A piezoelectric driving layered bionic scaffold for repairing diabetic bone defects and a preparation method and application thereof

Abstract

The application relates to the technical field of orthopedic clinical repair, and particularly discloses a piezoelectric driving layered bionic support for repairing diabetic bone defects as well as a preparation method and application thereof. The support is a double-layer bionic structure, and the support comprises an upper-layer piezoelectric bionic periosteum simulating a periosteum and a lower-layer conductive filling hydrogel simulating a defect filling matrix. The upper-layer piezoelectric bionic periosteum is a nanofiber membrane, and the lower-layer conductive filling hydrogel is a composite three-dimensional porous hydrogel. The scheme designs a kind of physical hierarchical structure capable of simulating bone tissue, and the composite bionic support can conduct surface electrical signals to the deep part of the defect, and can synergistically regulate local immune metabolism and angiogenesis. The physical hierarchical structure material of the support exhibits excellent mechanical compressive strength, low impedance conductivity, and ultrasound-responsive on-demand degradation and drug release behavior adapted to the bone tissue healing period, and has extremely high clinical conversion potential.

Description

Technical Field

[0001] This invention relates to the field of orthopedic clinical repair technology, and in particular to a piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects, its preparation method, and its application. Background Technology

[0002] With the rising global prevalence of diabetes, diabetic bone defects (DBD) have become a major challenge in orthopedic clinics. Diabetic patients are chronically hyperglycemic, leading to imbalances in local immune metabolism and the vascular microenvironment, triggering persistent oxidative stress and chronic inflammation. This disrupts the coupling process between angiogenesis and bone formation, ultimately resulting in delayed or nonunion of bone.

[0003] Currently, most biomaterials used for DBD repair focus on enhancing a single function (such as simply loading osteogenic factors, promoting angiogenesis, or reducing inflammation), which is insufficient to meet the spatial coordination and multi-level functional requirements of bone regeneration under pathological conditions. Natural bone repair is a systematic process involving clear spatial division: the outer periosteum provides barrier protection and vascular support, while the internal bone defect area requires matrix support for cell infiltration and signal transduction.

[0004] Specifically, the existing technology mainly faces the following technical challenges that urgently need to be addressed: (1) Existing homogeneous materials are insufficient to meet the spatial hierarchical and multi-functional coordination requirements of natural bone repair. The repair of natural bone tissue is a complex systemic process involving a clearly defined spatial division: the outer periosteum, as a highly vascularized active interface, needs to provide physical barriers, cell recruitment, and support for vascular network generation; while the core area of ​​the bone defect requires a porous matrix to support cell infiltration, tissue reconstruction, and signal transduction. Currently, most biomaterials used for the repair of diabetic bone defects adopt a homogeneous structure design, which is often limited to enhancing a single function (such as simple anti-inflammation, simple osteoproliferation, or angiogenesis). They cannot achieve spatial decoupling and synergy between surface vascularization support and deep tissue repair within the same material platform, making it difficult to meet the complex three-dimensional tissue regeneration requirements under pathological conditions.

[0005] (2) The electrical signals of traditional piezoelectric repair materials have "surface limitations," making it difficult to penetrate deep into the core area of ​​bone defects. Electrical stimulation has been proven to be an effective physical means of promoting bone regeneration. In recent years, piezoelectric biomaterials have attracted much attention because they can convert mechanical stimulation such as ultrasound into local electrical signals in situ. However, existing piezoelectric bone repair platforms face a fatal material bottleneck: the piezoelectric microcurrents they generate are often completely confined to the surface layer of the material. Due to the lack of continuous electron / ion conduction pathways inside the scaffold, the electrical signals generated on the surface cannot be effectively transmitted to the blood-deficient core area of ​​the defect, resulting in deep cells being unable to sense electrophysiological stimulation, which severely limits the overall repair efficacy of the material.

[0006] (3) Conventional materials are insufficient to fundamentally reverse the malignant pathological microenvironment of diabetes. The failure to repair bone defects in diabetes is not only a destruction of bone structure, but also a continuous collapse of the local immune, metabolic, and vascular microenvironment. Local hyperoxidative stress (ROS overload) and chronic inflammation caused by hyperglycemia can lead to a long-term pro-inflammatory metabolic state in immune cells such as macrophages, thereby hindering angiogenesis and bone matrix mineralization. Existing conventional biomaterials mostly adopt a strategy of passively releasing traditional anti-inflammatory drugs, which is not only prone to explosive release, but also cannot fundamentally correct this abnormal cellular metabolic pattern. Current technology lacks an intelligent material system that can actively respond to physical stimuli, target and clear deep ROS, and simultaneously release metabolic reprogramming factors, making it difficult to restore the immune-metabolic microenvironment homeostasis of the pathological defect area from the source.

