Vascular stent comprising a magnesium alloy coated with poly(l-lactide) and polycaprolactone and manufacturing method thereof

Abstract

The present invention relates to a method for manufacturing a triple-layer magnesium alloy coated with poly(L-lactide) and polycaprolactone. It has excellent advantages as it was confirmed that magnesium coated with MgF2 / PLLA-HAp / PCL(C-Mg) triple layer has increased re-lining and decreased new lining compared to fluorine-coated samples and uncoated Mg samples, and that the decomposition rate of the magnesium alloy is reduced.

Description

Technology Field

[0001] The present invention relates to a method for manufacturing a magnesium alloy coated with poly(L-lactide) and polycaprolactone, and to the use thereof. Background Technology

[0002] Among cardiovascular diseases, coronary artery disease causes 18 million deaths worldwide. During various coronary treatments, such as angiography, coronary stent implantation provides intra-arterial support. For stents to be ideally used clinically, several fundamental advantages, such as mechanical strength, ductility, and biocompatibility, are required. Nevertheless, a significant number of patients experience late restenosis after stent implantation. Broad-spectrum anticoagulant therapy is also not useful for these patients.

[0003] Non-degradable exposed metal stents (BMS) possess excellent mechanical properties and biocompatibility, but they still cause late thrombosis and restenosis due to chronic inflammation and endothelial hyperplasia. Drug-eluting stents (DES) are also used; while they reduce thrombosis and restenosis in the early stages, they may recur after 1–2 years because they implant permanently in the vessel. Biodegradable stents are a promising class that can mitigate the disadvantages of BMS and DES stents. Magnesium and its alloys are suitable metals for implants due to their biodegradability and mechanical properties. Spontaneous degradation and reabsorption by vascular tissue after implantation are promising strategies. Despite the significant biocompatibility of Mg stents, their rapid degradation hinders widespread clinical application.

[0004] Surface modification via surface coating is considered an efficient technique to reduce degradation rates. Various surface coating technologies, such as inorganic coatings, chemical conversion, and organic coatings, are used to achieve corrosion resistance. For a coating to be effective, it must be biocompatible and reduce the corrosion rate to an acceptable level. Among the wide range of coatings, chemical conversion is considered an advantageous method that offers simplicity, cost-effectiveness, and efficiency. Chemical conversion, involving treatment with hydrofluoric acid to form magnesium fluoride (MgF2), reduces the corrosion rate and enhances the rapid degradation of Mg. Among polymer coatings, poly(L-lactide) (PLLA) and polycaprolactone (PCL) are biodegradable polymers with excellent biocompatibility and mechanical properties. Conventional PLLA has been reported to exhibit superior adhesion strength to Mg substrates compared to PCL. However, relying solely on polymer coatings can have disadvantages, such as acidic degradation products and unwanted capillary pores, which may limit the control of Mg alloy degradation. Acidic degradation products can also cause inflammation in the surrounding tissue environment of the implant. The hydrophobicity of polymer coatings can hinder cell adhesion. HAp particles possess high biocompatibility, and due to the inorganic ions present in HAp, the degradation products are inherently alkaline. The hydrophilic nature of HAp facilitates cell adhesion and proliferation. When used alone on an Mg substrate, the HAp coating may break due to its brittle nature, which can lead to reduced adhesion to the metal surface. Combining polymeric organic coatings with inorganic HAp can help overcome the disadvantages of each while combining their respective advantages. Inorganic alkaline degradation products can be neutralized by acidic polymeric products, and intermolecular interactions between the different polymer layers and HAp can increase the adhesion strength of the coating. The degradation rate of the coating can be improved by combining polymeric organic coatings with inorganic HAp to enhance bioactivity.

[0005] To address the issues of restenosis and thrombosis, stents are coated with polymer layers containing antidiffusing agents such as paclitaxel (PTX) and sirolimus. However, these antidiffusing drugs prolong endothelial healing. Prolonged endothelial healing is considered undesirable and has been reported to cause various negative clinical complications. Endothelial cells play a crucial role in maintaining the vascular hemostatic environment. Approaches are needed to inhibit maladapted neoendothelial proliferation and stimulate reendothelialization. Several attempts have been made to introduce such approaches. For example, loading anti-CD34 antibodies can improve endothelialization, but unfortunately, these antibodies do not specifically target endothelialization. Furthermore, they can lead to the proliferation of other cells, resulting in nephromedial proliferation. In this regard, vascular endothelial growth factor (VEGF) is receiving attention due to its ability to induce reendothelialization by stimulating the migration and maturation of endothelial progenitor cells (EPCs). Therefore, an approach combining the release of antidiffusional drugs such as Ptx and VEGF can not only effectively address the problem of reendothelialization but also promote it. Previously, an interesting approach combining the effects of vascular growth factors and antidiffusional drugs for non-biodegradable systems has been reported. In the case of biodegradable Mg-based stents, the degradation rate of the material must be controlled. Therefore, a multi-purpose coating that inhibits the degradation of Mg-based stents and simultaneously releases vascular growth factors and antiproliferative agents is a suitable approach.

[0006] The inventors have developed a Mg alloy-based biodegradable cardiovascular stent in which the degradation of the Mg alloy is controlled and the simultaneous release of a growth factor (VEGF) and a drug (PTX) occurs. To this end, it was anticipated that combining various coating layers could reduce the rapid degradation rate of Mg; thus, an optimized coating (MgF2 / PLLA-HAp / PCL) effectively reduced the degradation rate of Mg, and PTX and VEGF were loaded onto the final coating layer of PCL. It was confirmed that the simultaneous release of the growth factor and the drug resulted in a decrease in neoendothelium along with an increase in re-endothelialization, as described in in vivo studies. Prior art literature

[0007] Republic of Korea Published Patent No. 10-2009-0067871 (June 25, 2009) The problem to be solved

[0008] The object of the present invention is to provide a method for manufacturing a triple-layer magnesium alloy with increased re-re

[0009] Another objective of the present invention is to provide a stent comprising a triple-layer magnesium alloy with increased re-endothelialization and reduced neoendothelialization produced by the above manufacturing method.

[0010] Other objects and advantages of the present invention will become more apparent from the following detailed description of the invention, claims, and drawings. means of solving the problem

[0011] The present invention relates to a method for preparing magnesium via an MgF2 / PLLA-HAp / PCL(C-Mg) trilayer coating to control the degradation rate of Mg alloys through the loading of drugs and growth factors. In vitro results indicate that the degradation rate was significantly reduced by this combination of coating layers, and that loading with a drug (PTX) and a growth factor (VEGF) increased reendothelialization and reduced in vivo neoendothelial (neoendothelial) formation. In vitro biocompatibility results indicate that the sample is suitable for in vivo transplantation. In vivo transplantation implied that PTX loading reduced neoendothelial formation by regulating smooth muscle cell proliferation, while VEGF increased the reendothelialization process by regulating inflammation. These results indicate that the degradation rate of magnesium alloys is simultaneously controlled, and neoendothelial formation and reendothelialization are controlled.

[0013] The present invention will be described in more detail below.

[0015] The present invention provides a method for manufacturing a triple-layer magnesium alloy with increased re-lining and reduced new lining, comprising the steps of: polishing a magnesium alloy; treating the polished magnesium alloy with a fluoride to produce a fluorine coating layer on the magnesium alloy; and coating the fluorine coating layer with a poly(L-lactide) / polycaprolactone solution to form a poly(L-lactide) and polycaprolactone layer.

[0016] The method for manufacturing a triple-layer magnesium alloy of the present invention, in which re-heathering is increased and new heathering is reduced, includes the step of polishing the surface of the magnesium alloy.

