A porous degradable bone substitute scaffold and a method of making the same

Abstract

This invention belongs to the field of bone replacement material technology, specifically relating to a porous biodegradable bone replacement scaffold and its preparation method. The porous biodegradable scaffold of this invention consists of pillars, each pillar having an outer shell and a core. The outer shell comprises ZnO, and the core comprises Zn. The porous biodegradable bone replacement scaffold of this invention has a nano- to millimeter-level hierarchical pore structure, which is more similar to the structure of human bone. Furthermore, its degradation performance, antibacterial properties, and biocompatibility are superior to existing biodegradable metallic bone implant materials (zinc and its alloys), which helps promote osteogenic differentiation.

Description

Technical Field

[0001] This invention belongs to the field of bone replacement material technology, specifically relating to a porous biodegradable bone replacement scaffold and its preparation method. Background Technology

[0002] The main bone substitutes currently available are magnesium and its alloys, iron and its alloys, and zinc and its alloys.

[0003] Magnesium and its alloys have been widely used in the biomedical field due to their complete biodegradability and excellent bioactivity. As an essential element for the human body, its degradation products are not only non-toxic, but magnesium ions can also stimulate sensory nerve endings in the periosteum to release more neurotransmitters, promoting osteogenic differentiation of stem cells in the periosteum and forming new bone in the periosteum. However, magnesium and its alloys also face significant challenges. Magnesium's low electrode potential makes it susceptible to corrosion in body fluids, leading to an excessively rapid degradation rate and a continuous weakening of its mechanical fixation and load-bearing functions. Simultaneously, the degradation process generates a large amount of hydrogen gas, causing a rapid increase in the surrounding pH value, resulting in biotoxicity.

[0004] Compared to magnesium, iron and its alloys possess superior mechanical properties, exhibit a slow degradation rate, and do not produce hydrogen gas through chemical reactions. However, iron's degradation rate is too slow to match bone growth, and its magnetism prevents the use of magnetic resonance imaging (MRI), significantly limiting its practical applications.

[0005] Zinc-based alloys have a standard electrode potential between that of magnesium-based and iron-based metals, exhibit moderate degradation rates, do not produce hydrogen gas, and their degradation products are harmless to the human body. As the second most abundant transition metal in the human body, zinc participates in many biological functions, such as nucleic acid metabolism, signal transduction, and gene expression. However, existing methods for preparing three-dimensional porous zinc and its alloys mostly employ laser powder bed fusion additive manufacturing processes, which cannot achieve nanoscale pores. Furthermore, the degradation rate of the prepared three-dimensional porous scaffolds is uncontrollable, they lack antibacterial properties, and exhibit poor biocompatibility.

[0006] Therefore, there is a need to provide an improved technical solution that addresses the shortcomings of the existing technology. Summary of the Invention

[0007] The purpose of this invention is to provide a porous biodegradable bone replacement scaffold and its preparation method, so as to help solve or improve at least one of the problems of uncontrollable degradation rate, lack of antibacterial properties and poor biocompatibility of zinc and its alloys used as bone replacement materials in the prior art.

[0008] To achieve the above objectives, the present invention provides the following technical solution: a porous biodegradable bone replacement scaffold, the porous replacement scaffold being composed of pillars, having an outer shell and a core, the outer shell being composed of ZnO and the core being composed of Zn.

[0009] The present invention also provides a method for preparing a porous biodegradable bone replacement scaffold, which adopts the following technical solution: A method for preparing a porous biodegradable bone replacement scaffold includes the following steps: (1) mixing zinc powder, hydroxypropyl methylcellulose and water to obtain a metal ink; (2) filling the metal ink into the barrel of an additive manufacturing equipment for printing to obtain a scaffold blank; (3) degreasing and sintering the scaffold blank to obtain the porous biodegradable bone replacement scaffold; in step (3), the degreasing and sintering are carried out under the condition of continuous air supply, the sintering temperature is 450-700℃, and the holding time is 0.5-12h.

[0010] Beneficial effects: The porous biodegradable bone replacement scaffold of the present invention has a nano- to millimeter-level hierarchical pore structure, which is more similar to the structure of human bone. Moreover, its degradation performance, antibacterial performance and biocompatibility are superior to existing biodegradable metal bone implant materials (zinc and its alloys), which helps to promote osteogenic differentiation.

[0011] The porous biodegradable bone replacement scaffold prepared by adding graphene oxide to the raw materials has near-infrared response performance, better antibacterial properties, and good biocompatibility. When the amount of graphene oxide added is appropriate, it helps to better promote osteoblast differentiation. Attached Figure Description

[0012] The accompanying drawings, which form part of this application, are used to provide a further understanding of the invention. The illustrative embodiments of the invention and their descriptions are used to explain the invention and do not constitute an undue limitation of the invention. Wherein: Figure 1 The images show SEM images and EDS analysis of the sintered samples in an argon atmosphere; (a1)-(a3) are samples obtained at a sintering temperature of 450℃, magnified sequentially; (b1)-(b3) are samples obtained at a sintering temperature of 600℃, magnified sequentially.

[0013] Figure 2 The images are SEM images and EDS analysis of the sintered samples in air atmosphere. Among them, (a1)-(a3) are the samples obtained at a sintering temperature of 450℃, magnified sequentially, and (a4) and (a5) are SEM images from different perspectives. (b1)-(b3) are the samples obtained at a sintering temperature of 600℃, magnified sequentially, and (b4) and (b5) are SEM images from different perspectives.

[0014] Figure 3 Characterization of the scaffolds obtained at different sintering temperatures in Example 1; wherein, (a) is a macroscopic image of the degreased and sintered scaffolds, (b) is an XRD pattern of the sample at 420℃ before and after sintering, and (c) is an XRD pattern of the scaffold sintered at 450-700℃.

[0015] Figure 4 The images show SEM and EDS images of the scaffold under different conditions; where (a) is the printed blank, (b) is the degreased sample, and (c) is the sintered sample at 420℃.

[0016] Figure 5 The images are SEM images of the sintered supports at different temperatures in Example 1; where (a1)-(a3) represent samples sintered at 500℃ and (b1)-(b3) represent samples sintered at 550℃.

[0017] Figure 6 The images are SEM images of the sintered supports at different temperatures in Example 1; where (a1)-(a3) are sintered at 650℃ and (b1)-(b3) are sintered samples at 700℃.

[0018] Figure 7 The images shown are cross-sectional SEM and EDS images of the sintered supports at different temperatures in Example 1; where (a) is 450℃, (b) is 500℃, (c) is 550℃, (d) is 600℃, (e) is 650℃, (f) is 700℃, (g) is the elemental composition obtained from the cross-sectional EDS image, and (h) is the oxide layer thickness measured from the cross-sectional SEM image.

[0019] Figure 8 The compression characteristics of the sintered scaffold at different temperatures during in vitro biodegradation in Example 1 are shown in Figure 1. (a) is the stress-strain curve, (b) is the compressive strength, (c) is the yield strength, and (d) is the elastic modulus.

[0020] Figure 9 These are microscopic SEM images of the sintered scaffolds at different temperatures during the in vitro biodegradation process in Example 1; where (a) is T500, (b) is T550, (c) is T600, and (d) is T650.

[0021] Figure 10 The images show the morphology and chemical composition of the sintered scaffolds from Example 1 after 28 days of in vitro degradation at different temperatures. Among them, (a) is a macroscopic image after 28 days of degradation, (b) is an XRD analysis after 28 days of degradation, (c) is the zinc ion concentration in the c-SBF solution during degradation, (d) is the calcium ion concentration in the c-SBF solution during degradation, and (e) is the phosphorus ion concentration in the c-SBF solution during degradation.

[0022] Figure 11 The images show cross-sectional SEM and EDS images of the sintered scaffolds at different temperatures after 28 days of in vitro degradation in Example 1; where (a) represents T500, (b) represents T550, (c) represents T600, (d) represents T650, (e) represents the elemental composition obtained from the cross-sectional EDS image, and (f) represents the weight loss during the biodegradation process.