[0007] In summary, while existing piezoelectric materials can promote bone repair through non-invasive physical stimulation, they generally suffer from the problem that the electrical signals are limited to the material surface and cannot penetrate deep into the core of the defect. Furthermore, simple electrical stimulation or anti-inflammatory interventions cannot fundamentally correct the abnormal metabolic reprogramming of immune cells in the diabetic microenvironment. Therefore, there is an urgent need to develop a composite biomimetic scaffold that can mimic the physical hierarchical structure of bone tissue, transmit surface electrical signals to the deep defect, and simultaneously synergistically regulate local immune metabolism and angiogenesis. Summary of the Invention

[0008] The purpose of this invention is to address the difficulties in repairing bone defects in the diabetic microenvironment, as well as the shortcomings of existing bone intervention materials in terms of structural homogenization, limited physical signal transmission, and insufficient ability to remodel the pathological microenvironment. This invention provides a piezoelectrically driven layered biomimetic scaffold for repairing diabetic bone defects, its preparation method, and its application, thereby solving the problems encountered in the aforementioned background art.

[0009] To achieve the above objectives, the technical solution of the present invention is as follows: A piezoelectrically driven layered biomimetic scaffold for repairing bone defects in diabetic patients is disclosed. The scaffold has a double-layer biomimetic structure, comprising an upper piezoelectric biomimetic periosteum that simulates the periosteum and a lower conductive filling hydrogel that simulates the defect filling matrix. The upper piezoelectric biomimetic periosteum is a poly(L-lactic acid) / zinc oxide composite electrospun nanofiber membrane (PLLA-ZnO-DFO) grafted with deferoxamine (DFO). The lower conductive filling hydrogel is a three-dimensional porous hydrogel composed of methacryloyl gelatin / transition metal carbide (GelMA / MXene) loaded with Ce-MOF@Arg nanoparticles.

[0010] In the above scheme, in the upper piezoelectric biomimetic bone membrane material, the ZnO nanoparticles in the PLLA-ZnO composite nanofiber membrane not only endow the fiber network with excellent electromechanical conversion (piezoelectric) properties, but also act as a heterogeneous nucleating agent to promote the crystallization of PLLA, effectively refining the diameter of the electrospun fibers and forming a dense and uniform biomimetic extracellular matrix network.

[0011] The DFO molecules are chemically bonded to the surface of the fiber membrane through surface oxygen plasma activation and carbodiimide (EDC / NHS) cross-linking chemistry, ensuring that the material exerts a long-lasting interfacial angiogenesis-promoting effect locally.

[0012] In the above scheme, in the lower conductive filling hydrogel material, the GelMA-MXene hydrogel system consists of two-dimensional MXene nanosheets cross-linked and assembled within a three-dimensional gel matrix, constructing a continuous low-impedance electronic conduction network. This conductive network forms a physical coupling with the upper piezoelectric film, enabling the deep and lossless transmission of the piezoelectric microcurrent signal generated in the upper layer to the interior of the hydrogel (i.e., the defective core region) under ultrasonic stimulation.

[0013] The Ce-MOF@Arg nanoparticles are cerium-based metal-organic framework materials internally loaded with amino acid drugs (arginine, Arg). This Ce-MOF nanozyme exhibits a specific Ce³⁺ / Ce⁻ ratio. 4 Its mixed valence structure endows it with reversible redox cycling ability, which can efficiently remove excess reactive oxygen species (ROS) in the diabetic microenvironment.

[0014] A method for fabricating a piezoelectrically driven layered biomimetic scaffold for repairing bone defects in diabetic patients, specifically including the following steps: Step 1: Preparation of the upper piezoelectric biomimetic periosteum (PLLA-ZnO-DFO) (1) Weigh out poly-L-lactic acid (PLLA) and dissolve it in a solvent in which N,N-dimethylformamide (DMF) and dichloromethane are mixed in a volume ratio of 1:2. Stir magnetically until completely dissolved to prepare a PLLA precursor solution with a mass fraction of 10%.

[0015] (2) Zinc oxide (ZnO) nanoparticles were ultrasonically dispersed in a mixed solvent of DMF and acetone in equal volume ratio. Then, they were added to a PLLA precursor solution so that the mass ratio of ZnO to PLLA was 1% (i.e., 1 wt% doping amount). The solution was magnetically stirred overnight at room temperature to obtain a uniform spinning solution.

[0016] (3) Electrospinning was performed under the conditions of 15kV positive voltage and -2kV negative voltage, working distance of 10cm, and collecting drum speed of 300r / min. The collected fiber membrane was vacuum dried at room temperature for 2 hours to remove residual solvent and obtain PLLA-ZnO composite membrane.

[0017] (4) The PLLA-ZnO composite membrane was subjected to oxygen plasma surface treatment for 5 minutes, and then immersed in a crosslinking system containing EDC, NHS and MES buffer for 2 hours for activation. After removal, it was transferred to a 1% (w / w) solution of deferoxamine mesylate (DFO) for reaction in the dark for 5 hours. After the reaction, it was washed alternately with 0.1M disodium hydrogen phosphate solution and deionized water, and then vacuum dried to finally obtain the PLLA-ZnO-DFO piezoelectric biomimetic bone membrane.

