen-consuming corrosion enable the stent both to mainly function as a mechanical support and then to degrade gradually, and the amount of hydrogen produced in the degradation process does not reach the level that can lead to a risk of air embolism.

Description:
TECHNICAL FIELD 
       [0001]    The present invention relates to a degradable implantable medical device, and particularly relates to an absorbable iron-based alloy device capable of degrading rapidly and controllably within a predetermined period. 
       BACKGROUND ART 
       [0002]    At present, the implantable medical devices are usually made from metals and their alloys, ceramics, polymers and the related composite materials, wherein the metal-based implantable medical devices are particularly popular because of their superior mechanical properties, such as high strength, and high toughness. 
         [0003]    Iron as an important element in the human body, is involved in many biochemical processes, such as in the delivery of oxygen. Easily corrosive pure iron stents each of the shape similar to that of a clinically used metal stent, made by Peuster M et al. through a laser engraving method, were respectively implanted to the descending aortas of 16 New Zealand rabbits. The animal experimental results showed that there was no thrombosis complication within 6 to 18 months, and also no adverse events occurred. The pathological examination confirmed that there were no inflammation in local blood vessel walls and no obvious proliferation on smooth muscle cells, preliminarily indicating that the degradable iron stent is safe and reliable and has good application prospects. But the study also found that the corrosion rate of pure iron was relatively slow in vivo, which cannot meet the clinical degradation time requirement for the degradable stent, thus the corrosion rate of iron needed to be accelerated. 
         [0004]    Various techniques for improving the corrosion rate of iron have been continuously developed, including alloying, changing the iron metallurgical structure, or coating of a degradable polyester coating layer on the surface of the iron-based alloy stent. For the method for increasing the corrosion rate of the iron-based material by the polyester, the literature disclosed that the degradable polyester coating would produce a product with a carboxyl group in the degradation process in the human body, so that the pH value of the local microenvironment near the stent implantation position dropped to form a local subacid environment, thereby the overpotential of hydrogen evolution reaction on the surface of the iron-based alloy substrate was reduced, and the hydrogen evolution corrosion was produced in the iron-based alloy substrate, thus producing an iron salt as a degradation product. The literature also indicated that the degradation process of iron-based alloy was accompanied by the oxygen-consuming corrosion process and the hydrogen evolution corrosion process, and because the highest oxygen-consuming corrosion rate of a solution in the local subacid environment is a constant value, it is difficult to improve the corrosion rate of the iron-based alloy by speeding up the oxygen-consuming corrosion rate. Namely, the literature believed that the corrosion rate of the iron-based alloy was improved only by the hydrogen evolution corrosion in the degradation of iron-based alloy. In addition, the literature only provided experimental data to indicate that the corrosion rate of iron-based alloy was increased under the action of polyester, and did not disclose the molecular weight and molecular weight distribution of the polymer, namely, did not disclose the match between the degradable polymer degradation and the iron-based alloy substrate corrosion. The literature also did not provide any experimental data to prove that the iron-based alloy stent can meet the clinical early mechanical property requirement after being implanted into the human body, and also did not disclose the corrosion period of the stent, so that whether the stent meets the clinical property requirement for the stent cannot be known by those skilled in the art. 
         [0005]    In fact, a large amount of hydrogen produced by hydrogen evolution corrosion will cause the tissue tolerance risk such as formation of air embolism, so that the stent cannot be used clinically. Our early experiments showed that after nitrogen was introduced and oxygen was removed in the corrosion environment, the corrosion rate of iron-based alloy was greatly reduced. i.e., the hydrogen evolution corrosion indicated by the above-mentioned literature was not produced in the iron-based alloy in the body. If the corrosion rate of iron is too rapid, it is possible that the iron-based alloy stent at early stage (such as 3 months) after the implanting would not have sufficient structural integrity, and it is difficult to reach the radial support force required by the blood vessel clinically, so that the stent loses its clinical application value. Otherwise, if an increase in the corrosion rate of iron caused by the polymer is limited, the corrosion period of the iron-based alloy is longer, and it is difficult to meet the clinical degradation time requirement for the degradable stent. Therefore, it is necessary to adopt a specific degradable polymer that is matched with the iron-based alloy substrate, realizing the rapid and controllable corrosion of the iron-based alloy, thus obtaining an absorbable iron-based alloy stent capable of meeting the clinical requirements. 
       SUMMARY OF THE INVENTION 
       [0006]    The technical problem to be solved by the present invention is to provide an absorbable iron-based alloy stent capable of degrading rapidly and controllably within a predetermined period after being implanted into the body in order to overcome the shortcomings of the prior art. 
         [0007]    As the first technical solution adopted by the present invention, the absorbable iron-based alloy stent comprises an iron-based alloy substrate and a degradable polyester in contact with the surface of the substrate, in which the degradable polyester has a weight average molecular weight in range of 20,000 and 1,000,000 and a polydispersity index of between 1.2 and 30. 
         [0008]    As the second technical solution adopted by the present invention, the absorbable iron-based alloy stent comprises an iron-based alloy substrate and a degradable polyester in contact with the surface of the substrate, in which the degradable polyester has a weight average molecular weight in range of 20,000 and 1,000,000 and a polydispersity index of more than 1.0 and less than 1.2, or more than 30 and less than or equal to 50. 
         [0009]    As the third technical solution adopted by the present invention, the absorbable iron-based alloy stent comprises an iron-based alloy substrate and a degradable polymer in contact with the surface of the substrate, in which the degradable polymer has a weight average molecular weight in range of 20,000 and 1,000,000 and a polydispersity index of more than 1.0 and more than or equal to 50. After the iron-based alloy stent is implanted into the body, the degradable polymer degrades to produce a carboxyl group. In the third technical solution, the degradable polymer may be a degradable polyester, a blend of the degradable polyester and a degradable polyanhydride, or a copolymer formed by copolymerizing monomers forming the degradable polyester and the degradable polyanhydride. 
         [0010]    The iron-based alloy substrate described in the present invention refers to a bare iron-based alloy stent, and the iron-based alloy substrate is selected from pure iron or a medical iron-based alloy. At least one of nutrient elements and harmless elements in the human body, or less toxic elements, such as C, N, O, S, P, Mn, Pd, Si, W, Ti, Co, Cr, Cu, and Re may be doped into the pure iron to form the medical iron-based alloy. 