[0023] Figure 12 The following are the potentiodynamic polarization curves of the sintered support at different temperatures in Example 1.

[0024] Figure 13 Images of different scaffolds for in vitro antibacterial activity; (a) pure zinc, (b) T500, (c) T550, (d) T600, (e) T650.

[0025] Figure 14 The MTT results are shown for different samples after 1 day, 3 days, and 7 days of co-culture with cells.

[0026] Figure 15 The images show the live and dead staining of different samples after 3 and 7 days of culture.

[0027] Figure 16 The ALP activity of different samples after 3 and 7 days of incubation.

[0028] Figure 17 These are stained photographs of sections of the T550 specimen and bone tissue after implantation experiments; (a) is an overall photograph of the wound, and (b)-(d) are magnified photographs of the areas within the red box in (a).

[0029] Figure 18 The images show SEM and EDS scans of the T550 specimen and bone tissue after in vivo implantation experiments.

[0030] Figure 19 H&E stained images of rat viscera sections after in vivo experiments with T550 sample.

[0031] Figure 20 The images shown are SEM and XRD patterns of the scaffold obtained by sintering in Comparative Example 2; where (a)-(c) are SEM images enlarged sequentially, and (d) is an XRD pattern.

[0032] Figure 21 The image shows the EIS test results of the sintered support obtained in Comparative Example 2.

[0033] Figure 22 The photographs of the support for Comparative Example 3 are physical images and SEM images; where (a)-(b) are physical images from different perspectives, and (c) is an SEM image.

[0034] Figure 23 In the figure, (a) is the MTT test result of the stent of Comparative Example 3, (b) is the live / dead staining result after 3 days, and (c) is the live / dead staining result after 7 days.

[0035] Figure 24SEM and EDS analyses of printed preforms with different GO contents in Example 2 are shown; where (a1)-(a2) are 1GO-H2O and (b1)-(b2) are 2GO-H2O.

[0036] Figure 25 SEM and EDS analyses of the scaffolds obtained from 1GO-H2O printed preforms sintered at different temperatures in Example 2 are shown; where (a1)-(a2) represent 450℃, (b1)-(b2) represent 500℃, and (c1)-(c2) represent 550℃.

[0037] Figure 26 The compressive strength of the support obtained by sintering 1GO-H2O printed blanks at different temperatures in Example 2 is shown.

[0038] Figure 27 The image shows a SEM image of the support obtained by sintering a 2GO-H2O printed preform at 550°C in Example 2; where (b) is an enlarged view of (a).

[0039] Figure 28 The XRD patterns are of the 1GO-H2O and 2GO-H2O printed preforms of Example 2 and the scaffolds (Sintered 1GO-H2O, Sintered 2GO-H2O) obtained by sintering at 550°C.

[0040] Figure 29 The surface morphology of Sintered 1GO-H2O and Sintered 2GO-H2O after in vitro degradation for 3 days and 7 days, respectively, in Example 2; where (a) is Sintered 1GO-H2O after 3 days of degradation, (b) is Sintered 2GO-H2O after 3 days of degradation, (c) is Sintered 1GO-H2O after 7 days of degradation, and (d) is Sintered 2GO-H2O after 7 days of degradation.

[0041] Figure 30 The surface morphology of Sintered 1GO-H2O and Sintered 2GO-H2O after in vitro degradation for 14 days and 28 days, respectively, in Example 2; wherein, (a) is Sintered 1GO-H2O after 14 days of degradation, (b) is Sintered 2GO-H2O after 14 days of degradation, (c) is Sintered 1GO-H2O after 28 days of degradation, and (d) is Sintered 2GO-H2O after 28 days of degradation.

[0042] Figure 31 This is the temperature change curve of the stent under near-infrared light.

[0043] Figure 32 This refers to the release of reactive oxygen species from the stent under near-infrared light.

[0044] Figure 33 Nyquist plot of Sintered 1GO-H2O scaffold after 28 days of in vitro degradation; where (a) is under no light and (b) is under near-infrared light.

[0045] Figure 34 Photograph of the in vitro near-infrared antibacterial coating on the stent.

[0046] Figure 35 The images show the live and dead staining of the scaffold and cells after 3 and 7 days of co-culture in the dark; (a1)-(a2) are T500, (b1)-(b2) are Sintered 1GO-H2O scaffolds, and (c1)-(c2) are Sintered 2GO-H2O scaffolds.

[0047] Figure 36 The images show the live and dead staining of the scaffolds after co-culturing with cells under NIR light for 3 and 7 days; where (a1)-(a2) are T500, (b1)-(b2) are Sintered 1GO-H2O scaffolds, and (c1)-(c2) are Sintered 2GO-H2O scaffolds.

[0048] Figure 37 The MTT results are shown after 1, 3, and 7 days of co-culture of scaffold and cells; (a) is under no light and (b) is under near-infrared light.

[0049] Figure 38 The ALP activity was measured after 1, 3, and 7 days of co-culturing the scaffold and cells; (a) was under no light and (b) was under near-infrared light. Detailed Implementation

[0050] This invention addresses at least one of the problems in the prior art where zinc and its alloys used as bone substitutes suffer from uncontrollable degradation rates, lack of antibacterial properties, and poor biocompatibility, by providing a porous biodegradable bone substitute scaffold.

[0051] The porous biodegradable bone replacement scaffold of this invention is composed of pillars, which have a shell and a core in a core-shell structure. The shell is composed of ZnO and the core is composed of Zn.

[0052] In a preferred embodiment of the porous biodegradable bone replacement scaffold of the present invention, the thickness of the outer shell of the strut is 0.1-20 μm; the radial dimension of the strut is 100-800 μm; and the spacing between two parallel and adjacent struts is 100-1000 μm.

[0053] Preferably, the thickness of the outer shell of the support column is 0.38-6.62 μm; the radial dimension of the support column is 587-700 μm; and the spacing between two parallel and adjacent support columns is 559-613 μm.

[0054] In a preferred embodiment of the porous biodegradable bone replacement scaffold of the present invention, the yield strength of the porous replacement scaffold is 1-10 MPa; the elastic modulus of the porous replacement scaffold is 150 MPa-500 MPa; and the compressive strength of the porous replacement scaffold is 1-50 MPa.

[0055] Preferably, the yield strength of the porous alternative stent is 1.03-1.08 MPa; the elastic modulus of the porous alternative stent is 192.49 MPa-240.79 MPa; and the compressive strength of the porous alternative stent is 3.3-13.31 MPa.

[0056] Preferably, the porous alternative scaffold component also includes graphene oxide.

[0057] The present invention also proposes a method for preparing a porous biodegradable bone replacement scaffold. The method for preparing a porous biodegradable bone replacement scaffold in the embodiments of the present invention includes the following steps: (1) mixing zinc powder, hydroxypropyl methylcellulose and water to obtain a metal ink; (2) filling the metal ink into the barrel of an additive manufacturing equipment for printing to obtain a scaffold blank; (3) degreasing and sintering the scaffold blank to obtain a porous biodegradable bone replacement scaffold; in step (3), degreasing and sintering are carried out in an air atmosphere, the sintering temperature is 450-700℃ (e.g., 450℃, 520℃, 620℃ or 700℃), and the holding time is 0.5-12h (e.g., 0.5h, 2h, 3h, 4h, 5h, 6h, 8h, 10h or 12h).

[0058] Preferably, in step (3), the sintering temperature is 450-650°C (e.g., 450°C, 500°C, 550°C, 600°C or 650°C). More preferably, in step (3), the sintering temperature is 500-650°C (e.g., 500°C, 550°C, 600°C or 650°C).