[0018] Step 2: Preparation of Ce-MOF@Arg nanoparticles, the underlying core functional material (1) Dissolve 354 mg of terephthalic acid in 12 mL of DMF, and then add 4 mL of cerium ammonium nitrate aqueous solution (concentration of 0.533 M). The mixture was placed at 100°C for 60 minutes. After the reaction, the product was collected by centrifugation, washed three times each with DMF and anhydrous ethanol, and dried overnight in a 60°C oven to obtain Ce-MOF nanozyme carrier (Ce-UiO-66).

[0019] (2) Weigh 100 mg of the above Ce-MOF and disperse it in 30 mL of deionized water. Add 17.4 mg of arginine (Arg) powder, disperse it by ultrasonication, and stir continuously for 3 hours.

[0020] (3) A total of 21 mg of carbonyl diimidazole (CDI) was added to the system in three batches as a coupling agent. After reacting for 30 minutes, the precipitate was collected by centrifugation, washed three times with DMF and ethanol, and dried overnight at 60°C to obtain arginine-functionalized Ce-MOF@Arg nanoparticles.

[0021] Step 3: Assembly of the lower conductive hydrogel and the double-layer biomimetic scaffold (1) Synthesis of GelMA: 20 g of gelatin was dissolved in 200 mL of PBS buffer (pH 7.4, 60°C), and after cooling to 40°C, 16 mL of methacrylic anhydride was slowly added dropwise at a rate of 0.5 mL / min. The reaction was stirred at a constant temperature in a 40°C water bath for 3 hours, and then an equal volume of PBS was added to terminate the reaction. The product was placed in a dialysis bag with a molecular weight cutoff of 14,000 Da, dialyzed at 37°C for 7 days, and then freeze-dried to obtain the lyophilized GelMA polymer.

[0022] (2) Preparation of hydrogel precursor: GelMA was dissolved in deionized water at a mass ratio of 1:10, MXene nanosheets were added at a mass ratio of MXene to GelMA of 3:10, and after mixing, Ce-MOF@Arg nanoparticles prepared above were added. Finally, 0.5wt% of lithium phenyl-2,4,6-trimethylbenzoyl phosphate (LAP) was added as a photoinitiator and mixed evenly.

[0023] (3) Integrated assembly: The above hydrogel precursor solution is injected into a custom mold and pre-crosslinked by 408nm ultraviolet light; then, the PLLA-ZnO-DFO piezoelectric bionic bone membrane obtained in step one is laid flat and attached to the surface of the hydrogel, and irradiated with 408nm ultraviolet light a second time to ensure that the hydrogel network is fully crosslinked and forms a tight interfacial interlock with the upper fiber membrane, thus obtaining the final piezoelectric driven layered bionic scaffold.

[0024] An application of a piezoelectrically driven layered biomimetic scaffold for the repair of diabetic bone defects, wherein the scaffold is used as a piezoelectric material for the repair of diabetic bone defects, simulating the physical hierarchical structure of bone tissue.

[0025] Compared with the prior art, the beneficial effects of the present invention are: (1) In this scheme, there is an innovative layered biomimetic structural design. The scaffold highly simulates the regeneration interface of natural bone in terms of physical structure. The upper electrospun membrane simulates the highly vascularized and barrier-functional periosteum, and the lower three-dimensional porous hydrogel simulates the filling matrix of the core area of ​​bone defect, realizing the spatial modular integration of tissue engineering materials in terms of structure and function.

[0026] (2) In this scheme, a piezoelectric-conductive cascade electrophysiological network was constructed, which broke through the bottleneck that the electrical signals of traditional piezoelectric materials could not penetrate into the core of the defect. Under ultrasound stimulation, the electromechanical conversion of the upper PLLA-ZnO film generates local piezoelectric microcurrents; the continuous three-dimensional conductive network constructed by MXene in the lower hydrogel acts as a signal relay station, efficiently and deeply transmitting the surface electrical signals to the core area of ​​the defect, thus constructing a biomimetic electrophysiological microenvironment conducive to bone regeneration.

[0027] (3) In this scheme, the material system achieves spatially differentiated controlled drug release through spatially specific targeted delivery and microenvironment remodeling. The DFO released from the upper membrane effectively recruits endothelial cells and promotes angiogenesis network reconstruction at the superficial surface of the defect; the lower hydrogel, through the unique Ce³⁺ / Ce of Ce-MOF nanozymes in the core area of ​​the defect... 4 The mixed valence state efficiently scavenges excess reactive oxygen species (ROS) induced by high glucose. Simultaneously, the arginine (Arg) released at an accelerated rate under ultrasound response can drive macrophages to undergo metabolic reprogramming from the glycolysis-dependent pro-inflammatory M1 type to the oxidative phosphorylation-dominated pro-repair M2 type, fundamentally reversing the persistent inflammatory microenvironment of diabetes.

[0028] (4) In this scheme, there are excellent physical and chemical properties and controllable degradation. The scaffold material exhibits excellent mechanical compressive strength, low impedance conductivity, and ultrasound-responsive on-demand degradation and drug release behavior that is adapted to the bone tissue healing cycle, and has extremely high clinical translation potential. Attached Figure Description

[0029] The disclosure of this invention is illustrated with reference to the accompanying drawings. It should be understood that the drawings are for illustrative purposes only and are not intended to limit the scope of protection of this invention. In the drawings, the same reference numerals are used to refer to the same parts. Wherein: Figure 1 This is a schematic diagram illustrating the fabrication of the piezoelectrically driven layered biomimetic structure of the present invention.