[0017] In the present invention, the magnesium alloy may be an alloy further comprising one or more metals selected from the group consisting of zinc, manganese, calcium, zirconium, yttrium, molybdenum, niobium, tantalum, titanium, strontium, chromium, silicon, phosphorus, nickel, and iron. The sum of the weights of the metals other than magnesium may be less than 10 weight percent of the total alloy, but is not limited thereto.

[0018] The method for manufacturing a triple-layer magnesium alloy with increased re-lining and reduced new lining according to the present invention includes the step of treating a polished magnesium alloy with fluoride to produce a fluorine coating layer on the magnesium alloy.

[0019] In the present invention, the fluoride may be hydrofluoric acid.

[0020] In the present invention, the magnesium alloy may have a reduced decomposition rate.

[0021] In the present invention, to fluorine-treat Mg, a prepared sample was immersed in 48 wt% hydrofluoric acid (HF, Sigma Aldrich, USA) at room temperature for 24 hours while continuously stirring, washed with distilled water, and vacuum dried to produce a fluorine-coated magnesium alloy, which was named MgF2.

[0022] The method for manufacturing a triple-layer magnesium alloy of the present invention, which has increased re-lining and reduced new lining, includes the step of coating the fluorine coating layer with a poly(L-lactide) / polycaprolactone solution to form a poly(L-lactide) and polycaprolactone layer.

[0023] In the present invention, the poly(L-lactide) and polycaprolactone solution can be prepared by dissolving poly(L-lactide) in dichloromethane.

[0024] In the present invention, the poly(L-lactide) and polycaprolactone layers may be PLLA-HAp / PCL(C-Mg).

[0025] In one embodiment of the present invention, the pH change of MgF2, PLLA-HAp, and PCL(C-Mg) coatings was analyzed, and it was confirmed that when a PLLA-HAp layered sample was coated with a PCL polymer, the pH change was significantly reduced.

[0026] In one embodiment of the present invention, the degradation rates of MgF2, PLLA-HAp, and PCL(C-Mg) coatings were analyzed, and it was confirmed that PCL(C-Mg), a triple-layer coating sample, exhibited the lowest degradation rate.

[0027] In one embodiment of the present invention, the amount of VEGF loaded onto samples coated with MgF2, PLLA-HAp, and PCL(C-Mg) was analyzed, and it was confirmed that the PCL(C-Mg) triple-layer coated sample showed the lowest degradation rate.

[0028] In one embodiment of the present invention, the VEGF release of MgF2, PLLA-HAp, and PCL(C-Mg) coatings was analyzed, and it was confirmed that Mg-C / PTX-VEGF and Mg-C / VEGF samples exhibited similar VEGF release profiles.

[0029] In one embodiment of the present invention, the VEGF release of MgF2, PLLA-HAp, and PCL(C-Mg) coatings was analyzed, and it was confirmed that Mg-C / PTX-VEGF and Mg-C / VEGF samples exhibited similar VEGF release profiles.

[0030] The present invention confirmed that compared to magnesium alloys with a fluorine coating alone or uncoated magnesium alloys, magnesium alloys with a triple layer of MgF2 / PLLA-HAp / PCL(C-Mg) coating showed increased re-innervation and reduced new innervation, which is superior.

[0031] In addition, the present invention provides a stent comprising a triple-layer magnesium alloy with increased re-endothelialization and reduced neoendothelialization produced by the above manufacturing method.

[0032] Meanwhile, since the corresponding features in the above details can be substituted in the section described above, their description is omitted. Effects of the invention

[0033] The present invention relates to a method for manufacturing a triple-layer magnesium alloy coated with poly(L-lactide) and polycaprolactone. It has excellent advantages as it was confirmed that magnesium coated with MgF2 / PLLA-HAp / PCL(C-Mg) triple layer has increased re-lining and decreased new lining compared to fluorine-coated samples and uncoated Mg samples, and that the decomposition rate of the magnesium alloy is reduced. Brief explanation of the drawing