[0059] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, in step (1), the mass ratio of hydroxypropyl methylcellulose to zinc powder is (0.1-2.0):100 (e.g., 0.1:100, 0.5:100, 1:100, 1.5:100 or 2:100); the mass ratio of water to zinc powder is (1-20):100 (e.g., 1:100, 5:100, 10:100, 15:100 or 20:100); and the particle size of the zinc powder is 10-50 μm (e.g., 10 μm, 20 μm, 30 μm, 40 μm or 50 μm). If the proportion of hydroxypropyl methylcellulose is too small, it will result in uneven thickness of the extruded metallic ink lines, with straight lines being thinner and curved lines being thicker, leading to distortion and collapse of the printed support structure. If the proportion of hydroxypropyl methylcellulose is too large, it will result in excessively high viscosity of the metallic ink, hindering the extrusion process and impairing printability. Preferably, the mass of hydroxypropyl methylcellulose is 4.0% of the mass of water (at this ratio, the printed support blank has uniform interlayer arrangement, its size is close to the design value, and the zinc particles are connected by a neck structure mediated by an adhesive).

[0060] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, in step (2), the additive manufacturing equipment has an extrusion device, a barrel, and a printing nozzle; the extrusion device includes a plug and an electric motor, the plug is disposed inside the barrel, the barrel is used to hold metal ink, and the electric motor is used to apply pressure to the plug; the printing nozzle is disposed on the barrel at one end away from the plug; during printing, the diameter of the extrusion needle is 0.2-1 mm (e.g., 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm, or 1 mm), the layer height is set to 0.1-1 mm (e.g., 0.1 mm, 0.2 mm, 0.4 mm, 0.6 mm, 0.8 mm, or 1 mm), the porosity is 40%-100% (e.g., 40%, 50%, 60%, 70%, 80%, 90%, or 100%), and the running speed is 10-50 mm / s (e.g., 10 mm / s, 20 mm / s, 30 mm / s, 40 mm / s, or 50 mm / s).

[0061] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, in step (3), the degreasing temperature is 300-400℃ (e.g., 300℃, 320℃, 340℃, 360℃, 380℃ or 400℃), and the degreasing time is 0.5-4h (e.g., 0.5h, 1h, 1.5h, 2h, 2.5h, 3h, 3.5h or 4h).

[0062] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, in step (2), the additive manufacturing equipment has an extrusion device and a printing nozzle, and the barrel is a replaceable barrel; the printing nozzle is connected to the barrel, and the plug at the top of the barrel is connected to an electric rotary machine, and pressure is applied to the plug through the rotary machine to extrude the metal ink (slurry).

[0063] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, in step (2), the scaffold blank is placed on the printing platform and the temperature of the printing platform is 70-80℃ (e.g., 70℃, 72℃, 74℃, 76℃, 78℃ or 80℃); after the printing of step (2) is completed, the scaffold is left to stand at room temperature for more than 24 hours before step (3) is performed.

[0064] In a preferred embodiment of the method for preparing the porous biodegradable bone replacement scaffold of the present invention, step (1) further includes the addition of graphene oxide; the mass of graphene oxide in the metallic ink is 0.1 wt.%-10.0 wt.% of the mass of water (e.g., 0.1 wt.%, 1 wt.%, 4 wt.%, 6 wt.%, 8 wt.%, or 10 wt.%). Preferably, the mass of graphene oxide in the metallic ink is 1.0 wt.%-2.0 wt.% of the mass of water (e.g., 1.0 wt.%, 1.2 wt.%, 1.4 wt.%, 1.6 wt.%, 1.8 wt.%, or 2.0 wt.%).

[0065] The porous biodegradable bone replacement scaffold of the present invention and its preparation method are described in detail below through specific embodiments. Unless otherwise specified, all raw materials used in the following embodiments are commercially available. The main raw materials used in this experiment are: zinc powder manufactured by Beijing October New Materials Technology Co., Ltd., with a purity of 99.88% and a particle size of 10-50 μm; hydroxypropyl methylcellulose (HPMC) was purchased from Shanghai Maclean Biochemical Technology Co., Ltd.; and graphene oxide was purchased from Jiangsu Pioneer Nanomaterials Technology Co., Ltd., XF021 type graphene nanosheets with a diameter of 5-10 micrometers, a thickness of 3-10 nm, a carbon content of >99.5%, and a specific surface area of ​​31.657 m². 2 / g.

[0066] Example 1 The preparation method of the porous biodegradable bone replacement scaffold in this embodiment includes the following steps: (1) HPMC powder and deionized water are mixed at a mass ratio of 4.0% (the mass of HPMC powder is 4% of the mass of deionized water), and mixed at room temperature with a magnetic stirrer at a speed of 100r / min for 30min. Then, the mixture is allowed to stand for 24h to defoam and ensure that the HPMC powder is fully dissolved to obtain a binder solution. (2) According to the mass ratio of Zn powder to binder solution of 7:1, Zn powder is added to binder solution and mixed evenly to prepare metal ink (metal ink is in slurry state at room temperature). (3) Draw a porous scaffold model (cube or cylinder, porosity 50%; that is, the volume of the pores is 50% of the total volume, calculated value) using Solidworks software, and export the drawn model as STL format; import the STL format model into Cura slicing software, and set the printing parameters through Cura software. The specific additive manufacturing parameters are as follows: extrusion needle diameter is 0.6mm, layer height is 0.4mm, running speed is 60mm / s, and initial layer thickness is 0.4mm; after setting all printing parameters, export it as a Gcode format model; then import it into the ink direct writing additive manufacturing equipment for use on the sample to be prepared. (4) Fill the metal ink into the additive manufacturing equipment barrel, connect the barrel to the printing nozzle, and fill the pressure plug connected to the rotary machine above the barrel. Adjust the printer parameters so that the platform movement speed is 2.8 mm / s and the barrel extrusion rate is 0.036 mL / min. Then adjust the printing platform temperature to between 70℃ and 80℃, and then start printing the product. The freshly printed sample is close to the shape of putty. (5) The printed sample was left to stand at room temperature for more than 24 hours to allow the internal moisture to evaporate and harden completely (the resulting sample is referred to as “Printed sample”); then the sample was degreased and sintered using an OTF-1200X tube furnace. The sample was heated to 350°C in an air atmosphere and held for 2 hours for degreasing (the resulting sample is referred to as “Debinded sample”). Then the temperature was raised to the sintering temperature and held at the sintering temperature for 6 hours (air was continuously introduced at a rate of 50 mL / min during the sintering process). Finally, a porous biodegradable bone replacement scaffold with a Zn@ZnO core-shell structure and regular pore shape was obtained in this embodiment.

[0067] Specifically, the sintering temperatures were set to 420℃, 450℃, 500℃, 550℃, 600℃, 650℃ and 700℃ respectively, and the corresponding samples were recorded as T420, T450, T500, T550, T600, T650 and T700 respectively.

[0068] Example 2 The only difference between this embodiment and Example 1 is that graphene oxide (GO) is introduced into the binder solution to prepare a Zn / GO composite scaffold (the corresponding samples are prepared sequentially at sintering temperatures of 450℃, 500℃, and 550℃). Specifically, step (1) of this embodiment is as follows: GO is first ultrasonically dispersed in water for 10 min, and then HPMC is added and mixed evenly to obtain a pre-made ink; the rest is consistent with Example 1.

[0069] Samples were prepared with GO added at 1.0 wt.% and 2.0 wt.% of the mass of water, respectively. The samples were processed according to steps (2)-(4) in Example 1 and dried at room temperature for 24 h. The resulting printed blanks were denoted as 1GO-H2O and 2GO-H2O.

[0070] After degreasing 1GO-H2O (same as in Example 1), it was sintered at 450℃, 500℃, and 550℃ respectively to obtain the corresponding samples (the sample sintered at 550℃ is denoted as Sintered 1GO-H2O). After degreasing 2GO-H2O (same as in Example 1), the sample sintered at 550℃ is denoted as Sintered 2GO-H2O.

[0071] Comparative Example 1 The only difference between this comparative example and the samples prepared at sintering temperatures of 450°C and 600°C in Example 1 is that degreasing and sintering are carried out under an argon atmosphere; all other aspects are consistent with Example 1.

[0072] Comparative Example 2 The only difference between this comparative example and Example 1 is that the degreasing and sintering processes are carried out in a box-type resistance furnace (the box-type resistance furnace is a sealed environment with only limited air inside, and no air is continuously introduced during the sintering process), and the sintering temperature is 550°C; all other aspects are consistent with Example 1.