         [0011]    The numerical interval is in accordance with the mathematical knowledge, namely. [a, b] means more than or equal to a and less than or equal to b; (a, b] means more than a and less than or equal to b; [a, b) means more than or equal to a and less than b. The same symbols shall apply hereinafter without need of repetition. 
         [0012]    The term “rapid” means that the degradable polyester can accelerate the corrosion of the iron-based alloy substrate, so that the iron-based alloy substrate can completely corrode within 5 years after being implanted into the body. 
         [0013]    The term “controllable” means that the corrosion of the iron-based alloy substrate caused by the degradable polyester ensures that the iron-based alloy stent has good mechanical properties at early stage after being implanted into the human body, and also enables the stent to produce a small amount of hydrogen or no hydrogen. A small amount of hydrogen refers to an amount that is not sufficient to form a risk of air embolism clinically. 
         [0014]    The term “complete corrosion” means that the mass loss rate W of the iron-based alloy stent is more than or equal to 90 percent. 
         [0015]    The complete corrosion is characterized by a mass loss test of an animal experiment. The mass loss test is carried out by implanting an iron-based alloy stent with an iron-based alloy substrate (i.e., a bare stent excluding a degradable polymer) of which the mass is M 0  into the abdominal aorta of a rabbit, cutting out the iron-based alloy stent implanted into an animal body and the tissue in which the iron-based alloy stent is placed at a predetermined observation point in time, soaking the tissue together with the stent in a solution of certain concentration (such as 1 mol/L of a sodium hydroxide solution) so that the tissue is digested, and then taking a stent strut out of the solution, putting the stent strut into a solution of a certain concentration (such as 3% of a tartaric acid solution, and/or an organic solution) to be ultrasonically cleaned so that corrosion products on the surface of the stent are all stripped off or dissolved in the solution, taking the residual stent strut out of the solution, drying and weighing the stent strut to obtain the mass M t . The mass loss rate W is expressed by a percentage of the weight loss difference value of the stent strut after corrosion and cleaning in weight of the iron-based alloy substrate, as shown in Formula 1-1: 
         [0000]    
       
         
           
             
               
                 
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         [0016]    W—Mass loss rate 
         [0017]    M t —Mass of the residual stent strut after corrosion 
         [0018]    M 0 —Mass of the iron-based alloy substrate 
         [0019]    When the mass loss rate W of the stent is more than or equal to 90 percent, the iron-based alloy stent is deemed to be completely corroded. 
         [0020]    The good mechanical properties obtained at early stage of implantation into the body are determined by specific clinical requirements. Generally. “early stage” refers to within 1 month, or 3 months, or 6 months after implantation into the body. The mechanical properties can be tested and verified by an animal experiment, and expressed by early OCT follow-up or a radial support force test. When the OCT follow-up is carried out, there is no obvious difference between the surrounding area of the stent and the surrounding area of the stent at the beginning of implanting, or when the radial support force test is carried out, the radial support force is more than 23.3 kPa (175 mm mercury column), indicating that the stent has good mechanical properties at early stage of implantation into the body. 
         [0021]    Compared with the prior art, the iron-based alloy stent provided by the present invention does not produce hydrogen at all or only produces a small amount of hydrogen in the whole corrosion process, which is specifically tested and verified by means of an animal experiment. For example, the iron-based alloy stent is implanted into the abdominal aorta of a rabbit, and the periphery of the support strut is observed by a microscope at the same amplification factor at the predetermined observation point in time, such as 1 month, 3 months, 6 months, 2 years, 3 years or 5 years after the stent is implanted into the body; if few stent strut coating slightly bulges, the stent is deemed to produce a small amount of hydrogen during corrosion; if the stent strut uniformly corrodes, and there are no bubbles around the stent strut, the stent is deemed to not produce hydrogen at all in the corrosion process. 
         [0022]    In the above-mentioned first to third technical solutions, the weight average molecular weight of the degradable polyester is more than or equal to 20,000 and less than 100,000, or more than or equal to 100,000 and less than 250,000, or more than or equal to 250,000 and less than 400,000, or more than or equal to 400,000 and less than 600,000, or more than or equal to 600,000 and less than or equal to 1,000,000. 
         [0023]    In the above-mentioned first technical solution, the polydispersity index of the degradable polyester is more than or equal to 1.2 and less than 2, or more than or equal to 2 and less than 3, or more than or equal to 3 and less than 5, or more than or equal to 5 and less than 10, or more than or equal to 10 and less than 20, or more than or equal to 20 and less than or equal to 30. 
         [0024]    In the above-mentioned third technical solution, the polydispersity index of the degradable polyester is more than or equal to 1.2 and less than 2, or more than or equal to 2 and less than 3, or more than or equal to 3 and less than 5, or more than or equal to 5 and less than 10, or more than or equal to 10 and less than 20, or more than or equal to 20 and less than 30, or more than or equal to 30 and less than or equal to 50. 
         [0025]    In the above-mentioned first and second technical solutions, the mass ratio of the iron-based alloy substrate to the degradable polyester is more than or equal to 1 and less than or equal to 200. Furthermore, in the above-mentioned first and second technical solutions, the mass ratio of the iron-based alloy substrate to the degradable polyester is more than or equal to 5 and less than or equal to 50. 
         [0026]    In the above-mentioned third scheme, the mass ratio of the iron-based alloy substrate to the degradable polymer is more than or equal to 1 and less than or equal to 200. 
         [0027]    Furthermore, in the above-mentioned third scheme, the mass ratio of the iron-based alloy substrate to the degradable polymer is more than or equal to 5 and less than or equal to 50. 
         [0028]    In the above-mentioned first and second technical solutions, the surface of the iron-based alloy substrate is coated with the degradable polyester in the form of a coating layer. In the above-mentioned third technical solution, the surface of the iron-based alloy substrate is coated with the degradable polymer in the form of a coating layer. 