[0034] Figure 1 is a schematic diagram showing the mesh size. Figure 1 shows the dimensions of the stent mesh. Since the mesh size is too small, it is a schematic diagram to show the dimensions. FIG. 2 is a schematic diagram of a stent-shaped rolled mesh, showing the front (A) and rear (C) sides. FIG. 2 illustrates a method of forming the mesh into a stent shape. FIG. 2a shows the upper layer of the stent, and FIG. 2b shows the lower layer of the rolled stent. FIG. 2c shows a cross-sectional image of the mesh. Figure 3 shows a cross-sectional interface of a Mg-C sample with A) SEM, B) AFM analysis of Mg, MgF2, MgF2 / PLLA-Hap, and C-Mg, and C) EDS mapping profile. Figure 3 evaluates the surface morphology of coated and uncoated magnesium samples. The cross-sectional images of the additionally coated magnesium show the elements of each coating layer. The coating layers of the magnesium alloy can be identified through the cross-sectional images. Figure 3A is the initial SEM image showing the magnesium alloy, MgF2, MgF2 / PLLA-Hap, and C-Mg supports. According to the SEM analysis, it can be confirmed that the surface morphology changes after coating formation. Figure 3B is an AFM image confirming the surface morphology through a 3D image. Figure 3C shows a cross-sectional SEM image of the magnesium sample after coating, which also includes the elemental analysis of each coating layer. Figure 4 shows photographs of samples used for cell compatibility evaluation. Figure 4 shows (A) Mg, (B) MgF2, and (C) Mg-C. This indicates that PCL formed a transparent coating film. Figure 4a shows the Mg alloy sample after polishing. Figure 4b shows a photograph of the MgF2-coated sample. Figure 4c shows the polymer-coated sample. These photographs demonstrate that the polymer coating was applied transparently. Figure 5 is an SEM image showing the coating thickness of an MgF-coated Mg sample. The coating thickness was measured to be 0.61 μm. Figure 5 is an SEM cross-sectional view of an Mg-C sample showing the coating thickness of an internal MgF2 sample. Figure 6 shows A) XRD, B) contact angle measurements of Mg, MgF2, MgF2 / PLLA-Hap, and Mg-C coated samples, C) chemical structural formulas of MgF2 / PLLA / PCL and D) MgF2 / PLLA-HAp / PCL, and E) adhesion strength of MgF2 / PLLA / PCL and MgF2 / PLLA-HAp / PCL (****P < 0.0001 and ns = not significant). Figure 6A shows the XRD analysis data of the coated surface. Coating formation is confirmed by differences in XRD patterns. Figure 6B shows the contact angle measurements of the coated material. Surface hydrophobicity or hydrophilicity can be confirmed through contact angle measurements. The coated material has a low contact angle, indicating that cells can adhere to the surface of the support. Figures C and D show the chemical structure of the coating layer. Figure C indicates the absence of HAp, while Figure D indicates the presence of HAp. Adhesive strength can be increased by adding HAp. This was confirmed through the scratch test in Fig. 6E. Figure 7 shows A) evaluation of degradability, B) measurement of pH change, C) evaluation of hydrogen gas generation of Mg, MgF2, MgF2 / PLLA-Hap, and Mg-C supports, and D) a schematic representation of the evaluation of degradability of Mg and coated samples (****P < 0.0001, **P < 0.01). Figure 7A shows the evaluation of the degradability of the supports. The Mg support exhibits the fastest degradability compared to other supports. It was confirmed that the degradability of Mg was significantly reduced by the coating treatment. Figure 7B shows the pH change of each sample. It was confirmed that the pH change of the coated sample was smaller than that of other samples. Figure 7C shows the amount of hydrogen gas generated. The amount of hydrogen gas generated in the coated sample was measured to be lower than that of Mg. Figure 7D shows a representative schematic diagram. The degradability of the coated sample was measured to be lower compared to Mg. Figure 8 is a graph showing that the coating adhesion of PLLA increased due to the coating of HAp particles. When only a polymer coating is present, the coating is not very strong. To prove this, a comparative analysis was conducted. In Figure 8a, the degradation power of the Hap-coated sample was measured to be lower compared to the polymer-coated sample. In terms of pH and hydrogen gas generation evaluations compared to the polymer-coated sample, the sample containing HAp particles showed better results. Figure 8 evaluates the effect of the presence of HAp on the degradation rate through A) degradation, B) pH change, and C) hydrogen gas generation of Mg-C and MgF2 / PLLA / PCL. When HAp is present in the coating, the degradation power became significantly slower compared to the polymer-coated sample, which exhibited strong adhesion. Figure 9 shows the SEM observation results of the decomposed samples after 14 days. (A) Mg, (B) MgF2, (C) MgF2 / PLLA-HAp, and (D) Mg-C. Figure 9 shows the SEM images of the EDS analysis after sample decomposition. This image indicates that the Mg and MgF2 samples have many cracks and can decompose rapidly. Here, the Mg-C sample had small cracks, and due to this characteristic, the decomposition rate was measured to be the lowest. Figure 10 shows A) a schematic representation of the Mg-C / PTX-VEGF sample; B) SEM image of Mg-C / PTX; C) SEM image of Mg-C / VEGF; and D) SEM image of Mg-C / PTX-VEGF. E) FT-IR analysis; and F) XPS analysis results of Mg, MgF2, MgF2 / PLLA-HAp, Mg-C, and Mg-C / PTX-VEGF. Figure 10A is a schematic diagram of growth factors and drugs loaded onto a PCL layer. Figures B, C, and D are SEM images showing changes in surface morphology after loading growth factors and drugs. Figure 10E shows the FT-IR spectra of each scaffold. Figure 10F shows the XPS evaluation of the drug-loaded coated scaffold. Figure 11 shows XPS analysis data. In the chemical characterization performed in the present invention, peaks of Zn, Ca, P, and N were identified. Figure 11 shows the high-resolution XPS spectra of (A) Zn2p, (B) Ca2p, (C) P2p, and (D) N1s peaks of Mg alloy, MgF2 / PLLA-HAp, and Mg-C / PTX-VEGF, respectively. Figure 12 shows A) VEGF loading amount, B) immunofluorescence images showing the VEGF distribution after coating, B1) Mg-C / VEGF, B2) Mg-C / PTX-VEGF, C) cumulative PTX release, and D) cumulative VEGF release results. Figure 12A shows the amount of growth factor loaded onto the sample. Figure 12B is an image of immunofluorescence staining of the growth factor loaded onto the coated sample. Figures 12A and 12B are graphs showing the cumulative release amounts of the drug and growth factor. The drug release graph indicates that it is being released continuously. Figure 13 shows SEM images representing platelet adhesion: A) Mg, B) Mg-C, C) Mg-C / VEGF, D) Mg-C / PTX, E) Mg-C / PTX-VEGF (yellow arrows indicate platelets). F) Schematic representation of platelet activation for Mg alloy and Mg-C samples. Figure 13 shows SEM images of platelet activation. In the case of Mg, it was confirmed that platelets were activated. However, in the case of the coated scaffold and the drug / growth factor-loaded scaffold, platelets were not activated. This implies that the likelihood of thrombosis is reduced after coating. Figure 14 shows A) observation of cell proliferation via fluorescence microscopy, B) in vitro cell viability test, C) expression of cell proliferation in Mg and coated Mg, and D) hemolytic rates of Mg, Mg-C, Mg-C / PTX, Mg-C / VEGF, and Mg-C / PTX-VEGF (****P < 0.0001, **P < 0.01, *P < 0.05). Figure 14 shows the results of the biocompatibility evaluation of the scaffolds. Figure 14A shows cell proliferation on the coated Mg scaffold. Drug loading and growth factors appear to have no adverse effects on cell growth. Figure 14B shows the cell viability of cells cultured on the coated scaffold compared to Mg. The Mg scaffold exhibits low biocompatibility due to a high degradation rate, a basic pH, and hydrogen gas generation. Figure 14C schematically illustrates the difference in biocompatibility between Mg and the coated scaffold. Figure 14D shows the hemolysis rate of the support. A hemolysis value of less than 5% is considered safe for animal use. The coated sample has low hemolytic activity compared to Mg. Figure 15 shows A) a schematic diagram of the surgery for transplantation, B) a schematic diagram of the stent-shaped Mg alloy and Mg-C scaffolds, C) a schematic representation of Mg-C and Mg-C / PTX-VEGF, and D) H&E images of Mg-C and Mg-C / PTX-VEGF scaffolds extracted after 1 and 2 months. Figure 15 shows the in vivo results. Figure 15A schematically illustrates the surgical method. Figure 15B shows the coated Mg stent and Mg stent used in cell experiments. Figure 15C shows the mechanism of neointima formation in Mg and coated scaffolds. Figure 15D shows H&E histological staining data of scaffolds extracted from mice after 1 and 2 months. It indicates that the Mg scaffold showed higher neointima formation compared to the coated scaffold. Figure 16 shows the immunohistological evaluation of artificial arteries fabricated with stents at 1 and 2 months post-implantation, including A) α-SMA, B) CD-68 and a graph of the positive region, C) α-SMA positive region, and D) CD-68 positive region. (**P < 0.01, *P < 0.05, ns = not significant). Figure 16 presents the immunohistological analysis of artificial arteries fabricated with stents extracted 1 and 2 months after implantation. Figure 16A confirmed that α-SMA was enhanced in the Mg scaffold due to high neointimal formation. Figure 16B shows CD-68, an inflammatory marker. Inflammation was measured to be higher in the Mg scaffold. Figures 16C and D show the positive regions for each marker. Calculating the positive regions for these markers is important in the study. Figure 17 shows the immunohistological evaluation of artificial arteries fabricated from stents extracted 1 and 2 months after transplantation, including A) FLKI, B) ICAM and a graph of the positive area, C) FLKI-positive area, and D) ICAM-positive area (ns = not significant). Figure 17 is the immunohistological evaluation of the endothelial markers FLKI and ICAM. As can be seen in the staining and images, re-endothelialization occurred more frequently in the coated scaffolds compared to the Mg scaffold. This is due to growth factor loading. Additionally, as can be seen in the stained slides, re-endothelialization of the coated Mg sample is very smooth, whereas in the case of Mg, it appears very abruptly. Specific details for implementing the invention

[0035] Hereinafter, the present invention will be described in detail with reference to examples to aid in understanding. However, the following examples are merely illustrative of the content of the present invention and the scope of the present invention is not limited to the following examples. The examples of the present invention are provided to more completely explain the present invention to those with average knowledge in the art.

[0037] 1. Experimental Materials and Methods

[0039] 1.1. Sample Preparation

[0041] In this invention, an Mg alloy (ZK-60) was used. ZK-60 was supplied by the Korea Institute of Materials Science (KIMS, Changwon, South Korea). The average hardness of ZK-60 was 63.6 Hv, the elastic modulus was 12.5 GPa, the compressive yield strength (CYS) was 146.1 MPa, and the maximum compressive strength (UCS) was 442.9 MPa. First, it was cut into thin sheets of 10 mm x 12 mm x 0.5 mm. The sample was first polished with 800 to 1500 grit silicon carbide paper. This maintained a uniform surface roughness of the sample. The sample was ultrasonically treated with acetone for 10 minutes and then washed with distilled water. Subsequently, the sample was dried at room temperature to remove residual particles from the surface of the Mg alloy. Afterward, the sample was treated with 48 wt% hydrofluoric acid (Sigma Aldrich, USA) at room temperature for 24 hours, and the hydrofluoric acid (HF) treated sample was washed appropriately with distilled water and vacuum dried to remove moisture from the surface.