[0030] Figure 2 This is a diagram showing the gross morphology and cross-sectional structure of a double-layered biomimetic periosteum-bone filling scaffold.

[0031] Figure 3 The physicochemical characterization and piezoelectric properties of the PLLA-ZnO-DFO biomimetic periosteum are shown in the figure. Among them, a) SEM images of pure PLLA, PLLA-1%ZnO and PLLA-2%ZnO electrospun films; b) PFM phase loops of pure PLLA, PLLA-1%ZnO and PLLA-2%ZnO films; c) PFM amplitude butterfly curves of pure PLLA, PLLA-1%ZnO and PLLA-2%ZnO films; d) Transient voltage output of pure PLLA, PLLA-1%ZnO and PLLA-2%ZnO films; e) SEM image and corresponding EDS elemental mapping image of PLLA-1%ZnO-DFO film, showing the distribution of C, O, S and Zn; f) PFM phase loop of PLLA-1%ZnO-DFO film; g) PFM amplitude butterfly curve of PLLA-1%ZnO-DFO film; h) Transient voltage output of PLLA-1%ZnO-DFO film.

[0032] Figure 4 This is the FTIR spectrum of the electrospun film.

[0033] Figure 5 This is the XRD pattern of the electrospun membrane.

[0034] Figure 6Characterization images of Ce-MOF@Arg nanoparticles and GelMA / MXene / Ce-MOF@Arg hydrogels are shown below. a) SEM images of Ce-MOF and Ce-MOF@Arg nanoparticles; b) TEM images of Ce-MOF and Ce-MOF@Arg nanoparticles; c) EDS elemental mapping images of Ce-MOF@Arg nanoparticles, showing the distribution of C, N, and Ce; d) SEM images of lyophilized GelMA, GelMA / MXene, and GelMA / MXene / Ce-MOF@Arg hydrogels; e) Solvent images of GelMA, GelMA / MXene, and GelMA / MXene / Ce-MOF@Arg hydrogels. f) Compression stress-strain curves of GelMA, GelMA / MXene, and GelMA / MXene / Ce-MOF@Arg hydrogels; g) Electrochemical impedance spectroscopy of GelMA, GelMA / MXene, and GelMA / MXene / Ce-MOF@Arg hydrogels; h) Degradation curves of hydrogels under PBS and PBS plus sonication conditions (n=3); i) Cumulative release curves of Arg under PBS and PBS plus sonication conditions (n=3); j) Cumulative release curves of DFO under PBS and PBS plus sonication conditions (n=3).

[0035] Figure 7 XRD characterization images of Ce-MOF and Ce-MOF@Arg.

[0036] Figure 8 The XPS full spectrum of Ce-MOF.

[0037] Figure 9 XPS spectrum of Ce-MOF@Arg.

[0038] Figure 10 This is an elemental mapping of the GelMA / MXene / Ce-MOF@Arg hydrogel. The scale bar is 50 μm.

[0039] Figure 11 This is a graph showing the mechanical compression test data of the hydrogel.

[0040] Figure 12 This is a graph showing the cell compatibility assessment of BMSCs. a, b) Live / dead cell staining and quantitative cell viability analysis after 48 hours of co-culture (n=5), scale bar = 200 μm. c) CCK-8 cell proliferation within 5 days (n=5).

[0041] Figure 13This image shows the cytocompatibility assessment of RAW264.7 macrophages. a, b) show live / dead cell staining and quantitative cell viability after 48 hours of co-culture (n=5), scale bar = 200 μm. c) shows CCK-8 cell proliferation over 5 days (n=5).

[0042] Figure 14 For GMCP UI Effects on macrophage polarization in vitro. a, b) Representative immunofluorescence images of CD86 and CD206 in macrophages treated with different methods. Scale bar: 100 μm. c, d) Quantitative analysis of CD86 and CD206 fluorescence intensity (n=5). e, f) TGF-β and TNF-α secretion levels measured by ELISA (n=5). gi) Flow cytometry analysis and corresponding quantitative proportions of M1 and M2 macrophages (n=5). j) Relative mRNA expression levels of Tnfa, Il1β, Tgfβ, and Il10 measured by RT-qPCR (n=4). *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the GM group; #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to the GM group. UI The groups showed significant differences.

[0043] Figure 15 For GMCP UI Graphs showing the effects of in vitro ROS clearance and mitochondrial metabolic reprogramming on BMSCs. a, e) Representative immunofluorescence images of intracellular ROS (green) in BMSCs after different treatments and quantitative assessment of ROS fluorescence intensity (n=5), scale bar = 200 μm. b, f) Representative fluorescence images of JC-1 staining and quantitative assessment of ROS fluorescence intensity (n=5), scale bar = 100 μm. c, g) Flow cytometry analysis of intracellular ROS levels using the DCFH-DA probe and corresponding quantitative results (n=5). d, h) Flow cytometry analysis of JC-1 staining and corresponding quantitative analysis (n=5). *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the GM group; #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to the GMCP group. UI The differences between the groups were statistically significant.