[0073] Comparative Example 3 This comparative example uses selective laser melting (SLM) to prepare pure Zn scaffolds. Specifically, the scaffolds are cylindrical in shape with dimensions of 10 mm in height and 3 mm in diameter, based on a diamond unit cell structure with a unit cell size of 0.4 mm and a support thickness of 0.4 mm. The laser power of the laser powder bed melting equipment is 500 W, the powder bed thickness is 60 μm, and the energy density is 39.0 J / mm². 3 After printing, the molded sample is cut from the steel substrate by wire EDM and ultrasonically treated with 96% ethanol for 20 minutes to remove residual powder particles inside the support.

[0074] Experimental Example Characterization and compositional analysis of the scaffold: The porous structure and surface morphology of the scaffold before and after sintering were observed using scanning electron microscopy (SEM), and the filament diameter and pore size in the SEM images of the scaffold were measured using ImageJ. The elemental composition of the scaffold surface or interface was detected using an energy dispersive spectroscopy (EDS) instrument attached to the SEM. Phase analysis of the scaffold was performed using a rotating target X-ray diffractometer (XRD) equipped with a graphite bent crystal monochromator. The porosity of the sintered samples was measured according to Archimedes' principle as described in ASTM standard B963-13.

[0075] In vitro immersion test of the scaffold: The in vitro degradation performance of the scaffold was evaluated by immersing it in conventional simulated body fluid (c-SBF). The immersion conditions were isothermal shaking incubation at 37±1℃, and the ratio of solution volume to sample surface area was 6 mL / cm². 2 The concentrations of Zn, Ca, and P ions in c-SBF after immersion for 1, 3, 7, 14, and 28 days were determined using inductively coupled plasma atomic emission spectrometry (ICP-AES). The mass loss of the stent after immersion for 7, 14, and 28 days was evaluated according to ASTM G1-03. The samples were ultrasonically treated in a 0.1 g / mL NH4Cl solution at 70°C for 5 minutes to remove corrosion products, followed by ultrasonic cleaning in deionized water for 15 minutes. The cleaned and dried samples were weighed.

[0076] Characterization of scaffold degradation products: At predetermined immersion time points, samples were removed from c-SBF and dried for further characterization. SEM and EDS were used to analyze the morphology and chemical composition of surface degradation products, and XRD was used for phase composition analysis. After embedding the sample in epoxy resin and exposing the strut cross-section through grinding and polishing, scanning electron microscopy with backscattered electron imaging and energy dispersive spectroscopy were used to observe the morphology and elemental composition of internal degradation products.

[0077] Mechanical performance testing of the stent: According to ISO 13314:2011 standard, the compressibility of the stent before and after in vitro immersion was tested using a universal testing machine at a crosshead speed of 1 mm / min. Key mechanical performance parameters were determined using stress-strain curves.

[0078] Electrochemical performance testing of the scaffold: The open-circuit potential and electrochemical impedance spectroscopy of the sample during in vitro immersion were monitored using an electrochemical workstation. A three-electrode system was used, with the scaffold as the working electrode, a graphite rod as the counter electrode, and a saturated calomel electrode as the reference electrode. The area of ​​the scaffold exposed to c-SBF was 8.84 cm². 2 The circuit was connected via copper wires. EIS measurements were performed on days 1, 3, 7, 14, 21, and 28 of immersion in c-SBF at 37±1℃, with OCP stabilized for 0.5 hours before each measurement. The frequency scan range was 10. 6 Hz to 10 -2Hz. The same three-electrode system was used for potentiodynamic polarization testing. After stabilization at OCP for 0.5 hours, the voltage was scanned from -0.3V to +0.5V (relative to OCP) at a scan rate of 1mV / s. The corrosion potential and corrosion current density were determined by linear extrapolation over the Tafel region.

[0079] In vitro antibacterial assay: The in vitro antibacterial performance of the scaffold was evaluated using the plate count method. Staphylococcus aureus was used as a model strain. After thawing frozen strains, they were inoculated into liquid culture medium and activated by shaking at 37°C and 250 rpm for 24 h. The activated bacterial solution was then used to adjust the bacterial concentration to an OD value of 0.2 with liquid culture medium. The sterilized straight scaffold was placed in a 24-well plate, and 1 mL of the above bacterial solution was added to each well. Three parallel samples were prepared and incubated at 37°C for 24 h. Subsequently, the bacterial solutions from each group were subjected to 10 μL of liquid culture medium. 5 Dilute 50-60 μL of the diluted bacterial solution evenly onto a solid culture medium plate. Invert the plate and incubate it in a 37°C incubator for 24 hours. Then, remove the plate and count the colonies. The antibacterial effect of the scaffold is evaluated by calculating the colony-forming units.

[0080] In vitro cell compatibility assay: Rat bone marrow mesenchymal stem cells (BMSCs) were used and cultured in an incubator at 37°C, 95% RH, and 5% CO2, using cells from passages 3 to 8. The complete culture medium consisted of 90% DMEM (low glucose) basal medium, 10% fetal bovine serum (FBS), and 1% penicillin / streptomycin (P / S). The medium was changed every two days for a seven-day pre-culture. Biocompatibility was evaluated by direct cell seeding. Samples (2.4 mm high, 8.6 mm in diameter) cleaned with acetone and isoacetone were sterilized in an autoclave (121°C, 20 min) and then incubated in DMEM (low glucose) basal medium for 3 days before direct cell seeding.

[0081] Scaffold cytotoxicity was assessed using the MTT assay, and cell status on the scaffolds was observed using a live-dead staining method: Before cell seeding, sterilized scaffolds were placed in the wells of a 24-well plate. Bone marrow mesenchymal stem cells were seeded at a rate of 3 × 10⁻⁶ cells / well. 4 Cells were seeded onto scaffolds at a concentration of [number] cells / mL, with a control group without scaffolds. After culturing for 1, 3, and 7 days, MTT reagent was added to the wells, and the culture plate was incubated in an incubator for four hours. The culture medium was then aspirated and replaced with dimethyl sulfoxide (DMSO). The culture plate was then placed in an incubator and shaken in the dark for 15 minutes. The absorbance at 490 nm was measured using a microplate reader and converted to relative activity. Live / dead staining experiments were performed on days 3 and 7, using a live / dead staining kit, following the same culture procedure as the MTT assay. The stained scaffold cells were observed under an inverted fluorescence microscope in the dark.

[0082] Osteogenic potential of the scaffold was assessed by alkaline phosphatase (ALP): cells were charged at a rate of 8 × 10⁻⁶. 4 Cells were seeded onto scaffolds at a concentration of [number] cells / mL. After 3 and 7 days of culture, the culture medium was aspirated and the cells were washed with phosphate-buffered saline (PFS). Lysis buffer (containing 1% Triton X-100) was then added, and the cells were lysed at 4°C for 30 min. The supernatant was collected by centrifugation. Following the kit instructions, the supernatant was incubated with p-nitrophenyl phosphate (pNPP) substrate solution at 37°C in the dark for 30 min to detect ALP activity. Immediately after adding stop solution, the absorbance at 405 nm was measured using a microplate reader. Enzyme activity (mM / min) was calculated using a standard curve generated from p-nitrophenol. The total protein content of the supernatant was then quantified using a BCA protein assay kit: absorbance was measured at 562 nm after incubation at 37°C for 30 min, and calculated using a protein standard curve. ALP activity was normalized to total protein content and expressed as U / mg protein.

[0083] In vivo biocompatibility test: SD rats (8 weeks old, 6 rats per group) underwent surgery to expose the femoral condyle. A 3mm wide and 3mm deep hole was drilled using a sterile drill bit, and the holes were rinsed twice with sterile saline. A sterile scaffold was placed into the femoral condyle defect, and the muscles and skin were sutured sequentially. Four weeks post-surgery, three rats were sacrificed, and femoral specimens, along with the heart, liver, spleen, lungs, and kidneys, were removed and fixed in 4% paraformaldehyde solution. For the bone containing the scaffold, hard tissue sections were prepared to preserve the integration of the scaffold with the defect site, and stained with fuchsin and methylene blue. Heart, liver, spleen, lungs, and kidney sections were stained with H&E for pathological observation. The group without the scaffold served as the control group.