[0042] After the fluoride coating was formed, a polymer layer was formed by dip coating. Before the dip coating process, the sample was heated at 50 °C to ensure the removal of trapped water molecules.

[0043] To form a uniform coating, the sample is 1 mm -1 The sample was mechanically extracted from the polymer solution at a constant rate. The sample was immersed in the polymer solution for 45 seconds. Afterward, it was suspended vertically in a vacuum drying oven (0.1 MPa, SAMHEUNG) at 25°C for drying. This process was repeated three times. Subsequently, after the three dip coatings were completed, the sample was vacuum dried for one day. The samples were named according to the polymer layer (Table 1).

[0045] SampleNo. SampleName Coating layer Layers description 1 Mg non Uncoated Mg Sample 2 MgF2 MgF2 1. Magnesium fluoride coating 3 MgF2 / PLLAHAp MgF2 / PLLA-HAp 1. Magnesium fluoride coating2. PLLA and HAp coating 4 Mg-C MgF2 / PLLA-HAp / PCL 1. Magnesium fluoride coating2. PLLA and HAp coating3. PCL coating 5 Mg-C / PTX MgF2 / PLLA-HAp / PCL-PTX 1. Magnesium fluoride coating2. PLLA and HAp coating 3. PCL containing PTX coating 6 Mg-C / VEGF MgF2 / PLLA-HAp / PCL-VEGF 1. Magnesium fluoride coating2. PLLA and HAp coating 3. PCL containing VEGF coating 7 Mg-C / PTXVEGF MgF2 / PLLA-HAp / PCL-PTX-VEGF 1. Magnesium fluoride coating2. PLLA and HAp coating 3. PCL containing VEGF and PTX coating

[0046] Table 1. Coating composition of the samples

[0048] 1.2. Coating Solution

[0050] To form the PLLA layer, a 2% PLLA (Poly(L-Lactide) mol wt. 85,000-160,000 Sigma Aldrich, USA) solution was prepared in dichloromethane (Sigma-Aldrich, USA). The solution was stirred for 4 hours to form a homogeneous mixture.

[0051] A PLLA-HAp solution was formed by dissolving 2% PLLA polymer in dichloromethane at room temperature. The mixture was stirred with a magnetic beater for 4 hours. 0.1% Hydroxyapatite Hap (particle size: 10-40 nm) was added to this solution and sonicated for 30 minutes to ensure a uniform distribution of HAp particles within the solution. Typically, a 2% PCL (polycaprolactone, Mn-70,000-90,000, Sigma-Aldrich, USA) solution was formed according to the same procedure as described for the PLLA solution.

[0053] 1.3. Drug Loading

[0055] 0.01 gml PTX (Paclitaxel, Sigma-Aldrich, USA) -1 It was dissolved in a 2% PCL solution in dichloromethane of a certain concentration.

[0057] 1.4. VEGF Loading

[0059] First, 0.25 ml VEGF solution (5 μg ml -1 An oil-in-water emulsion was prepared by magnetically stirring ) and 0.05 ml of Span 80 (Sigma-Aldrich, USA) in dichloromethane for 3 hours. Subsequently, 2% polycaprolactone was added, and the resulting emulsion was stirred at room temperature for 12 hours.

[0061] 1.5. Coating Characteristics

[0063] The coating morphology of the samples was determined using a scanning electron microscope (SEM) (JSM-635F, JEOL, Japan) and an energy-dispersive X-ray (EDX) spectrometer (Oxford Instruments, UK). Prior to SEM, the samples were sputter-coated with platinum (Cressington 108 Auto, JEOL, Japan) for conductivity. An acceleration voltage of 10 kV was used to obtain SEM images. A secondary electron detector was used to record all SEM images. Atomic force microscope (AFM, XE) in contact mode -1 The coated surface was analyzed using a 50, PSIA. Fourier Transform Infrared (FTIR) analysis was performed using a Nicolet iS10-Smart iTR (Thermo Fisher Scientific, USA) to characterize the functional groups of the coating. Typically, in transmittance mode, wavenumbers 4000–800 cm⁻¹ -1 4cm in the range -1 Thirty-two scans were used at a resolution of [resolution missing]. The FTIR instrument is equipped with OMNIC software. Data were visualized using Origin 8 software. The surface was analyzed using XPS (Thermo Scientific K-Alpha X-ray Photoelectron Spectrometer) with an AlKα radiation source (hγ = 1486.6 eV). The spot size of the irradiation scan was 1000 μm.

[0064] Phases present in Mg alloys and multilayer coatings were analyzed using Cu Kα radiation with an X-ray diffraction analyzer (XRD, D / MAX-250, Rigaku, Japan, operating at 40 KV and 30 mA). Samples were analyzed with 2θ scans from 20 to 80 at a scan rate of 2 / min.

[0065] MgF using a scratch tester (Anton Paar Micro Combo Tester, Austria) 2 / Adhesive strength was compared by measuring the adhesive scratch of PPLA / PCL and MgF2 / PPLA-HAp / PCL coatings. A progressive load of 10 mN–20,000 mN was applied at 10 μms -1 It was applied with a table speed of 1000 μm and a scanning length of 1000 μm. The radius of the Rockwell diamond ball was 100 μm.

[0066] The hydrophilicity of the coated samples was measured using a contact angle measurement system (DSA 100, KRUSS GmbH). 10 μl particle volume and 1 μls -1 Ultrapure water with a dosage range was dropped onto the surface of each sample. An average of 10 readings were taken for each sample. Each set of measurements was repeated 3 times.

[0068] 1.6. Decomposition into In-Bit

[0070] To determine the effect of the coating on the degradation rate of the samples, immersion tests were performed according to the ASTM G31-72 standard protocol. The initial weight (W0) of the samples was measured prior to immersion. Each sample was immersed in 5 ml of PBS (Amresco, Korea, pH -7.4) solution and maintained at 37°C. Three samples were used for the immersion test in each set. Samples were taken from the solution at different time intervals to determine weight loss. Changes in the surface morphology of the samples after different time intervals were observed via SEM and EDS analysis. Before measuring weight loss, each sample was washed with chromic acid for 5 minutes to remove residual corrosion products. Subsequently, the samples were washed with distilled water and dried at room temperature before measuring the weight (WL). The degradation rate was determined according to the following equation.

[0072]

[0073] WF, W0, and WL are the weight loss rate, weight before immersion, and weight after immersion, respectively.

[0075] Changes in pH over time due to the decomposition of the Mg alloy were also recorded. For this purpose, three samples from each set of coated Mg alloys were incubated in PBS at 37°C. Changes in pH were recorded at specific time intervals for 14 days using a pH meter (Thermo Scientific, Korea). The PBS solution was replaced with fresh solution daily. Hydrogen gas generation was recorded by immersing samples from each set in PBS at 37°C using an inverted graduated test tube. Changes in the solution level in the test tube were recorded daily for 14 days. Three samples from each set were used to measure hydrogen gas generation.

[0077] 1.7. Drug Release

[0079] To measure the loaded amount of PTX, paclitaxel-coated samples were immersed in 5 ml of DMC (dichloromethane, Sigma-Aldrich, USA) and vortexed for 3 hours. The resulting solution was evaporated to remove the solvent. Subsequently, 10 ml of PBS:methanol (70:30) (Sigma-Aldrich, USA) was added, and the absorbance for PTX was measured at 230 nm using a Biodrop spectrophotometer (Biochrom, UK). The absorbance values ​​were for PBS:methanol (70:30) (0.2–100 μg / ml). -1 The absorbance value of PTX was measured and converted using a calibration curve determined from the measurement. The calibration curve showed a linear relationship between the PTX concentration and the absorbance value.