[0044] Figure 16 For GMCP UIImmunomodulation promotes osteogenic-angiogenic coupling. a, b) Representative images and quantitative analysis of ALP staining in BMSCs on day 7 of osteogenic induction (n=5), scale bar = 200 μm. c, d) Representative images and quantitative analysis of ARS staining on day 21 of osteogenic induction (n=5), scale bar = 200 μm. e, f) Immunofluorescence staining and fluorescence intensity quantification of COL1 in BMSCs (n=5), scale bar = 100 μm. gk) Relative mRNA expression levels of osteogenic-related genes (Alpl, Runx2, Spp1, Col1a1, Bglap) determined by RT-qPCR (n=4). l) Representative immunofluorescence image of the endothelial marker CD31 in HUVECs (n=4), scale bar = 100 μm. m) Representative images of tube formation experiments (n=4), scale bar = 200 μm. n) Western blot analysis of HIF-1α and VEGF protein expression in HUVECs. *P<0.05, **P<0.01 and ***P<0.001 indicate significant differences compared to the GM group, #P<0.05, ##P<0.01 and ###P<0.001 indicate significant differences compared to the GMCP group. UI The groups showed significant differences.

[0045] Figure 17 Figure (a) shows the in vitro scratch healing experiment of HUVECs, and figure (b) shows the quantitative analysis of scratch closure rate (n=4). Scale bar = 500 μm. *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the GM group, and #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to GMCP. UI The differences between the groups were statistically significant.

[0046] Figure 18 The image shows the Transwell migration assay of HUVECs. Representative images of migrating cells stained with crystal violet (a) and the corresponding statistical quantitative analysis (b) (n=5) are shown. Scale bar = 100 μm. *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the GM group, and #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to the GMCP group. UI The groups showed significant differences.

[0047] Figure 19 For GMCP UIIn vivo bone regeneration assessment in a diabetic rat model of skull defects. a) Representative 3D Micro-CT reconstructed images of the skull defects at 4 and 8 weeks post-surgery. b, c) Quantitative analysis of BMD and BV / TV within the defect area (n=4). d) Representative H&E-stained histological sections of the defect site at 4 and 8 weeks post-surgery. *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the defect group; #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to the GMCP group. UI The groups showed significant differences.

[0048] Figure 20 Immunohistochemical analysis images are shown. a) Masson's trichrome staining of the defect area at 4 and 8 weeks. b, c) Immunohistochemical staining and corresponding quantitative analysis of pro-inflammatory TNF-α and healing-promoting TGF-β within the defect area (n=4), scale bar=100μm. df) Immunohistochemical staining and relative positive area quantification of osteogenic marker COL1A1 and angiogenesis marker CD31 at 4 and 8 weeks (n=4), scale bar=100μm. *P<0.05, **P<0.01, and ***P<0.001 indicate significant differences compared to the defect group; #P<0.05, ##P<0.01, and ###P<0.001 indicate significant differences compared to the GMCP group. UI The groups showed significant differences.

[0049] Figure 21 This is a diagram for the whole-body biosafety assessment. It shows H&E stained sections of major organs (heart, liver, spleen, lungs, and kidneys) 8 weeks post-implantation, with a scale bar of 200 μm. Detailed Implementation

[0050] To make the technical means, creative features, achieved objectives, and effects of this invention readily understandable, the invention will now be described in further detail with reference to the accompanying drawings. These drawings are simplified schematic diagrams, illustrating only the basic structure of the invention, and therefore only show the relevant components of the invention.

[0051] According to the technical solution of the present invention, without changing the essential spirit of the present invention, those skilled in the art can propose various interchangeable structural methods and implementations. Therefore, the following detailed embodiments and accompanying drawings are merely exemplary descriptions of the technical solution of the present invention, and should not be regarded as the entirety of the present invention or as a limitation or restriction of the technical solution of the present invention.

[0052] The technical solution of the present invention will be further described in detail below with reference to the accompanying drawings and embodiments.

[0053] like Figure 1 As shown, a piezoelectrically driven layered biomimetic scaffold for repairing bone defects in diabetic patients, its preparation method, and its application are described, specifically according to the following preparation scheme.

[0054] 1. Preparation scheme 1.1 Preparation of the upper piezoelectric biomimetic bone membrane, or more specifically, the preparation of the PLLA-ZnO-DFO electrospun membrane. Poly(L-lactic acid) (PLLA, Mw=80,000; Macklin) was dissolved in a mixed solvent of N,N-dimethylformamide (DMF) and dichloromethane in a volume ratio of 1:2, resulting in a final polymer concentration of 10% (w / w). The solution was magnetically stirred at room temperature for approximately 2 hours until the PLLA was completely dissolved. A pure PLLA membrane prepared from this solution was used as a control group.