         [0029]    In the above-mentioned first to third technical solutions, the wall thickness of the iron-based alloy substrate is more than or equal to 30 μm and less than 50 μm, and the thickness of the degradable polyester coating or degradable polymer coating is more than or equal to 3 μm and less than 5 μm, or more than or equal to 5 μm and less than 10 μm, or more than or equal to 10 μm and less than 15 μm, or more than or equal to 15 μm and less than or equal to 20 μm. 
         [0030]    In the above-mentioned first to third technical solutions, the wall thickness of the iron-based alloy substrate is more than or equal to 50 μm and less than 100 μm, and the thickness of the degradable polyester coating or degradable polymer coating is more than or equal to 5 μm and less than 10 μm, or more than or equal to 10 μm and less than 15 μm, or more than or equal to 15 μm and less than 20 μm, or more than or equal to 20 μm and less than or equal to 25 μm. 
         [0031]    In the above-mentioned first to third technical solutions, the wall thickness of the iron-based alloy substrate is more than or equal to 100 μm and less than 200 μm, and the thickness of the degradable polyester coating or degradable polymer coating is more than or equal to 10 μm and less than 15 μm, or more than or equal to 15 μm and less than 20 μm, or more than or equal to 20 μm and less than 25 μm, or more than or equal to 25 μm and less than or equal to 35 μm. 
         [0032]    In the above-mentioned first to third technical solutions, the wall thickness of the iron-based alloy substrate is more than or equal to 200 μm and less than or equal to 300 μm, and the thickness of the degradable polyester coating or degradable polymer coating is more than or equal to 10 μm and less than 15 μm, or more than or equal to 15 μm and less than 20 μm, or more than or equal to 20 μm and less than 25 μm, or more than or equal to 25 μm or less than 35 μm, or more than or equal to 35 μm and less than or equal to 45 μm. 
         [0033]    In the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the degradable polyester is any one of the following: polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate) (PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly (ethyleneglycol adipate) (PEA), poly (lactic-co-glycolic acid) (PLGA), and poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV). 
         [0034]    In the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the degradable polyester comprises at least two kinds of the same type of degradable polyester polymers. The “same type” refers to a general term of polymers with the same structural unit (i.e., the monomers are the same) and different weight average molecular weights. The first kind of degradable polyester polymer has a weight average molecular weight of more than or equal to 20,000 and less than 100,000 and the second kind of degradable polyester polymer has a weight average molecular weight that is greater than or equal to 100,000 and less than or equal to 1,000,000. The mass ratio of the first kind of degradable polyester polymer to the second kind of degradable polyester polymer is in the range of 1:9 to 9:1. The same type of degradable polyester polymer is selected from the group consisting of polylactic acid (PLA), polyglycolic acid (PGA), poly(butylene succinate)(PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly(ethyleneglycol adipate) (PEA), poly(lactic-co-glycolic acid) (PLGA), or poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV). The surface of the stent can be respectively coated with at least two kinds of the same type of degradable polyester polymers with different weight average molecular weights respectively, and can also be coated with the uniformly mixed degradable polyester polymers with different weight average molecular weights. 
         [0035]    Furthermore, in the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the mass ratio of the first kind of degradable polyester polymer to the second kind of degradable polyester polymer is in the range of 1:5 to 5:1. 
         [0036]    In the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the degradable polyester comprises at least two kinds of degradable polyester polymers with high molecular weights, wherein the weight average molecular weights of at least two kinds of degradable polyester polymers with high molecular weights are more than or equal to 100,000 and less than 200,000, or more than or equal to 200,000 and less than 400,000, or more than or equal to 400,000 and less than 600,000, or more than or equal to 600,000 and less than or equal to 1000,000. 
         [0037]    In the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the degradable polyester is a mixture of at least two of polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate) (PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly(ethyleneglycol adipate) (PEA), poly(lactic-co-glycolic acid) (PLGA), and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), or a copolymer formed by copolymerizing at least two of monomers forming polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate)(PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly(ethyleneglycol adipate) (PEA), poly(lactic-co-glycolic acid) (PLGA), and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV). As an example, the degradable polyester may comprise polylactic acid (PLA) and poly (lactic-co-glycolic acid)(PLGA), wherein the weight average molecular weight of PLGA is more than or equal to 20,000 and less than 300,000, the weight average molecular weight of PLA is more than or equal to 20,000 and less than or equal to 1,000,000, and the content ratio of the two is in the range of 1:9 to 9:1. 
         [0038]    In the above-mentioned first to second technical solutions and the third technical solution, when the degradable polymer is a degradable polyester, the degradable polyester is a mixture of at least two kinds of degradable polyester polymers with different crystallinity, wherein the content of degradable polyester polymer with crystallinity in the range of 5% to 50% is in the range of 10% to 90% in percentage by mass. The degradable polyester polymer is selected from the group consisting of polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate)(PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly (ethyleneglycol adipate) (PEA), poly (lactic-co-glycolic acid) (PLGA), and poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV). 
         [0039]    At least two kinds of degradable polyester polymers with different crystallinity include the following: the degradable polyester may be a mixture of crystalline and non-crystalline degradable polyester polymers, or a blend of degradable polyester polymers with low crystallinity and high crystallinity. For example, the degradable polyester comprises polylactic acid (PLA) with crystallinity in the range of 5% to 50%, and the content of the polylactic acid (PLA) is in the range of 10% to 900% in percentage by mass. Preferably, the polylactic acid (PLA) may be poly (DL-lactic acid) or poly (L-lactic acid). 
         [0040]    In the above-mentioned third technical solution, when the degradable polymer is a mixture of a degradable polyester and a degradable polyanhydride, or a degradable copolymer formed by copolymerizing monomers forming the degradable polyester and the degradable polyanhydride, the degradable polyester and the degradable polyanhydride have weight average molecular weights more than or equal to 20,000 and less than or equal to 1,000,000, and polydispersity indexes of more than 1.0 and less than or equal to 50, the polyanhydride is selected from the group consisting of poly (1, 3-bis(p-carboxyphenoxy) propane-sebacic acid) and poly (erucic acid dimer-sebacic acid) or poly (fumaric-sebacic acid), and the degradable polyester is any one of the following: polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate)(PBS), poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly(ethyleneglycol adipate) (PEA), poly(lactic-co-glycolic acid) (PLGA), and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV). The mass ratio of the degradable polyester to the polyanhydride is in the range of 1:9 to 9:1. 