[0080] To determine PTX release over time, PTX-loaded samples were immersed in 3 ml of PBS at 37°C for 24 days. The PBS solution was regularly changed to a fresh PBS solution. To determine the absorbance value of PTX in the PBS solution, methanol was added while maintaining a PBS:methanol ratio of 70:30. After adding methanol, the solution was vortexed and the absorbance was measured.

[0082] 1.8. VEGF Release

[0084] VEGF loading was confirmed by immunohistochemistry. The amount of VEGF loaded in the samples was measured by immersing the samples in RIPA lysis buffer (Radioimmunoprecipitation assay buffer, Thermo-Fisher Scientific, South Korea) at 4°C for 24 hours. The amount of VEGF was subsequently determined using a VEGF ELISA kit (R&D systems, USA) according to the manufacturer's instructions. VEGF release was determined by immersing the samples in 3 ml of PBS at 37°C for 24 days. VEGF release was determined by a VEGF ELISA kit.

[0086] 1.9. Platelet Adhesion and Hemolysis

[0088] Blood was obtained from rats (Sprague-Dawley) and stored in an anticoagulant. The anticoagulated blood was centrifuged to obtain Platelet-rich Plasma (PRP) (10 min, 2000 rpm). The PRP layer was carefully separated. The separated PRP was applied to the coated sample and incubated at 37°C for 1 hour. After incubation, the sample was washed with PBS and fixed with 3% glutaraldehyde. Different concentrations of ethanol solution (40–100%) were applied for dehydration. After drying, the sample was observed using SEM.

[0089] The hemolysis test was performed by immersing each sample in 5 ml of PBS and incubating at 37°C for 30 minutes. 0.1 ml of diluted blood was added to each sample. Subsequently, the samples were incubated at 37°C for 1 hour. For the positive and negative protocols, distilled water and PBS (5 ml) were used, respectively. After incubation, the samples were centrifuged at 5000 rpm for 5 minutes. The supernatant was removed, and the absorbance was measured at 545 nm. The hemolysis rate was calculated using the following formula.

[0091]

[0092] Ht is the absorbance value of the sample, and Hpc and Hnc are the absorbances of the positive and negative controls, respectively.

[0094] 1.10. In-vitro biocompatibility

[0096] In the present invention, small pulmonary artery endothelial cells (CPAE, the American Type Culture Collection, USA) were used. CPAE cells were cultured in a cocktail of minimal essential medium α-MEM (Gibco, USA), 20% fetal bovine serum, FBS (Sigma Aldrich, USA), and 1% penicillin-streptomycin, PS (Bio-Whittaker, USA) and placed in an incubator at 37°C with 5% CO2.

[0097] Indirect cell analysis was used to measure cell viability using the ISO 10993-12 protocol. All sets of samples were first sterilized with UV radiation for 40 minutes. Each sample was immersed in α-MEM medium at 37°C for 1 day to obtain a 100% extract. After culturing for 1 day, the supernatant was removed and centrifuged. The extracts were further diluted to different concentrations (12.5%, 25%, and 50%), and cytotoxicity was evaluated. CPAE cells (1 x 10⁻¹⁰ 4Cells (1 / ml) were seeded into a 96-well plate and incubated at 37°C for 1 day. After 24 hours, the cell medium was decanted and replaced with sample extracts of different concentrations. Medium without sample extracts was used as a control. The 96-well plate containing the medium extracts and cells was incubated at 37°C for 1 day. 100 μl of MTT [3-(4,5-dimethylthiazole-2-yl)-2,5-diphenyltetrazolium bromide] (Gibco, CA) solution was added to each well and incubated at 37°C for 4 hours. Finally, 100 μl of DMSO (dimethyl sulfoxide, Samchun Chemical, South Korea) was added to each well. After 1 hour of incubation, absorbance was recorded at 595 nm using an ELISA plate reader (EL 312, Biokinetics reader, Bio-Tek instrument).

[0098] Cell proliferation behavior was observed in CPAE cells (1 x 10⁻¹⁰) in a 24-well culture plate in cell culture medium for 24 hours. 4 cells / ml) were seeded and observed. After 24 hours, the cell culture medium was replaced with a medium containing a 100% extract of the coated sample. Cell-seeded culture plates were cultured for 1, 3, and 7 days at 37°C in a humid atmosphere of 5% CO2. For each sample, the cell-seeded culture plates were washed three times with PBS and fixed with 4% paraformaldehyde (Sigma Aldrich, USA) for 10 minutes. Subsequently, the samples were permeated by adding 0.5% Triton X-100 (Sigma Aldrich, USA) for 10 minutes. The samples were blocked using 2.5% BSA (bovine serum albumin, Sigma-Aldrich, USA) at room temperature for 1 hour. After blocking with BSA, the cell membrane was FITC (fluorescein isothiocyanate conjugated phalloidin, 25 μg / ml -1After immunostaining with Hoechst (Sigma Aldrich, USA) solution, the mixture was incubated at 4°C for 12 hours. 10 μg / ml -1 Nuclei were stained using a Sigma Aldrich, USA. Subsequently, samples were analyzed using a confocal microscope (Olympus, FV10i-W, USA) with FV10i-ASW2.0 software.

[0100] 1.11. In-vivo Biocompatibility

[0102] To evaluate biocompatibility, in vivo transplantation was performed on male rats (Rattus norvegicus, 12 weeks old, body weight 200–250 g) purchased from Dayun (Animal Center). Animal experiments were conducted according to the standard protocol of the Animal Ethics Committee of Soonchunhyang University, Korea (Approval No. SCH22-0041), in compliance with the EU Directive 2010 / 63 / EU on Animal Experimentation and the National Research Council's Guide for the Care and Use of Laboratory Animals. In accordance with the guidelines, the animals were fed daily (Zeigler Bros., Inc.) and monitored to ensure they were adapting normally, were healthy, and consumed sufficient food and / or water to maintain their health. The animals were kept in separate cages. The rats were divided into two groups: Mg-C stents and Mg-C / PTX-VEGF stents. For the in vivo experiments, laser-cut Mg mesh was manually rolled into the stent to mimic the shape of the stent (Figs. 1–2). A stent was implanted in the abdominal aorta of rats. Isoflurane (Piramal Critical Care Inc., USA) was used to anesthetize the rats. For the stent implantation, a midline laparotomy was performed first to isolate the abdominal aorta. Both sides of the aorta (diameter 2-2.5 mm) were beveled to avoid bleeding. Subsequently, the incision was made, and the midline incision on the arterial side was sutured (intermediate suture using 9-0 monofilament nylon suture ETHILON, ETHICON, USA) to implant a stent (diameter 1.5 mm, length 5 mm). To close the abdominal wound, continuous suturing was performed using 4-0 monofilament nylon suture (SURGIFIT, Ailee Company Limited, Korea). To prevent infection and minimize pain, the antibiotic Baytril (Bayer Co., Korea) and the analgesic Maritrol (Jeil Pharmaceutical, Korea) were administered for 3 days. After 1 and 2 months, rats were sacrificed (n=3 for each group at each time point).All samples were fixed in paraformaldehyde. CO2 inhalation was used to sacrifice rats. After fixation, samples were dehydrated with a series of ethanol solutions followed by dehydration using a xylene series. Paraffin wax (Leica Biosystem, Germany) was used for the embedding process. After sectioning by microtome (5 ± 2 μm thickness), deparaffinization was performed while subsequently staining with hematoxylin and eosin (H&E). Immunohistochemical staining was performed using a DAB staining kit (3'-Diaminobenzidine, Agilent Technologies, US). Deparaffinized, rehydrated, and blocked samples were stained with primary antibodies (Anti-CD68, Anti-alpha smooth muscle actin, Anti-ICAM). -1 , and Anti-FLKI -1 The samples were treated overnight at 4°C. The samples were treated with HRP-conjugated secondary antibodies for 1 hour, and the antigen-antibody reaction was determined by applying DAB chromosomes. The samples were observed and photographed using a microscope (BX53 Olympus) and a camera (DP72 Olympus), respectively. Image J software was used for the quantification of the immunohistochemically stained samples.