[0055] To prepare the nanocomposite electrospun membrane, ZnO was first dispersed in a mixed solvent of DMF and acetone in equal volume ratios, resulting in a final concentration of 10% (w / w). Subsequently, ZnO was introduced into the PLLA precursor solution at mass ratios of 10 wt% ZnO-PLLA and 20 wt% ZnO-PLLA. The resulting mixtures were magnetically stirred overnight at room temperature to ensure uniform dispersion and thorough mixing.

[0056] The electrospinning solution was loaded into a 10 mL standard syringe equipped with a 20G flat-tipped needle and fed at a constant rate of 0.04 mm / s. A positive voltage of 15 kV was applied between the needle tip and the receiver at a working distance of 10 cm. The receiver was wrapped with aluminum foil and maintained at −2 kV. Electrospinning was performed at room temperature and a relative humidity of 40–80%, with the roller receiver rotation speed set at 300 r / min. After spinning, all films were vacuum dried at room temperature for 2 hours to remove residual solvent.

[0057] To fix DFO, the electrospun membrane was first treated with oxygen plasma for 5 minutes, then immersed in a crosslinking solution containing EDC-NHS-MES for 2 hours. Subsequently, the membrane was transferred to a 1% deferroamine methanesulfonate solution and reacted for 5 hours, allowing the carboxyl groups on PLLA to couple with the amino groups of deferroamine methanesulfonate. After the reaction, the membrane was washed three times sequentially with 0.1M disodium hydrogen phosphate solution and deionized water, and then vacuum dried to finally obtain the PLLA-ZnO-DFO piezoelectric biomimetic bone membrane.

[0058] 1.2 Preparation of GelMA-MXene-Ce-MOF@Arg hydrogel, which is the preparation of the underlying core functional material. 1.2.1 Synthesis of MOF@Arg A total of 354 mg of terephthalic acid was dissolved in 12 mL of N,N-dimethylformamide (DMF), followed by the addition of 4 mL of cerium ammonium nitrate solution (0.533 M). The reaction mixture was kept at 100°C for 60 minutes. The resulting product was collected by centrifugation, washed three times with DMF and ethanol, and dried overnight in an oven at 60°C to obtain the Ce-MOF nanozyme carrier, namely Ce-UiO-66.

[0059] To load arginine, 100 mg Ce-UiO-66 was dispersed in 30 mL of deionized water, followed by the addition of 17.4 mg arginine. After ultrasonic dispersion, the suspension was stirred for 3 hours. Then, 21 mg carbonyl diimidazole (CDI) was added in three portions as a coupling agent. After reacting for 30 minutes, the product was collected by centrifugation, washed three times with DMF and ethanol, and dried overnight at 60°C to obtain Ce-UiO-66 / Arg.

[0060] 1.2.2 Synthesis of GelMA and Preparation of MXene / GelMA Hydrogel The synthesis of methacrylamide gelatin (GelMA) is as follows: Briefly, 20 g of gelatin was dissolved in 200 mL of phosphate-buffered saline (PBS, pH 7.4) at 60°C with stirring until completely dissolved. The temperature was then lowered to 40°C, and 16 mL of methacrylic anhydride was slowly added at a rate of 0.5 mL / min. The reaction was carried out with continuous stirring in a 40°C water bath for 3 hours. To terminate the methacrylamide reaction, the mixture was diluted with 200 mL of PBS. The resulting solution was transferred to a dialysis bag with a molecular weight cutoff of 14,000 Da and dialyzed at 37°C for 7 days. After dialysis, the product was freeze-dried to obtain GelMA.

[0061] To prepare the hydrogel, GelMA was dissolved in deionized water at a weight ratio of 1:10, and then MXene powder was added at a mass ratio of 3:10 (MXene-GelMA). After thorough mixing, 0.5 wt% (relative to the total mass) of LAP was added as a photoinitiator. The precursor solution was then poured into a mold and photocrosslinked under 408 nm UV irradiation.

[0062] To prepare the biomimetic periosteum-bone construct, a hydrogel precursor was first laid on the surface of a mold and pre-gelled under 408 nm ultraviolet light. After demolding, the hydrogel side was irradiated again to ensure complete cross-linking. A second 408 nm ultraviolet irradiation was performed to ensure complete cross-linking of the hydrogel network and the formation of a tight interfacial interlock with the upper fibrous membrane, thus obtaining the final piezoelectrically driven layered biomimetic scaffold.

[0063] 1.3 Characterization The surface morphology of the samples was observed using a scanning electron microscope (SEM, model HITACHI Regulus 8100). All samples were sputter-coated with gold prior to imaging. Elemental composition and distribution were analyzed using energy-dispersive X-ray spectroscopy (EDS), and elemental mapping images were acquired using an IXRF system (model 550i). The microstructure of the samples was further examined using a transmission electron microscope (TEM, model JEM-2100F).

[0064] The crystal structure of the samples was characterized using X-ray diffraction (XRD) with CuKα radiation (λ = 0.15406 nm) at a working current of 40 mA and a voltage of 40 kV. Chemical functional groups were analyzed using Fourier transform infrared spectroscopy (FT-IR, Thermo Fisher Scientific Nicoleti S20). Surface elemental composition and oxidation state were examined using X-ray photoelectron spectroscopy (XPS) under ultra-high vacuum, and full-spectrum and high-resolution spectra were collected for elemental identification and quantification.