         [0041]    In the above-mentioned third technical solution, in the mixture of the degradable polyester and the degradable polyanhydride, the content of degradable polyester or degradable polyanhydride with crystallinity in the range of 5% to 50% is in the range of 10% to 90% in percentage by mass, the degradable polyester polymer is selected from the group consisting of polylactic acid (PLA), polyglycolic acid (PGA), poly (butylene succinate) (PBS) and poly (beta-hydroxybutyrate) (PHB), polycaprolactone (PCL), poly (ethyleneglycol adipate) (PEA), poly (lactic-co-glycolic acid) (PLGA), and poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV), and the polyanhydride is selected from the group consisting of poly (1, 3-bis(p-carboxyphenoxy) propane-sebacic acid) and poly(erucic acid dimer-sebacic acid) or poly(fumaric-sebacic acid). 
         [0042]    In the above-mentioned first to third technical solution, the degradable polyester or degradable polymer may also be in contact with the surface of the iron-based alloy substrate in the form of a non-coating layer, for example, the iron-based alloy substrate is provided with gaps or grooves, and the degradable polyester or the degradable polymer is arranged in the gaps or grooves; or the iron-based alloy substrate is provided with an inner cavity, and the degradable polyester or the degradable polymer is filled in the inner cavity. At least one of the non-coating contact form and the coating form in the above-mentioned example is selected. Namely, in the above-mentioned first to third technical solutions, the “surface” of “in contact with the surface of the substrate” not only refers to the outer surface, but also refers to all circumstances in which the degradable polyester or the degradable polymer has a contact point or contact surface with the iron-based alloy substrate. 
         [0043]    In the above-mentioned first to third technical solutions, the degradable polyester or the degradable polymer may also be mixed with an active drug, and the mass ratio of the degradable polyester or the degradable polymer to the drug is in the range of 0.1 to 20. The active drug may be a vascular proliferation inhibiting drug such as paclitaxel, rapamycin and their derivatives, or an antiplatelet drug selected from cilostazol, or an antithrombotic drug such as heparin, or an anti-inflammatory drug such as dexamethasone, or a mixture of the above-mentioned drugs. Furthermore, the mass ratio of the degradable polyester or the degradable polymer to the drug is in the range of 0.5 to 10. 
         [0044]    Compared with the prior art, the specific degradable polymer used by the absorbable iron-based alloy stent provided by the present invention can allow the iron-based alloy substrate to undergo oxygen-consuming corrosion under the action of the degradable polymer, with minimal or no hydrogen produced, thus avoiding the clinical air embolism risk caused by a large amount of hydrogen being produced by hydrogen evolution corrosion in the prior art, and also meeting the clinical early mechanical property requirements for the stent. 
     
    
     
       BRIEF DESCRIPTION OF THE ACCOMPANYING DRAWINGS 
         [0045]      FIG. 1  is a sectional schematic diagram of a stent strut after an iron-based alloy stent provided by Example 5 of the present invention is coated with a degradable polyester coating layer; 
           [0046]      FIG. 2  is a microphotograph illustrating that a small amount of hydrogen is produced in a stent strut in Example 7 of the present invention; 
           [0047]      FIG. 3  is a microphotograph illustrating that a small amount of hydrogen is produced in a stent strut in Example 9 of the present invention; 
           [0048]      FIG. 4  is a microphotograph illustrating that a small amount of hydrogen is produced in a stent strut in Example 12 of the present invention; 
           [0049]      FIG. 5  is a microphotograph illustrating that a small amount of hydrogen is produced in a stent strut in Example 13 of the present invention; 
           [0050]      FIG. 6  is a microphotograph illustrating that a small amount of hydrogen is produced in a stent strut in Example 14 of the present invention; 
           [0051]      FIG. 7  is a microphotograph illustrating that a large amount of hydrogen is produced in a stent strut in the corrosion process in Control Example 2. 
       
    
    
     DETAILED DESCRIPTION OF THE EMBODIMENTS 
       [0052]    According to the absorbable iron-based alloy stent provided by the present invention, animal experiments are adopted to test and verify whether the iron-based alloy stent can rapidly and controllably corrode under the action of a degradable polymer or not, whether the iron-based alloy stent controllably corrodes or not is mainly determined by early mechanical properties and whether a large amount of hydrogen is produced at predetermined observation points in time or not, and whether the iron-based stent rapidly corrodes or not is determined by a mass loss test. 
         [0053]    Specifically after the iron-based alloy stent containing the degradable polymer was implanted into an animal body, the test was carried out at each of the predetermined observation points in time. For example, at 3 months from the date of implantation, an OCT follow-up test was carried out, it was found that there was no obvious difference between the surrounding area of the stent strut at this observation point and that of the original stent strut at the beginning of implanting, or animals were killed humanely, the stent and the tissue in which the stent is placed were taken out of the body, and the stent together with the blood vessel in which the stent is placed were subject to a radial support force test to determine if the stent meets the early mechanical properties; a stent was taken out at 2 years to measure the mass loss of the stent so as to observe the corrosion situation of the stent. The stent taken out of the body was axially split at each observation point in time, and the periphery of the support strut of each observed stent was observed by a microscope at the same amplification factor so as to determine whether a large amount of hydrogen is produced in the corrosion process of the stent. 
         [0054]    The radial support force can be tested by means of a radial support force tester RX550-100 produced by the MSI Company; namely, the radial support force can be obtained by taking out the stent implanted into the animal body at a predetermined observation point in time together with the blood vessel and then directly testing. 