[0104] 1.12. Statistical Analysis

[0106] GraphPad Prism version 8.0 (GraphPad Software, Inc., USA) was used to perform statistical analysis. For the comparative analysis of different groups t - Tests and two analyses of variance (ANOVA) were used. Confidence levels are as follows, unless otherwise noted. p < It was considered 0.05. Unless otherwise specified, 3 times ( n =3) The experiment was performed.

[0108] 2. Experimental Results

[0110] 2.1. Coating Characteristics

[0112] Mg alloy-based stents are a type of biodegradable stent; however, the high degradation rate of Mg is a limitation to its widespread applicability despite its many advantages. A second problem encountered in the use of cardiovascular stents is restenosis and thrombosis within the stent. Various coating layers have been used to address the rapid degradation of Mg metal. A chemically treated MgF2 layer provided an anchor for the additional layer. PLLA possesses a greater number of oxygen atoms compared to PCL, enhancing electrostatic intermolecular interactions to allow PLLA to adhere more strongly. Therefore, the second layer consisted of PLLA-HAp. The HAp binding helped improve biocompatibility and increase adhesion strength. PCL was used to lock in the uneven surface created by the HAp addition and to load the drug (PTX) and the growth factor VEGF.

[0113] SEM images along with EDS mapping are shown in Figs. 3A and 3C. The surface morphology of the exposed Mg alloy exhibited sharp vertical lines (Fig. 3A). As the magnesium fluoride coating formed, the initial metallic color of the Mg alloy changed from gray to black (Fig. 4). Bonding HAp to the PLLA layer roughened the surface and resulted in a uniform distribution of particles. The PCL coating laminated over the PLLA-HAp coating appeared as a highly porous transparent film (Fig. 4), as depicted in the SEM image. Using the PCL coating resulted in a smoother surface morphology compared to the PLLA-HAp coating layer. The observed vertical curves were due to the polished Mg surface. Although there were three coating layers, the total thickness of the coating was very small to completely cover the vertical lines generated during the polishing process. Fig. 3B shows an AFM image of the coated surface. Surface analysis revealed rough peaks after the formation of the PLLA-HAp coating on MgF2, where HAp was bonded to the PLLA. After the PCL coating, these peaks became smooth. The surface roughness found for Mg, MF2, MgF2 / PLLA-Hap, and C-Mg is 53.29 nm, 50.62 nm, 56.03 nm, and 38.96 nm, respectively. A rough surface provides a larger surface area for corrosive attack. The degradation rate can be reduced by reducing the degradation surface roughness.

[0114] Figure 3C shows a cross-sectional image of the coating using an Mg alloy base. It can be observed that the coating is consistent and free of cracks. EDS data revealed partial concentrations of Zn in the Mg alloy. The MgF2 coating layer was confirmed by the presence of F in the EDS mapping. The thickness of the internal MgF2 coating was approximately 0.61 μm (Figure 5). The presence of O and C in both coating layers confirmed the formation of PLLA and PCL coatings. HAp particles could be detected by the presence of P and Ca. .

[0115] Meanwhile, the thicknesses for Mg, C-Mg, and Mg-C / PTX-VEGF were summarized as follows.

[0116] Sample Name Thickness (μm) Mg 103.76 C-Mg 117.41 Mg-C / PTX-VEGF 118.55

[0117] Table 2: Thickness of Mg, C-Mg, and Mg-C / PTX-VEGF

[0119] The XRD diffraction pattern of the Mg alloy shows a specific peak for Mg (ICDD No.: 01-071-6543) (Fig. 6A). The XRD pattern of the MgF2-coated sample shows the presence of peaks mainly corresponding to a tetragonal structure (ICDD No.: 01-071-6543). Due to the thin coating, the MgF2 peak exhibited very low intensity compared to Mg. At a test angle of 1°C, the depth of the CuKα line for Mg is approximately 0.8 μm. In the PLLA-HAp-coated sample, a characteristic peak for HAp was observed (ICDD No.: 01-074-0566), indicating the presence of HAp particles in the PLLA layer. Additionally, after PCL coating, characteristic broad peaks appeared at 21°C and 24°C, indicating the amorphous nature of the polymer coating. The Mg peak predominated in all samples, indicating the thin nature of the protective coating.

[0120] Changes in the wettability of the samples and photographs of water droplets on the surface are illustrated in Fig. 6B. The contact angles of Mg, MgF2, MgF2 / PLLA-HAp, and Mg-C were 106.49 ± 1.72˚, 102.8667 ± 0.98˚, 86.7 ± 0.85˚, and 86 ± 1.1˚, respectively. Compared to clean Mg, each coating reduced the contact angle, indicating increased hydrophilicity. Adhesive interactions between the water droplet and the sample surface increased with increasing hydrophilicity, resulting in a decrease in the contact angle. Polymer coatings PLLA and PCL possess hydrophobic properties that hinder cell attachment to the Mg surface. Incorporating HAp into the polymer coating helps increase the hydrophilicity of the composite coating. Increased hydrophilicity may be indirectly related to increased biocompatibility, cell adhesion, and proliferation.

[0121] Figure 6E shows the scratch test results for the PLLA / PCL coating and the PLLA-HAp / PCL coating of MgF2. The results show that the adhesion strength of the PLLA-HAp / PCL coating is higher than that of the PLLA / PCL coating. Figures 6C and D show the chemical formation of the two coatings. After the bonding of HAp particles, the number of electrostatic interactions increased. This implies that increased electrostatic interactions can lead to strong adhesion of the HAp-bonded coating.

[0123] 2.2. Decomposition into In-Bit

[0125] To determine the effect of the coating on the degradation rate, various parameters such as pH change (Fig. 7B), hydrogen gas generation (Fig. 7C), and degradation rate over time (Fig. 7A) were studied.

[0126] After a 14-day immersion period, the pH of the samples (Mg, MgF2, MgF2 / PLLA-HAp, and Mg-C) was 10.70, 9.50, 8.80, and 7.98, respectively (Fig. 7B). The significant decrease in pH of the uncoated Mg sample compared to the coated sample indicates that the Mg-C layered coating reduced the degradation rate. The internal MgF2 coating facilitates local alkalization, thereby reducing the polymer (PLLA or PCL). After 2 days of incubation, the pH of the samples was found to vary depending on the degradation of Mg. The pH of the MgF2 coating was higher than that of the MgF2 / PLLA-HAp and Mg-C samples, which may be due to the rapid degradation of the MgF2 substrate. The presence of organic coatings beyond MgF2 did not significantly affect the increase in the sample pH. Furthermore, when the PLLA-HAp layered sample was coated with the PCL polymer, the change in pH was significantly reduced. This can be established from the results that the Mg-C stacking sequence helps effectively protect the Mg alloy from corrosion by preventing the penetration of immersion fluids.