[0084] Test results: The effect of combustion atmosphere on support properties: In Comparative Example 1, the support was degreased at 350°C under an Ar atmosphere and then sintered at 450°C or 600°C. Figure 1 SEM images and EDS results of the sintered scaffold at different magnifications show that the scaffold lost its structural integrity and fragmented into powder. Although some deformed particles agglomerated locally, almost no sintering necks were observed. Compared with the relatively smooth surface obtained by sintering at 450℃, the scaffold sintered at 600℃ exhibited a large number of rough, hollow shell structures. Figure 1(a2), (b2)). If the oxide film ruptures during sintering, molten zinc will be drawn into the neck region between adjacent particles by capillary action. Therefore, during sintering at 450℃, due to the high Young's modulus (228 GPa) of ZnO, the molten zinc is completely encapsulated within the unruptured oxide shell. When the temperature rises to 600℃, the internal thermal stress exceeds the fracture strength at the thinnest point of the oxide shell, causing the shell to crack. However, the molten zinc does not flow out but evaporates, leaving a large number of hollow shells, while the zinc vapor condenses to form a rough surface morphology. Figure 1 It can also be observed that most particles maintain an irregular morphology, and the surface is sparsely distributed with flaky features, which, according to EDS analysis, are carbon residues produced by the decomposition of HPMC. Figure 1 (a3), (b3)). Under argon atmosphere sintering, the oxide film on the surface of Zn particles is difficult to remove, and it can be observed that the oxide shell protrusions of the support are intertwined at high temperature (600℃).

[0085] Figure 2 The support structure of Example 1 was sintered in air at 450°C and 600°C. Figure 2 (a1) and (b1) show that the porous structure of the sintered support is consistent with the structure of the printed preform, and the layers are tightly bonded together without collapse. Compared with the support formed at 450℃, the support formed at 600℃ has fewer pores on the pillars, the width of the pillars is increased, and the bonding between Zn particles is tighter. Figure 2 (a2), (b2)), the needle-like substances on the surface are coarser and denser, and the interweaving between them is more complex. Figure 2 (a3), (b3)). This may be because Zn particles are oxidized during sintering, and the oxides grow outwards on the oxide layer. However, there are not enough pores between the contacting particles for the oxides to grow, resulting in a tightly interlocked bond between the oxides. This phenomenon becomes more pronounced as the temperature increases and the morphology of the oxides changes. EDS analysis shows that the Zn content decreases and the O content increases with increasing temperature, indicating that the oxidation of Zn particles increases. This suggests that sintering the support in an air atmosphere not only maintains the structure of the support body but also allows for control of the support performance by adjusting the degree of oxidation through temperature control.

[0086] SEM image of the scaffold obtained by sintering in Comparative Example 2 is shown below. Figure 20 As shown in the figure, some Zn particles expand in volume during the melting process, forming irregular protrusions on the surface of the support (Figures (a)-(b)). Some particles combine with each other after melting (Figure c), and the surface is covered with a layer of granular material and a small amount of needle-like structure.

[0087] The XRD pattern of the scaffold obtained by sintering in Comparative Example 2 is shown below. Figure 20As shown in (d); XRD test results show that ZnO is formed in the support after sintering, the relative intensity of the ZnO peak is low, and the ZnO content is lower than that of the sample sintered at 550℃ in a tube furnace (T550 in Example 1; see details). Figure 3 (c)).

[0088] Comparative Example 3: Physical images ((a)-(b)) and SEM images ((c)) of the pure Zn scaffold. Figure 22 As shown.

[0089] Example 2: SEM and EDS analysis results of printed preforms with different GO contents are as follows: Figure 24 As shown in the SEM images ((a2), (b2)), it can be seen that large flake-like materials are evenly distributed on the support of the green body, and no GO agglomerates are observed, indicating that the added GO can be evenly dispersed on the support. The EDS analysis results ((a1), (b1)) show that the carbon content in the support is high, and the carbon content in the 2GO-H2O sample is even higher, indicating that the flake-like GO was successfully added to the support.

[0090] The SEM and EDS analysis results of the scaffolds obtained by sintering the printed preform 1GO-H2O at different sintering temperatures in Example 2 are as follows: Figure 31 As shown, the number of needle-like ZnO particles gradually increases with increasing sintering temperature. Simultaneously, mechanical testing results also show an increase in compressive strength. Figure 26 ).

[0091] The SEM and EDS analysis results of the scaffold obtained by sintering the printed preform 2GO-H2O at 550℃ in Example 2 are as follows: Figure 27 As shown, the microstructure and elemental composition of the 2GO-H2O scaffold are not significantly different from those of the 1GO-H2O scaffold, but the C content of the 2GO-H2O sample is higher.

[0092] The XRD patterns of the printed preforms 1GO-H2O and 2GO-H2O from Example 2, and the scaffolds (Sintered 1GO-H2O and Sintered 2GO-H2O) obtained by sintering at 550°C are shown below. Figure 28 As shown in the figure, due to the low content of GO, only the diffraction peak of Zn was detected on the printed blank, while the diffraction peak of ZnO appeared on the sintered support.

[0093] The effect of sintering temperature on the microstructure of the support: Figure 3The macroscopic morphology and phase composition of the molded scaffold of Example 1 within the selected temperature range are analyzed. It can be seen that the color of the scaffold gradually lightens with increasing sintering temperature ((a); in the figure, "Debinded sample" refers to the sample in Example 1 that was degreased but not sintered). The scaffold sintered at 700℃ exhibits significant dimensional shrinkage, while the scaffold sintered at lower temperatures basically maintains its original geometry. XRD analysis shows that due to the low HPMC content and limited zinc oxidation during degreasing and sintering, only Zn diffraction peaks exist in the printed and degreased scaffolds ((b)), while the sintered scaffold shows corresponding ZnO diffraction peaks ((b), (c)). With increasing temperature, the ZnO diffraction peaks gradually strengthen, while the Zn diffraction peaks weaken. At 700℃, the zinc diffraction peaks almost disappear.

[0094] In the scaffolds formed in Example 1, the degreased scaffolds and the scaffolds sintered at 420°C are brittle and easily break into powder. Figure 4 Compared to the printed sample, the debinded sample and the 420℃ sintered support have fewer particle bonding sites, the C content on the particle surface gradually decreases, and the O content gradually increases due to oxidation, resulting in mass losses of 0.76±0.13% and 0.56±0.11%, respectively.

[0095] The scaffold sintered at 450-650℃ possesses sufficient mechanical strength to maintain structural integrity. As the sintering temperature increases, the width of the scaffold supports gradually increases from 601.72±14.71μm to 689.06±10.33μm, while the pore spacing decreases from 602.53±11.28μm to 568.75±9.23μm. Figure 5 , 6 and Figure 2 The (a4), (a5), (b4), and (b5) values ​​in the model lead to a decrease in the overall porosity of the support structure, while the weight change of the support structure gradually increases from 0.49±0.16% to 14.17±0.30%. Regarding the microstructure evolution, the morphology of the zinc particles undergoes significant changes: in the 450-600℃ temperature range, they exhibit an irregular particle morphology modified with needle-like zinc oxide, evolving into a plate-like zinc oxide structure at 650℃. Notably, sintering at 700℃ induces significant size shrinkage, with the support width and pore spacing decreasing sharply to 573.44±6.04μm and 565.63±16.05μm, respectively. Figure 6 (b1)-(b3)), the support structure consists of a fractured hollow shell, accompanied by a significant mass loss of 73.13±4.29%.