         [0055]    The complete corrosion is characterized by a mass loss test of an animal experiment. The mass loss test is carried out by implanting an iron-based alloy stent with an iron-based alloy substrate (i.e., a bare stent excluding a degradable polymer) of which the mass is M 0  into the abdominal aorta of a rabbit, cutting out the iron-based alloy stent implanted into an animal body and the tissue in which the iron-based alloy stent is placed at a predetermined observation point in time, soaking the tissue together with the stent in a solution of certain concentration (such as 1 mol/L of a sodium hydroxide solution) so that the tissue is digested, and then taking a stent strutout of the solution, putting the stent strut into a solution of certain concentration (such as 3% of a tartaric acid solution, and/or an organic solution) to be ultrasonically cleaned so that corrosion products on the surface of the stent are all stripped off or dissolved in the solution, taking the residual stent strut out of the solution, drying and weighing the stent strut to obtain the mass M t . The mass loss rate W is expressed by a percentage of the weight loss difference value of the stent strut after corrosion and cleaning in weight of the iron-based alloy substrate, as shown in Formula 1-1: 
         [0000]    
       
         
           
             
               
                 
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         [0056]    W—Mass loss rate 
         [0057]    M t —Mass of the residual stent strut after corrosion 
         [0058]    M 0 —Mass of the iron-based alloy substrate 
         [0059]    When the mass loss rate W of the stent is more than or equal to 90 percent, the iron-based alloy stent is deemed to completely corroded. The weight average molecular weight and the polydispersity index of the degradable polymer were tested by using an eight-angle laser light scattering instrument produced by the Wyatt Technology Corporation. 
         [0060]    The absorbable iron-based alloy stent provided by the present invention is further illustrated in conjunction with the following accompanying drawings and examples. It should be understood that the following examples are only preferred examples of the present invention described herein, but not to limit the present invention. Any modifications, equivalent replacements, improvements. etc. made within the spirit and principles of the present invention should fall in the scope of the present invention described herein. 
       Example 1 
       [0061]    A pure iron stent comprises a pure iron substrate and a degradable polymer coating with which the surface of the pure iron substrate is coated, wherein the mass ratio of the pure iron substrate to the degradable polymer is 5:1. The degradable polymer is polyglycolic acid (PLA) with a weight average molecular weight of 200,000 and a polydispersity index of 1.8, and the wall thickness of the iron substrate is between 80 μm and 90 μm, and the thickness of the degradable polymer coating is between 15 μm and 20 μm. The stent was implanted into the abdominal aorta of a rabbit. The stent and the tissue in which the stent was placed were taken out at 3 months after the date of implantation, a radial support force test was carried out, and the test result that the radial support force was 70 kPa was obtained, indicating that the degradable polymer was well matched with the iron-based alloy substrate, and the early mechanical properties of the stent could be ensured; the periphery of the stent strut was observed with a microscope, no hydrogen bubbles were found. After 2 years from the date of implantation, a mass loss test was carried out by sampling again, and the mass loss rate of the stent was 95 percent, indicating that the stent completely corroded; no hydrogen bubbles were found by observing the periphery of the stent strut with a microscope. 
       Example 2 
       [0062]    The surface of a bare nitrided pure iron stent (i.e., a nitrided pure iron substrate) of which the wall thickness is between 65 μm and 75 μm was uniformly coated with a 10 to 12 μm thick degradable polymer coating, wherein the mass ratio of the nitrided pure iron substrate to the degradable polymer is 25, and the degradable polymer coating is a poly(DL-lactic acid) coating with a weight average molecular weight of 100,000 and a polydispersity index of 3. The absorbable iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the coronary artery of a pig. At 3 months from the date of implantation, it was found that there was no difference between the surrounding area of the stent strut and the surrounding area of the stent strut at the beginning of implanting by OCT follow-up. The stent was taken out at 1 year after the implanting, the mass loss rate of the stent was 92 percent by a mass loss test, indicating that the stent completely corroded. At 3 months and 1 year after the date of implantation, the stent was taken out, and then no hydrogen bubbles were produced by observing the periphery of the stent strut with a microscope respectively. 
       Example 3 
       [0063]    The surface of a bare electrodeposited pure iron (550° C. annealing) stent (i.e., an electrodeposited pure iron substrate) of which the wall thickness is between 40 μm and 50 μm was uniformly coated with a 3 to 5 μm thick mixture coating of polycaprolactone (PCL) and paclitaxel, wherein the mass ratio of the electrodeposited pure iron substrate to the degradable polymer was 35:1, the polycaprolactone (PCL) was formed by mixing two kinds of polycaprolactones (PCL) with weight average molecular weights of 30,000 and 80,000 according to a ratio of 1 to 2, the polydispersity index of the mixed polycaprolactones (PCL) was 25, and the mass ratio of polycaprolactones (PCL) to paclitaxel was 2 to 1. An absorbable iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the abdominal aorta of a rabbit. The stent was taken out at a corresponding observation point in time, the surface of the stent was observed with a microscope, and the radial support force and the mass loss percentage of the stent were tested. The test results showed that the radial support force was 60 kPa at 3 months after the date of implantation; after 1 year from the date of implantation, the mass loss rate of the stent was 98 percent, indicating that the stent completely corroded, and there were no hydrogen bubbles around the stent strut by observing with a microscope at the two observation points in time. 
       Example 4 
       [0064]    The outer wall surface of a bare carburized iron stent (i.e., a carburized iron substrate) obtained after heat treatment was coated with a poly (L-lactic acid) coating, wherein the wall thickness of the carburized iron substrate is between 140 μm and 160 μm, the thickness of the poly (L-lactic acid) coating is between 30 μm and 35 μm, and the mass ratio of the carburized iron substrate to the poly (L-lactic acid) is 120. The coating comprises two layers. i.e., a PLLA coating with crystallinity of 50 percent as a bottom layer and a PLLA coating with crystallinity of 5 percent as a top layer, the weight average molecular weights of the two layers are 600,000, and the polydispersity indexes of the two layers of poly (L-lactic acid) are 1.2. The mass ratio of the degradable polymer coating with crystallinity of 50 percent to the degradable polymer coating with crystallinity of 5 percent is 1:1. An absorbable iron-based alloy stent was obtained after drying. The stent was implanted into the abdominal aorta of a rabbit. The stent was taken out at a corresponding observation point in time, the surface of the stent was observed with a microscope, and the radial support force and the weight loss percentage of the stent were tested. The test results showed that the radial support force was 45 kPa at 6 months after the implanting; after 3 years from the date of implantation, the mass loss rate of the stent was 92 percent, and there were no hydrogen bubbles around the stent strut at the above-mentioned two observation points in time. 