[0127] The generation of hydrogen gas is directly related to changes in pH. The results regarding hydrogen gas generation are shown in Fig. 7C. For the uncoated Mg sample, hydrogen gas generation was highest, indicating rapid decomposition of the Mg sample. MgF2 significantly reduces the rate of hydrogen gas generation, which is further reduced by the subsequent PLLA-HAp and PCL layers. These results resonate with the results regarding pH change. The final coating of Mg-C effectively inhibited liquid penetration into the surface and significantly reduced the corrosion rate.

[0128] The white loss of the samples was measured at different time points of 3, 7, 10, and 14 days after immersion in PBS at 37 °C (Fig. 7A). The degradation rate of the exposed Mg alloy samples was higher compared to the coated samples. Among the coated samples, the MgF2 coating showed the lowest protection against degradation. When MgF2 was coated with PLLA-HAp, the degradation rate decreased significantly. The triple-layer coated sample, Mg-C, exhibited the lowest degradation rate. The descending order of degradation rates was Mg > MgF2 > MgF2 / PLLA-HAp > The samples were Mg-C. After immersion in PBS at 37°C for 14 days, the weight loss rates for the samples (Mg, MgF2, MgF2 / PLLA-HAp, and Mg-C) were 28%, 16.9%, 5.4%, and 3.1%, respectively. As the immersion time increased, the degradation rate for all samples also increased. The trend of continuously increasing degradation with increasing degradation time indicates that while the coating layer significantly protects the Mg alloy against corrosion, prolonged exposure over time leads to delamination of the protective layer. The results also indicate that the Mg alloy exhibited an accelerated corrosion rate within 3 days compared to prolonged exposure to PBS. This can be attributed to the fact that a large surface area is initially exposed to PBS, enhancing the corrosion rate due to exothermic reactions. Increased initial exposure to chloride ions and other salts may also lead to increased corrosion. In other cases of the coated samples, the degradation rate increased with increasing PBS exposure time. This can be attributed to the fact that prolonged exposure to PBS may cause the adhesion of the coating layer to detach or weaken. In the degradation study, it was found that the Mg-C coating layer exhibited the least degradation compared to other sample sets. The presence of HAp enables the coating to adhere more strongly to the Mg alloy compared to the sole polymer coating. To evaluate this, the degradation rate of Mg-C was compared with that of the PLLA / PCL coating (Fig. 8). As a result, Mg-C showed a lower degradation rate compared to the polymer coating.

[0129] Figure 7D schematically compares the decomposition of the exposed Mg alloy and Mg-C. In the case of the Mg-C sample immersed in PBS, the penetration of PBS occurred through the pores of the coating layer. As soon as the PBS reached the surface of the Mg alloy, cathodic and anodic reactions began.

[0131] Mg = Mg 2+ +2e - (Bipolar) (1)

[0132] 2H2O+2e -= H2 + 2OH - (Cathode) (2)

[0133] Mg 2+ +2H2O = Mg(OH)2+H2(g) (3)

[0135] The formation of Mg(OH)2 generated H2 gas along with an increase in pH. Since the deposition of Mg(OH)2 occurs at the interface between the Mg alloy and the coating, the H2 gas displaces the coating. This reduces the adhesion between the coating layer and Mg, causing the coating to unravel and crack. Because this phenomenon is accompanied by aggressive salts such as MgCl2, the protective coating is damaged and degrades over time. Figure 9 shows SEM images of the samples after immersion for 14 days. The Mg and MgF2 samples exhibited numerous cracks, whereas the final Mg-C sample showed a single crack.

[0137] 2.3. Drug and Growth Factor Loading

[0139] Figure 10A shows a schematic diagram of Mg-C / PTX-VEGF. Figures 10B–D show SEM images after loading with PTX, VEGF, and PTX / VEGF. The morphology of the coating showed changes after loading the growth factors and drugs. After loading the drugs, small needle-shaped crystals were observed, and there was no significant change in appearance. Since VEGF loading also did not alter the appearance, it can be seen that no obvious changes were observed. The FT-IR spectra of the Mg alloy and the coating are shown in Figure 10E. The FT-IR of the clean Mg alloy does not absorb infrared light, indicating the absence of chemical bonds on the surface. For MgF2, 3000 cm⁻¹ -1 and 1600cm -1 The characteristic peaks in the periphery are attributed to the presence of water. The PLLA-coated sample is approximately 2900 cm⁻¹ -1 to 2800 cm -1 IR peaks were observed, which can be attributed to -CH and -CH2 stretching vibrations. 1700 cm⁻¹ -1The peak in the vicinity is attributed to the stretching vibration of C- - O. The CO stretching vibration is at 1100 cm. -1 It was found in the vicinity. A distinct shift in the peak was observed due to the binding of HAp particles. The characteristic peak of -PO4-2 is at 1090 cm⁻¹. -1 It can be observed at. In the case of PCL-coated samples, the stretching of the carbonyl group is approximately 1650 cm⁻¹ -1 It was observed at -PO4-2. A prominent peak for stretching was at approximately 1090 cm⁻¹. -1 It can be observed at 1708 cm for PTX--O, CC stretching, CH3 bending, CN stretching, and CO stretching, respectively. -1 , 1646cm -1 , 1370cm -1 , 1241cm -1 and 1072cm -1 It showed a peak at 1700 cm⁻¹. In the case of VEGF, 1700 cm⁻¹ -1 The peak at is amide I, 1450 cm⁻¹ -1 Silver amide II, 1300 cm -1This is attributed to amide III. The surface chemical state was measured using XPS analysis (Fig. 10F). XPS analysis of the Mg alloy showed peaks for Mg, Zn, O, and C. The formation of Mg(OH)2 or some carbohydrates due to moisture in the air may be the reason for the appearance of the O peak. C may appear due to environmental pollution. In the MgF2-coated samples, the peak for F can be observed alongside the Mg peak, confirming its formation. The PLLA-HAp-coated samples showed peaks for Ca, P, C, and O. High-resolution peaks for Ca2p and P2p are shown in Figs. 11B-C. The C-Mg and Mg-C / PTX-VEGF samples show peaks for C and O. High-resolution analysis showed an N1s peak for the Mg-C / PTX-VEGF-coated samples (Fig. 11D). This can be aligned with the presence of PTX in these samples.

[0141] 2.4. VEGF and Drug Release

[0143] Figure 12A shows the amount of VEGF loaded onto the Mg-C / PTX-VEGF and Mg-C / VEGF samples. Figures 12C and D show the cumulative VEGF and drug release profiles of Mg-C. The release profiles showed an initial burst release followed by sustained release. The Mg-C / VEGF-PTX coated stents exhibited releases of VEGF and PTX in the ranges of (30%–49%) and (10%–35%) within 1 to 24 days, respectively. The results also indicated that the Mg-C / PTX-VEGF and Mg-C / VEGF samples exhibited similar VEGF release profiles. A comparison of the Mg-C / PTX-VEGF and Mg-C / PTX samples showed that PTX release was not affected by the co-loading of VEGF, and vice versa. Immunohistological staining revealed a uniform distribution of VEGF on the coated metal surface (Figure 12B). The slow degradation process caused by the coating can be attributed to the sustained release of the drug and growth factors.

[0145] 2.5. Platelet Adhesion and Hemolysis

[0147] To evaluate the hemolytic compatibility of biomaterials, investigating platelet activation and aggregation is an essential parameter. The activation phases of platelets generally correspond to platelet shapes: round, dendritic, diffuse dendritic, and fully spread. SEM images showed that platelets were fully spread on the exposed magnesium alloy, whereas the Mg-C, Mg-C / VEGF, Mg-C / PTX, and Mg-C / PTX-VEGF samples exhibited dendritic or round platelet shapes (Fig. 13). This implies that the coating and loading of drugs and growth factors reduced platelet adhesion and aggregation on the metal surface. The hemolysis rate shows that the exposed Mg alloy had a value of 6.5%, which decreased sharply after coating (Fig. 14D).