[0096] The cross-sectional microstructure and corresponding energy dispersive spectroscopy analysis of the scaffold obtained by sintering at 450-650℃ in Example 1 showed that the irregularly shaped zinc particles were encapsulated by an oxide layer, forming a distinct Zn@ZnO core-shell structure. Figure 7 The interparticle bonding is mainly achieved through oxides formed during sintering. At the lower sintering temperatures of 450℃ and 500℃, the oxide layer is relatively thin (thicknesses of 0.52±0.14μm and 0.94±0.12μm, respectively), with low oxygen content, and effective interparticle bonding only occurs locally (as shown by the red arrows). As the sintering temperature increases to 550℃, 600℃, and 650℃, the oxide layer thickness increases significantly, reaching 2.49±0.40μm, 3.49±0.59μm, and 5.14±1.48μm (h), respectively. At a sintering temperature of 700℃, the support is mainly composed of a hollow zinc oxide shell, with almost no residual metallic zinc cores, and energy dispersive spectroscopy analysis detected a sharp decrease in zinc content (g).

[0097] Mechanical performance testing of the support: Figure 8 The mechanical properties of the sintered scaffold of Example 1 before and after in vitro biodegradation are shown in the figure. As can be seen from the figure, the compression curves of the scaffolds obtained by sintering at 500℃, 550℃, 600℃, and 650℃ exhibit yield characteristics ((a)). The yield strength and elastic modulus of each sample are comparable, ranging from 1.03 MPa to 1.08 MPa and from 192.49 MPa to 240.79 MPa, respectively. The compressive strength significantly increases from 3.86 ± 0.56 MPa at 500℃ to 5.52 ± 0.86 MPa at 550℃, 9.75 ± 0.81 MPa at 600℃, and 11.56 ± 1.75 MPa at 650℃. The stress-strain curves of the scaffolds sintered at 400℃, 450℃, and 700℃ exhibited brittle fracture characteristics ((a)), with low compressive strengths of 0.35±0.09MPa, 1.95±0.32MPa, and 1.18±0.25MPa, respectively. The mechanical properties of the scaffolds sintered at 500℃, 550℃, 600℃, and 650℃ (named T500, T550, T600, and T650, respectively) underwent a 4-week in vitro degradation process. The compressive strength and yield strength of the scaffolds did not show a significant decrease during the 28-day immersion period. The elastic modulus of the T500 sample decreased from 192.49±8.01MPa to 115.26±6.08MPa ((d)), still within the applicable range for bone scaffold mechanical properties. The elastic modulus of the remaining groups decreased slightly in the early stage of immersion, with T550 sample decreasing by 9.04%, T600 sample decreasing by 1.46%, and T650 sample decreasing by 5.42%. Subsequently, they all increased by about 2%, and did not show a continuous decline in mechanical properties throughout the entire in vitro immersion period.

[0098] The test results of the compressive strength of the scaffold obtained by sintering the printed preform 1GO-H2O at different sintering temperatures in Example 2 are as follows: Figure 26 As shown.

[0099] Stent degradation performance test: Figure 9 The images show SEM images and energy dispersive spectroscopy (EDS) spectra of the sintered scaffold from Example 1 during its in vitro biodegradation. It can be seen that corrosion products preferentially nucleate and grow at the substrate of needle-like and plate-like oxides, gradually filling the pores within the oxide layer. Simultaneously, fine particles are gradually deposited on top of the oxides (as indicated by the arrows), with the deposition rate of sample T500 significantly higher than other groups. In addition to Zn, O, and C, EDS analysis also detected the presence of Ca, P, and Cl. The higher Cl content in samples T500 and T550 indicates a faster degradation rate of the scaffold.

[0100] After a 4-week in vitro immersion experiment, white corrosion products were observed to form on the surface and inside the pores of the scaffold in Example 1. Figure 10 (a) XRD analysis of the immersed scaffold confirmed the formation of various degradation products, including Zn5(CO3)2(OH)6 and Ca. 10 (PO4)6(OH)2 and Zn5(OH)6Cl6 Figure 10 (b)). Simultaneously, monitoring the release of zinc ions from c-SBF directly confirmed the scaffold degradation behavior. Figure 10 (c)). Zn in the T500 bracket 2+ The release rate was consistently significantly higher than other samples, with the highest cumulative ion concentration reaching 26.32±1.02 mg / L; followed by T600 (13.37±0.24 mg / L), T550 (10.97±0.16 mg / L), and T650 (7.01±0.08 mg / L). The calcium and phosphorus ion concentrations in c-SBF of all samples showed a gradually decreasing trend, indicating the gradual deposition of the calcium-phosphorus phase. Figure 10 (d)-(e)).

[0101] Figure 11The images show cross-sectional SEM and EDS images of the scaffold from Example 1 after 28 days of immersion, as well as the weight loss during degradation. The cross-sectional SEM images of the scaffold show significant degradation of the oxide shell, with localized porosity, loose structure, and discontinuities (circled in (a)-(d)). In the T500, T550, and T600 scaffolds, a clear partial separation between the residual zinc core and the surface oxide layer was observed (as indicated by arrows). Corrosion products (mainly composed of Zn, O, and Cl) were distributed both inside and on the outer surface of the oxide shell. In the T500 sample, a large amount of corrosion products filled the interparticle pores, leading to increased O and Cl content ((e)). In contrast, the T550 and T600 scaffolds contained relatively few corrosion products, with Cl content below 1%. In the T650 sample, the zinc core remained tightly bonded to the oxide layer, with only a small amount of Cl-containing corrosion products forming at the interface. The T500 stent had the highest total mass loss after 28 days, reaching 7.41±0.82%; while the T550, T600 and T650 stents had significantly lower mass losses, at 3.93±0.69%, 4.95±0.51% and 2.82±0.37%, respectively.

[0102] Figure 12 The table shows the potentiodynamic polarization curves of the stent in Example 1. The corrosion current densities of samples T500, T550, T600, and T650 are 69.7±5.29, 47.6±3.09, 74.8±6.34, and 66.8±7.06 μA / cm², respectively. 2 The calculated biodegradation rates were 1.04±0.07, 0.71±0.04, 1.11±0.09 and 0.99±0.08 mm / year, respectively.

[0103] The EIS results of the stent prepared in Comparative Example 2 are as follows: Figure 21 As shown; the test results indicate that the impedance modulus of the sample remained at 8 × 10⁻⁶ throughout the immersion process. 4 Ω·cm 2 The value around 100° reflects that the total polarization resistance of the system remains essentially unchanged. The semicircle radius of the Nyquist plot remains essentially unchanged with immersion time, indicating that the sample does not undergo significant degradation within 28 days, failing to meet the requirements for degradable materials.

[0104] The SEM images and EDS results of the scaffolds (Sintered 1GO-H2O, Sintered 2GO-H2O) obtained by sintering the printed preforms 1GO-H2O and 2GO-H2O at 550°C after 3 and 7 days of in vitro degradation are shown below. Figure 29As shown, in addition to the basic elements Zn, C, and O, Ca and P were detected on the scaffolds, confirming the deposition of calcium-phosphorus compounds on the scaffold surface and indicating that the material has good ability to induce apatite deposition. The presence of Cl was also detected, indicating simultaneous degradation of the scaffold. Compared with the immersion process of the T550 scaffold without GO, Sintered 1GO-H2O and Sintered 2GO-H2O deposited large amounts of Ca, P, and Cl in the early stages of immersion, indicating that the degradation rate of the GO-added scaffold was higher than that of the unadded scaffold in the early stages of immersion. Furthermore, with increasing GO content, the Ca, P, and Cl contents of the Sintered 2GO-H2O scaffold were higher than those of the Sintered 1GO-H2O scaffold, suggesting that the addition of GO promoted the degradation rate and the ability to induce apatite deposition.

[0105] As the soaking time increases, the morphology of the support surface becomes more complex. Figure 30 Compared to Sintered 1GO-H2O, the Sintered 2GO-H2O scaffold had a lower Zn content and a higher Cl content. After 28 days of immersion, the porous structure between Zn particles was almost invisible on the Sintered 2GO-H2O scaffold, and agglomeration of spherical particles could be observed on the surface of the deposit covering the scaffold. In contrast, the deposit on the surface of the Sintered 1GO-H2O scaffold did not completely obscure the original morphology of the scaffold. At this point, the Cl content in the Sintered 2GO-H2O scaffold was nearly twice that of the Sintered 1GO-H2O scaffold, indicating that the addition of GO accelerated the degradation of the scaffold.