       Example 5 
       [0065]    A bare iron-manganese alloy stent (i.e., an iron-manganese alloy substrate) was polished so that grooves were distributed in the surface of the stent. As shown in  FIG. 1 , a stent strut of the stent has a thickness of between 100 μm and 120 μm, and grooves  2  are arranged in the surface of the stent strut  1 . A degradable polyester polymer mixture coating  3  was uniformly coated on the surface of the stent strut  1  and in the grooves  2 . The degradable polyester polymer coating was formed by mixing poly (L-lactic acid) with a weight average molecular weight of 1,000,000 and poly (lactic-glycolic acid) with a weight average molecular weight of 20,000 (the molar ratio of lactic acid to glycolic acid is 50:50) according to a mass ratio of 5 to 1, wherein the polydispersity index of the polyglycolic acid is 10 after mixing, the thickness of the mixture coating is between 20 μm and 25 μm, and the mass ratio of the iron-based alloy substrate to the degradable polymer is 40:1. An absorbable iron-based alloy stent was obtained after drying. The stent was implanted into the abdominal aorta of a pig. A stent was taken out at a corresponding observation point in time, and then the mass loss rate and the radial support force of the stent were tested. The test results showed that the radial support force was 60 kPa at 3 months after the implanting; after 2 years from the date of implantation, the mass loss rate of the stent was 95% by a mass loss test, and there were no hydrogen bubbles around the stent strut by observing at the above-mentioned two observation points in time. 
       Example 6 
       [0066]    The outer surface of a bare iron-carbon alloy stent (i.e., an iron-carbon alloy substrate) with a thickness of between 30 μm and 40 μm, excluding the inner wall of a tubular cavity of the stent, was uniformly coated with a 5 to 8 μm thick poly(butylene succinate) (PBS) coating, wherein the mass ratio of the iron-carbon alloy substrate to poly(butylene succinate) (PBS) is 12:1, and the poly(butylene succinate) (PBS) has a weight average molecular weight of 60,000 and a polydispersity index of 2. An absorbable iron-based alloy stent was obtained after drying. The stent was implanted into the abdominal aorta of a rabbit. The stent was taken out at a corresponding observation point in time, and a mass loss test and a radial support force test of the stent were carried out. The results showed that the radial support force of the stent was 50 kPa at 1 month after the implanting; after 1.5 years from the date of implantation, the mass loss rate of the stent was 99 percent, and no hydrogen bubbles were found by observing the periphery of the stent strut with a microscope at the above-mentioned observation points in time. 
       Example 7 
       [0067]    The surface of a bare sulfurized iron stent (i.e., a sulfurized iron-based alloy substrate) with a wall thickness of between 250 μm and 270 μm was uniformly coated with a 35 to 45 μm thick degradable polymer coating. The mass ratio of the sulfurized iron-based alloy substrate to the degradable polymer is 50:1, and the degradable polymer is formed by mixing polylactic acid (PLA) and PLGA, wherein the polylactic acid has a weight average molecular weight of 30,000, crystallinity of 40 percent, the content of 90 percent, and a polydispersity index of 1.8, and the PLGA has a weight average molecular weight of 30,000, a polydispersity index of 4, crystallinity of 5 percent and the content of 10 percent. An absorbable iron-based alloy stent was obtained after drying. The stent was implanted into the abdominal aorta of a pig. The stent was taken out at a corresponding observation point in time, and a mass loss test of the iron-based alloy stent was carried out. The test results showed that the radial support force of the stent was 50 kPa at 6 months after the date of implantation, and it was found that the periphery of few stent strut slightly bulges. i.e., a small amount of hydrogen was produced by observing the periphery of the iron-based alloy stent strut with a microscope as shown in  FIG. 2 ; after 4.5 years from the date of implantation, the mass loss rate of the stent was 90 percent, and no hydrogen bubbles were found. 
       Example 8 
       [0068]    The surface of a bare iron-manganese alloy stent (i.e., an iron-manganese alloy substrate) with a wall thickness of between 120 μm and 150 μm was coated with a 15 to 20 μm thick coating. The coating was formed by mixing poly (beta-hydroxybutyrate) (PHB), poly (fumaric-sebacic acid) and heparin according to a mass ratio of 8 to 1 to 1, wherein the mass ratio of the iron-based alloy substrate to a degradable polymer. i.e., the mass sum of poly (beta-hydroxybutyrate) (PHB) and poly (fumaric-sebacic acid) is 80. PLLA has a weight average molecular weight of 300,000, crystallinity of 30%, and a polydispersity index of 2, and the PLGA has a weight average molecular weight of 100,000 and a polydispersity index of 45. An absorbable iron-based alloy stent was obtained after drying. The stent was implanted into the abdominal aorta of a rabbit. The stent was taken out at a corresponding observation point in time, and then the radial support force test and the mass loss test of the iron-based alloy stent were carried out. The results showed that that the radial support force of the iron-based alloy stent was 60 kPa at 3 months after the date of implantation; after 3 years from the date of implantation, the mass loss rate of the stent was 95 percent, and no hydrogen bubbles were found by observing the periphery of the stent strut with a microscope at the above-mentioned observation points in time. 
       Example 9 
       [0069]    The surface of a bare carburized iron stent (i.e., a carburized iron substrate) with a wall thickness of between 50 μm and 70 μm was coated with a degradable polymer coating with an average thickness of between 12 μm and 15 μm. The degradable polymer coating was formed by mixing poly (DL-lactic acid) (PDLLA) and rapamycin according to a mass ratio of 2 to 1, wherein the PDLLA has a weight average molecular weight of 200,000 and a polydispersity index of 1.6, and the mass ratio of the carburized iron substrate to the degradable polymer coating is 30. An absorbable iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the coronary artery of a pig. The iron-based alloy stent was taken out at a corresponding observation point in time to undergo a mass loss test and a radial support force test. The results showed that at 3 months after the date of implantation, the radial support force was 60 kPa, and it was found that a small amount of hydrogen was produced in the local stent strut, and few iron-based alloy stent coating slightly bulges by observing the periphery of the iron-based alloy stent strut with a microscope as shown in  FIG. 3 . At 2 years after the date of implantation, the mass loss rate of the stent was 98 percent, and no hydrogen bubbles were found. 