[0148] The hemolysis rates of Mg-C, Mg-C / VEGF, Mg-C / PTX, and Mg-C / PTX-VEGF ranged from 0.26 to 0.22%. According to ISO protocol (10993-4:2002), a hemolysis rate of less than 5% is considered safe for biomedical use. In the case of the exposed Mg alloys, the rapid release of Mg ions and hydrogen gas generation inhibited cell proliferation and induced hemolysis. As indicated by the degradation rate and pH changes, a decrease in pH was also observed after the degradation rate of the coating decreased.

[0150] 2.6. In-vitro biocompatibility

[0152] To determine the biocompatibility of CPEA, cells were cultured on Mg-C, Mg-C / VEGF, Mg-C / PTX, and Mg-C / PTX-VEGF samples. According to previous studies, cell proliferation on exposed Mg alloys is very low due to hydrogen gas generation and rapid release of Mg ions from uncoated Mg alloy surfaces. An increase in pH adversely affects cell proliferation (Fig. 14C). To measure cell viability, an indirect method was used with various concentrations of cell extracts after 1 day of culture (Fig. 14B). Cell viability decreased as the extract concentration increased from 12.5% ​​to 100%. Compared to the exposed Mg alloy, the coated alloys increased cell viability. The cell proliferative capacity exhibited by the implant is a direct indicator of biocompatibility. Confocal images show enhanced cell proliferation by the coated Mg samples (Fig. 14A). Loading of PTX did not adversely affect the proliferation of CPEA cells. According to previous studies, PTX inhibits the proliferation of smooth muscle cells when coated on cardiovascular stents. With the loading of growth factors on coated Mg samples, cell proliferation increased, indicating that VEGF can promote the proliferation of endothelial cells.

[0154] 2.7. In-vivo Biocompatibility

[0156] Mg-C and Mg-C / PTX-VEGF were implanted into the rat aorta for 1 and 2 months, respectively. To mimic the shape of a stent, a laser-cut Mg mesh was rolled into a stent shape (Fig. 15B, S2). The stent was implanted into the rat aorta (Fig. 15A). Fig. 15D shows an H&E image of a cross-section of the stented artery. To evaluate the effects of the growth factor and the drug, the side opposite to the sutured side was considered as shown in Fig. 15C.

[0157] As a result, one month after transplantation, the neoendothelium formed by the Mg-C stent was intensive and heterogeneous, whereas the Mg-C / PTX-VEGF stent formed thin and homogeneous neoendothelium. Even after two months after transplantation, the C / PTX-VEGF stent still produced thin and homogeneous neoendothelium, whereas the C-Mg stent resulted in the proliferation of heterogeneous neoendothelium.

[0158] Vascular tissues were stained with anti-α-SMA, anti-CD68, anti-ICAM, and anti-F1k1 antibodies. The α-SMA factor represents smooth muscle cells (SMCs). PTX-loaded samples showed a decrease in α-SMA expression compared to non-drug-loaded samples. Controlled remodeling of SMCs is essential for arterial contractility, but uncontrolled proliferation can lead to restenosis. Samples without drug loading exhibited higher levels of uncontrolled SMC proliferation (Fig. 16A, C). CD68 expression was determined to assess the rate of inflammation (Fig. 16B, D). As a result, samples treated with the drug showed lower inflammation compared to samples without the drug. Even two months after transplantation, samples without the drug showed a very high rate of inflammation. In the early stages, inflammation aids in regeneration. However, prolonged inflammation can lead to adverse effects.

[0159] To indicate reendothelialization, two markers, FLKI and ICAM 1, were used (Fig. 17). The results showed that VEGF-loaded samples exhibited higher expression of both markers compared to non-growth factor-loaded samples 2 months after transplantation. In VEGF-loaded samples, the expression of endothelial markers indicated endothelialization along the outer layer. In non-VEGF-loaded samples, the expression of markers was scattered and heterogeneous, indicating the absence of continuous endothelium. Consequently, this demonstrates that dual drug and growth factor loading resulted in the regulation of inflammation, neoendothelial formation, and reendothelialization in Mg-C / PTX-VEGF coated stent samples.

[0160] The dual effect of drugs and growth factors is a promising strategy for regulating neoendothelium and accelerating reendothelialization. Previous studies have shown that using endothelial progenitor cells and the antibody CD34 can accelerate reendothelialization but not reduce neoendothelium. Although there are few reports introducing methods to simultaneously regulate neoendothelium and promote reendothelialization, attempts are being made regarding non-biodegradable stent materials. There are several factors to consider when handling biodegradable materials such as Mg. There are limitations to biocompatible manipulation while addressing the rapid degradation of Mg. Gao Fan et al. created a bioactive layer on an Mg alloy to address the rapid corrosion and reendothelialization of the material. Similarly, Tang Hongyan et al. coated an Mg alloy with poly(1,3-trimethylene carbonate) to control rapid corrosion and deliver drugs. Compared to these existing reports, the present invention addresses both aspects of biocompatibility, namely re-endothelialization and controlled neoendothelialization, but indicates a novel aspect of biodegradable Mg stents by reducing the rapid degradation of the Mg alloy.

[0162] Foregoing, specific parts of the present invention have been described in detail. It will be apparent to those skilled in the art that such specific descriptions are merely preferred embodiments and do not limit the scope of the invention. Accordingly, the actual scope of the invention is defined by the appended claims and their equivalents.

Claims

Claim 1 A vascular stent comprising a magnesium alloy having a triple-layer coating structure in which the MgF2 layer, PLLA-HAp layer, and PCL layer are sequentially laminated on the substrate, comprising: (i) a magnesium fluoride coating layer (MgF2 layer) formed by treating a magnesium alloy ZK-60 substrate with hydrofluoric acid; (ii) a poly(L-lactide)-hydroxyapatite composite coating layer (PLLA-HAp layer) laminated on the MgF2 layer; and (iii) a polycaprolactone coating layer (PCL layer) laminated on the PLLA-HAp layer. Claim 2 A vascular stent according to claim 1, characterized in that the magnesium alloy having the triple-layer coating structure has a reduced decomposition rate compared to an uncoated magnesium alloy of the same composition. Claim 3 A vascular stent according to claim 1, wherein the PCL layer comprises paclitaxel (PTX), vascular endothelial growth factor (VEGF), or a combination thereof. Claim 4 A vascular stent according to claim 1, wherein the magnesium alloy having the triple-layer coating structure is characterized by increased reendothelialization of vascular endothelial cells and decreased neoendothelial formation after vascular grafting, compared to an uncoated magnesium alloy of the same composition or a magnesium alloy of the same composition with only a fluorine coating (MgF2) formed. Claim 5 delete Claim 6 A method for manufacturing a vascular stent comprising a magnesium alloy having a triple-layer coating structure in which the MgF2 layer, the PLLA-HAp layer, and the PCL layer are sequentially laminated on the polished substrate, comprising: a step of polishing a magnesium alloy ZK-60 substrate; a step of treating the polished magnesium alloy with hydrofluoric acid to form a magnesium fluoride coating layer (MgF2 layer); a step of coating the MgF2 layer with a solution mixed with poly(L-lactide) and hydroxyapatite (HAp) to form a PLLA-HAp composite coating layer (PLLA-HAp layer); and a step of coating the PLLA-HAp layer with a polycaprolactone (PCL) solution to form a PCL coating layer (PCL layer).

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