[0106] In vitro antibacterial properties of Zn-based scaffolds: Figure 13 Images of in vitro bacterial coatings on different scaffolds are shown. (a)-(e) represent pure Zn (Comparative Example 3), T500, T550, T600, and T650 samples, respectively. As can be seen from the figures, the colony counts of the T500, T550, T600, and T650 samples are all lower than those of the pure Zn sample, indicating that the Zn@ZnO core-shell structure has stronger antibacterial properties. This antibacterial property may originate from the larger specific surface area resulting from the nanoscale porous structure of the samples and the disruption of bacterial cell membranes caused by the unique needle-like / plate-like structure of ZnO.

[0107] MTT test: Figure 14The MTT results are shown for cells cultured directly on the scaffold of Example 1 for 7 days. As can be seen from the figure, the cell viability of the T600 scaffold in Example 1 was significantly lower than that of the T550 and T650 samples on day 1. However, with the extension of culture time, the metabolic activity of cells on the T600 scaffold increased from 70.14±1.64% to 96.40±7.40%, and was significantly higher than other samples after day 3, indicating better cell compatibility. Furthermore, the cell viability of all scaffolds exceeded 75% after 7 days of culture, indicating that all scaffolds had excellent cell compatibility. The MTT test results of the pure Zn scaffold in Comparative Example 3 are shown below. Figure 23 As shown, the cell survival rate of pure Zn scaffolds after 7 days of co-culture with cells was only 35%, indicating poor cell compatibility.

[0108] Live / dead staining test of direct culture: The live / dead staining results showed that BMSCs directly seeded on the scaffold surface of Example 1 were evenly distributed in all samples. Figure 15 Cells on the T500 sample were relatively sparse after 3 days of culture, with some isolated, round cell clusters observed ((a1)). Abundant green fluorescence was detected in the strut pores of other scaffolds ((b1)-(d1)), indicating enhanced cell proliferation. From day 3 to day 7, the density of viable cells increased on the T500, T550, and T600 samples, while cell death was minimal, confirming the high cell compatibility of these scaffolds ((a2)-(c2)). However, the number of dead cells increased on day 7 in the T650 sample ((d2)).

[0109] The results of live / dead staining of the pure Zn scaffold in Comparative Example 3 are as follows: Figure 23 As shown in (b), only a few isolated green fluorescences were observed on the scaffold after 3 days of culture, and even fewer fluorescences were observed after 7 days, indicating that the cell survival rate on the pure Zn sample surface was low and the cell compatibility of the scaffold was poor.

[0110] Directly cultured cell differentiation assay: Figure 16 The ALP activity of the scaffold and cells in Example 1 after 3 and 7 days of co-culture is shown. ALP activity increased significantly on all scaffolds from day 3 to day 7, demonstrating that cells can undergo early osteogenic differentiation on the scaffold surface. The ALP activity of the T600 scaffold was significantly higher than that of other scaffolds, indicating its strong ability to promote osteogenic differentiation.

[0111] In vivo biocompatibility of Zn-based scaffolds: Figure 17The images show stained sections (magenta and methylene blue) of the T550 specimen and bone tissue 4 weeks after implantation into the femoral condyle of rats. In (a), the yellow circle represents the 3mm diameter notch removed from the rat femoral condyle before the implantation experiment, and the 3mm diameter specimen was completely implanted into this area. The black area represents the remaining T550 scaffold, showing significant degradation of the specimen 4 weeks after implantation. Neoplastic bone-like tissue can be seen around the remaining T550 scaffold in the green area. (b), (c), and (d) are magnified views of the area within the red box in image a. (b) Obvious neovascularization can be observed (within the blue circle), providing conditions for the transport of nutrients; (c) Granulation tissue formation plays a role in bone fixation, and the purple area in the upper right corner is relatively mature bone tissue, indicating that it has entered the callus calcification stage and the bone tissue is healing well; (d) Osteoblasts are more obvious between newly formed osteoid and newly formed bone trabeculae (blue area indicated by the blue arrow); Figures (c)-(d) show a large number of spindle cells, which are slightly mature fibroblasts that will gradually form fibrous callus, paving the way for the subsequent formation of bony callus and promoting the repair and regeneration of damaged bone tissue. In (a), the formation of bone trabeculae is clearly observed in the area between the yellow and green circles (pink area), and the purple area is the medullary cavity, indicating that the sample degradation rate is relatively fast. As the sample degrades, the bone tissue healing is good, and the sample has excellent in vivo biocompatibility. Figure 18 The images show cross-sectional SEM and local EDS scans of the T550 scaffold and bone tissue after in vivo implantation experiments (red boxes in the SEM images). The Zn element indicates that the lighter-colored area in the center represents the remaining T550 scaffold. Besides Zn, O, and C, the scaffold also contains Cl, Ca, and P, indicating that its in vivo degradation products are similar to those in the in vitro immersion experiment. The small pieces of scaffold remaining in regions 1 and 2 indicate that the scaffold in the surrounding areas has been degraded, and the scaffold volume has significantly decreased. The T550 scaffold is surrounded by substances containing Ca and P, which should be newly formed bone trabeculae, and the pores in the trabeculae are the medullary cavities.

[0112] Figure 19 The images show H&E-stained sections of rat internal organs after in vivo experiments. Compared with the control group, no obvious organic damage was observed in the T550 group, indicating that the sample has no obvious biotoxicity to rats and has good biosafety.

[0113] Photothermal performance characterization: A near-infrared laser (980 nm, 0.64 W / cm²) was used. 2The sample was irradiated. During irradiation, the temperature of the sample was recorded in real time using a thermal imager. The distance between the irradiating light source and the sample coating was kept constant throughout the photothermal experiment. The laser aperture was adjusted to 9 mm and kept constant, while the distance between the thermal imager and the sample remained constant. The temperature changes of the sample were recorded during the three cycles of heating for 5 minutes, cooling for 5 minutes, and the photothermal temperature-time curve was plotted.

[0114] Test results are as follows Figure 31 As shown, the near-infrared photothermal response of pure Zn in SLM (Comparative Example 3) is relatively poor. During irradiation for 5 minutes, its temperature rise rate and amplitude (only 10±1℃) are significantly lower than other samples. The temperature rise amplitude of the T550 sample (Example 1) is 26±2℃, indicating that it has significant photothermal performance. After adding GO (Sintered 1GO-H2O and Sintered 2GO-H2O in Example 2; represented by "1GO" and "2GO" in the figure, respectively), the temperature rise amplitude of the sample is significantly improved, and the temperature rise amplitude increases accordingly with the increase of GO content, indicating that GO can enhance the near-infrared photothermal performance of the scaffold.

[0115] Near-infrared photodynamic properties: The reactive oxygen species (ROS) generation capacity of the samples was detected using a 2,7-dichlorofluorescein diacetate (DCFH-DA) reactive oxygen species (ROS) kit. Samples were placed in 24-well plates, washed twice with PBS, and 10 μM DCFH-DA working solution (diluted from 10 mM DMSO stock solution with serum-free culture medium) was added. The sample was analyzed using an optical density of 0.64 W / cm². 2 Irradiate the sample with a 980 nm laser for 20 min. After irradiation, remove the working solution, wash with PBS 2-3 times, and measure the fluorescence intensity using a fluorescence microplate reader at an excitation wavelength of 485 nm and an emission wavelength of 525 nm. A sample placed in the dark for 20 min serves as a control group. The level of reactive oxygen species (ROS) generation is expressed as relative fluorescence intensity.