[0065] Piezoelectric microscopy (PFM) images were acquired using a Bruker Dimension Icon system. The piezoelectric output signal was recorded using a Tektronix MSO545-BW-1000 oscilloscope. Surface wettability was assessed by static water contact angle measurement using an IC2000D2G contact angle meter.

[0066] The swelling behavior of the hydrogel was evaluated in PBS at 37°C. Briefly, equal volumes of hydrogel samples were freeze-dried and weighed to obtain a dry weight (Wd). The samples were then immersed in PBS at 37°C for different time intervals. At each time point, excess liquid on the surface was gently removed with filter paper, and the swollen sample was weighed (Wt). The swelling ratio was calculated using the following formula: Swelling ratio = (Wt − Wd) / Wd.

[0067] The tensile properties of hydrogels were evaluated using a CMT6103 universal testing machine (MTS Industrial System) via uniaxial tensile testing. Hydrogel samples were cast into standard dumbbell-shaped molds according to ASTM standards. After curing, the samples were carefully removed and mounted on the testing machine, where tensile force was applied at a constant strain rate until fracture, thus obtaining stress-strain curves.

[0068] Rheological measurements (including viscosity and viscoelasticity) were performed at 25°C. Shear rates varied from 0.1 to 100 s⁻¹. Storage modulus (G′) and loss modulus (G″) of GelMA and the corresponding hydrogel scaffolds were measured at 25°C with an angular frequency range of 0.1–100 rads⁻¹.

[0069] 1.4 In vitro biocompatibility and targeted cell migration assessment Rat bone marrow mesenchymal stem cells (BMSCs) and RAW264.7 macrophages were isolated and cultured in α-MEM and DMEM (Thermo Fisher Scientific, Waltham, MA, USA) supplemented with 10% fetal bovine serum and 1% penicillin / streptomycin, respectively, and stored in a humidified incubator at 37°C and 5% CO2. To assess cell compatibility, cells were seeded in multi-well plates and co-cultured with scaffold extract. After 48 hours of co-culture, cell viability was observed under an inverted fluorescence microscope (Zeiss, Oberkochen, Germany) using a live / dead cell staining kit (Beyotime, Haimen, China). Cell proliferation kinetics were quantitatively determined using a cell counting kit-8 (CCK-8, Beyotime), and optical density was measured at 450 nm using a microplate reader.

[0070] To evaluate the scaffold's ability to recruit endogenous endothelial cells, scratch healing and Transwell migration assays were performed. For the scratch healing assay, standardized linear scratches were created on a confluent human umbilical vein endothelial cell (HUVEC) monolayer using sterile 200 μL pipette tips. Cell migration to the exposed area was recorded by photography at 0, 6, and 12 hours. For the Transwell assay, HUVECs seeded in the upper chamber migrated to the scaffold extract in the lower chamber within 24 hours. Cells that migrated to the lower membrane surface were fixed, stained with 0.5% crystal violet, and quantitatively analyzed using ImageJ software to determine cell recruitment efficiency.

[0071] 1.5 Macrophage polarization assessment To simulate the pathological microenvironment of diabetes in vitro, RAW264.7 macrophages were cultured in high-glucose medium (35 mM glucose) supplemented with lipopolysaccharide (LPS, 1 μg / mL). Cells were co-cultured with scaffold extract and subjected to exogenous sonication daily to activate the piezoelectric effect. Phenotypic transitions were assessed by immunofluorescence staining using primary antibodies against CD86 (a classic pro-inflammatory M1 marker) and CD206 (a pro-repair M2 marker), with nuclear counterstaining performed using DAPI. Images were acquired using an inverted fluorescence microscope, and fluorescence intensity was quantified using ImageJ. Flow cytometry was used to further validate the accurate M1 / M2 phenotypic transition ratio.

[0072] In addition, quantitative analysis of pro-inflammatory inflammatory markers was performed using RT-qPCR. Tnfα , Il1β ) and anti-inflammatory ( Tgfβ , Il10 The relative mRNA expression levels of these biomarkers were determined. Simultaneously, the secretion profiles of these cytokines in the culture supernatant were measured using a specific ELISA kit.

[0073] 1.6 Evaluation of in vitro osteogenic and angiogenesis in co-culture systems A Transwell co-culture system was established to investigate the effects of scaffold-mediated macrophages on osteogenic and angiogenesis. Macrophages were seeded in the upper chamber and exposed to the scaffold, while BMSCs or HUVECs were cultured in the lower chamber. To assess osteogenic differentiation, BMSCs were cultured in osteogenic induction medium. On day 7, early osteogenic activity was observed using a BCIP / NBTALP staining kit and quantified using an ALP assay kit. On day 21, extracellular matrix mineralization was assessed using Alizarin Red S (ARS) staining; calcium nodules were dissolved in 10% cetylpyridinium chloride, and absorbance was measured at 560 nm.