       Example 10 
       [0070]    The surface of a bare pure iron stent (i.e., a pure iron substrate) of which the wall thickness is between 50 μm and 60 μm was uniformly coated with a 8 to 12 μm thick degradable polymer coating, wherein the mass ratio of the pure iron substrate to the degradable polymer coating is 20:1, the degradable polymer as a bottom layer in the degradable polymer coating is PLLA with a weight average molecular weight of 300,000 and a thickness of about 6 to 8 μm, the top layer is PDLLA with a weight average molecular weight of 300,000, and the polydispersity index of the degradable polymer coating is 15. An iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the coronary artery of a pig. OCT follow-up was carried out at 3 months after the date of implantation, the OCT testing results showed that there was no difference between the surrounding area of the iron-based alloy stent and the surrounding area of the iron-based alloy stent at the beginning of implanting. After 2.5 years from the date of implantation, a mass loss test of the stent was carried out by sampling, the mass loss rate of the stent was 98 percent, and it was found that no hydrogen bubbles were produced by observing the periphery of a stent strut by sampling at the above-mentioned two observation points in time. 
       Example 11 
       [0071]    The surface of a bare nitrided iron stent (i.e., a nitrided iron substrate) of which the wall thickness is between 60 μm and 90 μm was coated with a 10 to 15 μm thick poly(3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) coating with a weight average molecular weight of 400,000 and a polydispersity index of 3, and then coated with a mixed coating comprising polylactic acid (PLA), poly (erucic acid dimer-sebacic acid) and cilostazol by praying again. The mixed coating with a thickness of about 10 μm was mainly sprayed on the outer wall and the side walls of the nitrided iron substrate, wherein the polylactic acid (PLA) has a weight average molecular weight of 50,000 and a polydispersity index of 1.6, the poly (erucic acid dimer-sebacic acid) has a weight average molecular weight of 20,000 and a polydispersity index of 10, the mass ratio of the polylactic acid (PLA), the poly (erucic acid dimer-sebacic acid) to the cilostazol is 1:1:1, and the ratio of the mass of the nitrided iron substrate to the mass sum of the two degradable polymer coatings is 35:1. An iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the coronary artery of a pig. OCT follow-up was carried out at 3 months after the date of implantation, and the OCT testing results showed that there was no difference between the tubular cavity area of the iron-based alloy stent and the tubular cavity area of the iron-based alloy stent at the beginning of implanting. After 1.5 years from the date of implantation, the mass loss rate of the stent was 95 percent, and it was found that no hydrogen bubbles were produced around the stent strut by sampling at the two points in time. 
       Example 12 
       [0072]    A stent strut of a bare nitrided iron stent (i.e., a nitrided iron substrate) of which the wall thickness is between 220 μm and 240 μm was treated so that micropores and grooves are arranged in the stent strut, and poly (butylene succinate) (PBS) is uniformly filled in the micropores and the grooves, wherein the poly (butylene succinate)(PBS) has a weight average molecular weight of 150,000 and a polydispersity index of 5, and the mass ratio of the nitrided iron substrate to the poly (butylene succinate)(PBS) is 5:1. An iron-based alloy stent was obtained after drying. The stent was implanted into a rabbit body. The iron-based alloy stent was taken out at a corresponding observation point in time, and a mass loss test and a radial support force test were carried out. At 2 months after the date of implantation, the radial support force of the stent was 75 kPa, and it was found that a small amount of bubbles was produced by observing the periphery of the stent strut with a microscope shown in  FIG. 4 . The mass loss rate of the iron-based alloy stent was 90 percent and no hydrogen bubbles were found around the stent strut at 3 years after the date of implantation. 
       Example 13 
       [0073]    An iron-cobalt alloy stent comprises an iron-cobalt alloy substrate and a degradable polymer coating covering the surface of the substrate, wherein the wall thickness of the iron-cobalt alloy substrate is in the range of 280 μm to 300 μm, the degradable polymer coating is a copolymer coating formed by copolymerizing monomers forming PLLA and PGA, the mass ratio of the monomers forming the two kinds of degradable polymers is 9:1, the copolymer has a weight average molecular weight of 50,000, a polydispersity index of 1.1 and crystallinity of 50 percent, and the thickness of the copolymer coating is in the range of 35 μm to 45 μm. The copolymer coating also comprises rapamycin, the ratio of the mass sum of the two kinds of polymers to the mass of the drug is 0.1:1, and the mass ratio of the iron-cobalt alloy substrate to the polymer coating is 25:1. The iron-cobalt alloy stent was implanted into the abdominal aorta of a pig. The radial support force was tested by sampling, and the periphery of the stent strut was observed with a microscope at 3 months and 4.5 years after the date of implantation. The test results showed that at 3 months after the date of implantation, the radial support force of the iron-based alloy stent was 45 kPa and a small amount of hydrogen bubbles was produced around the stent strut as shown in  FIG. 5 ; the mass loss rate of the stent strut was 90 percent, and no hydrogen bubbles were produced around the stent strut at 5 years after the date of implantation. 
       Example 14 
       [0074]    The surface of a bare iron-carbon alloy stent (i.e., an iron-carbon alloy substrate) was coated with a degradable polyester coating, wherein the wall thickness of the iron-carbon alloy substrate is in the range of 180 μm to 200 μm, the thickness of the degradable polyester coating is in the range of 20 μm to 25 μm. The degradable polyester coating is formed by mixing poly (butylene succinate) (PBS) and polyglycolic acid (PGA) according to a mass ratio of 9:1, and the blend has a weight average molecular weight of 250,000 and a polydispersity index of 2. The degradable polyester coating may also be mixed with heparin, wherein the mass ratio of the degradable polyester to the heparin is 20:1, and the mass ratio of the iron-carbon alloy substrate to the degradable polyester is 40:1. An iron-based alloy stent was obtained after drying. The iron-based alloy stent was implanted into the abdominal aorta of a pig. The radial support force was tested by sampling, and the periphery of a stent strut was observed with a microscope at 3 months and 3 years after the date of implantation. The test results showed that the radial support force was 75 kPa at 3 months after the date of implantation, and a small amount of hydrogen bubbles was produced around the stent strut as shown in  FIG. 6 ; the mass loss rate of the stent strut was 95 percent, and no hydrogen bubbles were produced around the stent strut at 3 years after the date of implantation. 