[0116] Figure 32The figure shows the reactive oxygen species (ROS) generation of the samples after 20 min of near-infrared light irradiation. It can be seen that without irradiation, Zn (Comparative Example 3), the T550 scaffold (Example 1), and the composite scaffolds (Sintered 1GO-H2O and Sintered 2GO-H2O in Example 2; represented by "1GO" and "2GO" respectively) all produced fluorescence at 525 nm with the same intensity. After 20 min of NIR light irradiation, the fluorescence intensity of the pure Zn scaffold remained essentially unchanged, indicating that the Zn scaffold lacks near-infrared photodynamic activity. The T550 scaffold showed a significant increase in ROS generation under near-infrared light irradiation, indicating that the scaffold possesses near-infrared photodynamic activity. The addition of GO significantly increased ROS generation, indicating that GO can improve the photodynamic performance of the scaffold, but the GO content had no significant effect on the fluorescence intensity.

[0117] Near-infrared degradation behavior test: Irradiate the sample with a near-infrared light source for 30 minutes every day; Figure 33 The Nyquist plots show the degradation of 1GO (Sintered 1GO-H2O) samples under no light (a) and NIR light (b) conditions for 28 days in vitro. Both samples exhibit a single capacitive arc characteristic, indicating that the corrosion process is primarily controlled by interfacial charge transfer. In Figure (a), the capacitive arc radius gradually increases with immersion time, reaching its maximum at 28 days, indicating that the corrosion product layer on the sample surface continuously thickens and densifies, the charge transfer resistance continuously increases, and the corrosion resistance steadily improves over time. In Figure (b), the capacitive arc radius first increases and then decreases, reaching its peak at 21 days, and then slightly declines, indicating that near-infrared light irradiation causes a dynamic evolution of the sample's corrosion resistance along with the corrosion product layer. The capacitive arc radius of the NIR-illuminated sample is significantly lower than that of the unilluminated sample, indicating that its degradation rate is faster.

[0118] In vitro near-infrared antibacterial experiment: Figure 34 The images show the in vitro antibacterial activity of the GO-added scaffold, with the T550 scaffold serving as the control group. As shown, in the absence of light, the colony density of the T550 sample was significantly higher than that of the 1GO (Sintered 1GO-H2O) and 2GO (Sintered 2GO-H2O) samples. The 2GO sample had the fewest colonies and the sparsest distribution, indicating that the introduction of GO significantly enhances the antibacterial ability of the composite scaffold, and the antibacterial effect increases with increasing GO content. (Using an optical density of 0.64 W / cm²) 2 Irradiation with a 980 nm laser for 20 minutes significantly reduced the colony density of the samples. Almost no colonies were observed in the 2GO sample, indicating that the samples all have near-infrared light-responsive antibacterial properties, and GO further enhanced the photo-antibacterial activity of the samples.

[0119] Effects of GO on the in vitro cell compatibility of scaffolds: Cell proliferation activity: The effect of scaffold extract on cell proliferation activity was assessed using a live-dead staining method. Figure 35 Images show the live and dead cell staining of T550 ((a1)-(a2)), 1GO (Sintered 1GO-H2O) ((b1)-(b2)), and 2GO ((Sintered 2GO-H2O) (c1)-(c2)) scaffolds under no light. After 3 and 7 days of culture, all groups showed predominantly green fluorescence with only a very small number of red dead cells, indicating good cell compatibility in all samples. With increasing time, the number of live cells in each group significantly increased from 3 to 7 days, with cells exhibiting a relaxed morphology and tight arrangement. There was no significant difference in live cell density and morphology among the three groups, and the number of dead cells was extremely low, indicating that GO did not cause cytotoxicity within the 0-2.0 wt% addition range, and the composite scaffolds formed by Zn and GO exhibited good cell compatibility.

[0120] The cell viability and mortality staining images after co-culturing the composite scaffold extract with cells under NIR light are shown below. Figure 36 As shown. Similarly, after 3 and 7 days of culture, all groups showed predominantly green live cells with only a small number of red dead cells. The number of live cells increased significantly with prolonged culture time, and the cells were expanded and tightly packed. Compared to the live / dead staining images without light exposure (…),… Figure 35 At 3 days of culture, cells under NIR light were smaller and had sparser intercellular connections, and this phenomenon became more pronounced with increasing GO addition. At 7 days of culture, there was no significant difference in cell proliferation. This indicates that NIR light may affect cell spreading in the short term, but the effect of NIR light weakens as cells continue to grow and proliferate. In other words, neither NIR light nor GO doping had a significant impact on cell growth and proliferation.

[0121] Cytotoxicity evaluation: Quantitative MTT assay of the composite scaffold was performed to more directly evaluate the impact of GO addition on scaffold cell compatibility. For example... Figure 37 As shown, the cell viability of each group increased with the extension of culture time and remained above 70%, indicating that the composite scaffold with added GO in Example 2 has good cell compatibility.

[0122] Cell differentiation activity: The effect of composite scaffolds with different GO doping levels on osteogenic differentiation was evaluated by ALP activity assay. Results showed ( Figure 38 The Zn-based composite scaffold with an appropriate amount of GO has a good ability to promote osteocyte differentiation, and this ability remains good even under the inhibitory effect of NIR light.

Claims

1. A porous biodegradable bone replacement scaffold, characterized in that, The porous alternative support consists of a support column having a shell and a core, the shell being composed of ZnO and the core being composed of Zn.

2. The porous biodegradable bone replacement scaffold as described in claim 1, characterized in that, The radial dimension of the support column is 100-800μm, and the distance between two parallel and adjacent supports is 100-1000μm. The thickness of the outer shell of the support column is 0.1-20 μm.

3. The porous biodegradable bone replacement scaffold as described in claim 1, characterized in that, The yield strength of the porous alternative scaffold is 1-10 MPa; The elastic modulus of the porous alternative scaffold is 150MPa-500MPa; The compressive strength of the porous alternative stent is 1-50 MPa.

4. The porous biodegradable bone replacement scaffold according to any one of claims 1-3, characterized in that, The porous alternative scaffold also includes graphene oxide.

5. The method for preparing a porous biodegradable bone replacement scaffold as described in claims 1-4, characterized in that, Includes the following steps: (1) Zinc powder, hydroxypropyl methylcellulose and water are mixed to obtain metallic ink; (2) The metal ink is filled into the barrel of the additive manufacturing equipment for printing to obtain a bracket blank; (3) The scaffold blank is degreased and sintered to obtain the porous biodegradable bone replacement scaffold; In step (3), the degreasing and sintering are carried out in an air atmosphere, the sintering temperature is 450-700℃, and the holding time is 0.5-12h.

6. The method for preparing a porous biodegradable bone replacement scaffold as described in claim 5, characterized in that, In step (1), the mass ratio of the hydroxypropyl methylcellulose to the zinc powder is (0.1-2.0):100; The mass ratio of water to zinc powder is (1-20):100; The zinc powder has a particle size of 10-50 μm.

7. The method for preparing a porous biodegradable bone replacement scaffold as described in claim 5, characterized in that, In step (2), the additive manufacturing equipment includes an extrusion device, a barrel, and a printing nozzle; The extrusion device includes a stopper and an electric motor. The stopper is disposed inside the barrel, which is used to hold metallic ink. The electric motor is used to apply pressure to the stopper. The printing nozzle is disposed on the barrel at one end away from the stopper. During printing, the diameter of the extrusion needle is 0.2-1mm, the layer height is set to 0.1-1mm, the porosity is 40%-100%, and the running speed is 10-50mm / s.

8. The method for preparing a porous biodegradable bone replacement scaffold as described in claim 5, characterized in that, In step (3), the degreasing temperature is 300-400℃ and the degreasing time is 0.5-4h.

9. The method for preparing a porous biodegradable bone replacement scaffold as described in claim 5, characterized in that, In step (2), the support blank is placed on the printing platform, and the temperature of the printing platform is 70-80℃; After printing in step (2), let the paper stand at room temperature for more than 24 hours before proceeding to step (3).

10. The method for preparing a porous biodegradable bone replacement scaffold as described in claim 5, characterized in that, Step (1) also includes the step of adding graphene oxide; The mass of graphene oxide in the metallic ink is 0.1 wt.% to 10.0 wt.% of the mass of water.

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