       Example 15 
       [0075]    An iron-nitrogen alloy stent comprises an iron-nitrogen alloy substrate and a degradable polymer coating covering the surface of the substrate, wherein the wall thickness of the iron-nitrogen alloy substrate is in the range of 90 μm to 100 μm, and the thickness of the degradable polymer coating is in the range of 15 μm to 20 μm. The coating is formed by mixing polylactic acid (PLA) and poly(ethyleneglycol adipate) (PEA) according to a mass ratio of 1 to 5, wherein the weight average molecular weights of the polylactic acid (PLA) and the poly(ethyleneglycol adipate) (PEA) are 500.000 and 300,000 respectively, the polydispersity index of the degradable polyester coating is 3, and the mass ratio of the iron-nitrogen alloy substrate to the degradable polyester coating is 10:1. The iron-based alloy stent was implanted into the abdominal aorta of a rabbit. The radial support force was tested by sampling, and the periphery of a stent strut was observed by a microscope at 3 months and 3 years after the date of implantation. The test results showed that the radial support force of the iron-based alloy stent was 50 kPa, and no hydrogen bubbles were produced around the stent strut at 3 months after the date of implantation; the mass loss rate of the stent strut was 95 percent, and no hydrogen bubbles were produced around the stent strut at 3 years after the date of implantation. 
       Example 16 
       [0076]    An iron-palladium alloy stent comprises an iron-palladium alloy substrate and a degradable polymer coating covering the surface of the substrate, wherein the wall thickness of the iron-palladium alloy substrate is in the range of 70 μm to 90 μm, and the thickness of the degradable polyester coating is in the range of 10 μm to 15 μm. The degradable polyester coating is formed by mixing polylactic acid (PLA) and polyglycolic acid (PGA) according to a mass ratio of 5 to 1, wherein the weight average molecular weights of the polylactic acid (PLA) and the polyglycolic acid (PGA) are 800,000 and 20,000 respectively, the polydispersity index of the mixture is 50, and the mass ratio of the iron-palladium alloy substrate to the degradable polyester coating is 15:1. The iron-based alloy stent was implanted into the abdominal aorta of a rabbit. The radial support force was tested by sampling, and the periphery of a stent strut was observed with a microscope at 2 months and 2 years after the date of implantation. The test results showed that the radial support force of the stent was 80 kPa, and no hydrogen bubbles were produced around the stent strut at 2 months after the date of implantation; the mass loss rate of the stent was 98 percent, and no hydrogen bubbles were produced around the stent strut at 2 years after the date of implantation. 
         [0077]    There were thickness differences between every part of each absorbable iron-based alloy stent in the preparation process, therefore, the coating thickness and the wall thickness of the iron-based alloy substrate were interval values in the Examples 1 to 16, and whether hydrogen bubbles were produced around the stent or not at predetermined observation points in time was observed with a microscope at the same magnification factor in each Example. 
       Control Example 1 
       [0078]    A bare pure iron stent (a pure iron substrate, uncoated with any coating on the surface) of which the wall thickness is between 60 μm and 70 μm was implanted into the abdominal aorta of a rabbit. After 3 months from the date of implantation, the stent was taken out, and the radial support force was 120 kPa by testing; at 3 years after the date of implantation, the stent was taken out to undergo a mass loss test, and the mass loss rate of the stent was 25 percent at the moment, indicating that the bare pure iron stent corroded slowly. 
       Control Example 2 
       [0079]    The surface of a bare pure iron stent (i.e., a pure iron substrate) of which the wall thickness is in the range of 60 μm to 70 μm was coated with a 25 μm to 35 μm thick polylactic acid (PLA) coating, wherein the mass ratio of the pure iron substrate to the polylactic acid (PLA) is 10:1, and the polylactic acid (PLA) has a weight average molecular weight of 15,000 and a polydispersity index of 1.8. An iron-based stent was obtained after drying. The iron-based stent was implanted into the abdominal aorta of a rabbit. After 1 month from the date of implantation, the stent was taken out and axially split, and the periphery of the support strut was observed with a microscope at the same amplification factor as those of the above-mentioned Examples; it was found that a large amount of hydrogen was produced around the stent strut in the corrosion process, and larger hydrogen bulges were formed as shown in  FIG. 7 , indicating that there was a relatively large risk of air embolism formation. The radial support force was 20 kPa by testing at 3 months after the date of implantation, and the mass loss test of the stent showed that the mass loss rate of the stent was 100 percent at 6 months after the date of implantation, indicating that the stent completely corroded, the corrosion was too rapid, and the clinical mechanical property requirement could not be met at predetermined points in time. 
         [0080]    It can be seen from the test results of the above-mentioned Examples 1 to 16 and Control examples 1 to 2 that the absorbable iron-based alloy stent provided by the present invention uses the degradable polymer with a weight average molecular weight in the range of 20,000 to 1000,000 and a polydispersity index of more than 1.0 and less than or equal to 50 to achieve a result that no hydrogen is produced or only a small amount of hydrogen is produced in 5 years after the iron-based alloy substrate is implanted into the body. i.e., the oxygen-consuming corrosion mainly occurs, and the internal corrosion rate is improved by means of oxygen-consuming corrosion, thereby overcoming the technical biases that the corrosion rate of an iron-based alloy can only be improved by hydrogen evolution corrosion in the degradation of the iron-based alloy and that the degradation speed of the iron-based alloy is not easily improved by increasing the oxygen-consuming corrosion rate in the prior art, and further avoiding the clinical air embolism risk brought by a large amount of hydrogen produced from the iron-based alloy substrate due to hydrogen evolution corrosion in the prior art. In addition, the mass loss rate of the absorbable iron-based alloy stent provided by the invention is not less than 90 percent in 5 years after the implantation into the body, thus meeting the clinical corrosion period requirements for the degradable stent; when the OCT follow-up was carried out, there was no obvious difference between the surrounding area of the stent and the surrounding area of the stent at the beginning of implanting, or the early radial support forces were all more than 23.3 kPa (175 mm mercury column), thus meeting the clinical early mechanical property requirements for the stent after the implantation into the body