[0074] Key osteogenic markers were determined by RT-qPCR. Runx2 , Col1a1 , Alpl , Spp1 and Bglap The relative mRNA expression of type I collagen (COL1, IHC-P 1:500, ABclonal) was measured. Immunofluorescence staining of type I collagen (COL1, IHC-P 1:500, ABclonal) further confirmed osteogenic differentiation.

[0075] Since the diabetic environment often induces severe oxidative stress, the DCFH-DA fluorescent probe was used to detect ROS levels in co-cultured cells, and the results were analyzed by flow cytometry. Mitochondrial functional integrity was assessed by monitoring membrane potential changes using JC-1 fluorescent staining.

[0076] To investigate angiogenesis potential, HUVECs co-cultured under identical conditions were collected and seeded onto a low-growth-factor matrix gel for tube formation experiments. After 6 hours, the formation of capillary-like networks was recorded and quantified. Finally, angiogenesis was confirmed by immunofluorescence staining with the endothelial-specific marker CD31.

[0077] 1.7 In vivo assessment of bone regeneration in diabetes All animal experiments were conducted in strict accordance with animal protection and institutional ethical guidelines. Six-week-old male Sprague-Dawley rats were fed a high-fat diet for four weeks, followed by a single intraperitoneal injection of streptozotocin (STZ, 35 mg / kg) to induce type 2 diabetes mellitus (T2DM). Rats with a sustained fasting blood glucose level above 16.7 mmol / L were classified as successful diabetes models. Under general anesthesia, critical-sized full-thickness defects (5 mm in diameter) were created bilaterally using a sterile trephine. Scaffolds were precisely implanted into the defects; untreated defects served as a blank control group.

[0078] At 4 and 8 weeks post-surgery, rats were euthanized, and skull specimens were carefully collected. High-resolution micro-computed tomography (Micro-CT) was used to assess the 3D morphology of newly formed bone. Bone mineral density (BMD) and bone volume fraction (BV / TV) in the defect area were quantitatively analyzed. After standard decalcification, specimens were sectioned and stained with hematoxylin-eosin (H&E) and Masson's trichrome to assess histological structure and collagen matrix deposition.

[0079] To verify the in vivo bone immunomodulatory efficacy, immunohistochemical (IHC) staining was performed to track the in situ expression of TNF-α, TGF-β, COL1A1, and CD31. Finally, H&E staining of major organs (heart, liver, spleen, lung, and kidney) confirmed the system's biocompatibility and verified that the implanted biomaterial did not have systemic toxicity.

[0080] 2. Results 2.1 Physicochemical properties and piezoelectric properties of PLLA-ZnO-DFO biomimetic periosteum To simultaneously meet the needs for structural support, local electrical stimulation, immune metabolic regulation, and pro-angiogenic activity during the repair of diabetic bone defects, an integrated bilayer biomimetic periosteum-bone filling scaffold was constructed. This system consists of an upper PLLA-ZnO-DFO electrospun nanofiber membrane and a lower GelMA-MXene-Ce-MOF@Arg functional hydrogel, designed to simulate the natural periosteum interface and the defect-filling matrix microenvironment, respectively.

[0081] Figure 1 This is a schematic diagram illustrating the fabrication of a piezoelectrically driven layered biomimetic structure system. (Example:) Figure 2 As shown, the resulting double-layer scaffold exhibits a complete overall architecture, clear layering, and good structural stability. Figure 3 Image a shows surface SEM images of fiber films with different ZnO doping ratios. Pure PLLA, PLLA / 1%ZnO, and PLLA-2%ZnO all formed continuous and uniform nanofiber networks, indicating that the incorporation of ZnO did not disrupt the electrospinning process.

[0082] It is noteworthy that the average fiber diameter gradually decreases with increasing ZnO content, indicating that ZnO nanoparticles alter the charge density and rheological behavior of the spinning solution, thereby enhancing jet stretching under an electric field and promoting fiber refinement.

[0083] The piezoelectric response of the thin film was further evaluated using piezoelectric microscopy (PFM). Figure 3 As shown in b, the PFM phase diagram of the ZnO-containing group exhibits a more pronounced phase switching than that of pure PLLA, indicating enhanced local piezoelectric activity after ZnO doping. Consistently, Figure 3The amplitude butterfly curve in c shows that the response of the ZnO-doped group is significantly higher than that of the pure PLLA group, further confirming the improved electromechanical conversion capability of the composite membrane.

[0084] Transient voltage output curve as shown Figure 3 As shown in d, the voltage output of both ZnO-doped groups was higher than that of pure PLLA, indicating that the incorporation of ZnO enhanced the macroscopic piezoelectric output of the film. Since the overall output of the 1% ZnO and 2% ZnO groups was comparable, 1% ZnO was chosen as the final doping ratio for subsequent studies.

[0085] Figure 3 The image shows the SEM morphology and elemental mapping of the PLLA-1%ZnO-DFO film. After DFO functionalization, the PZ1D film still maintains a continuous fiber morphology and an intact fiber network. EDS mapping shows that C, O, S, and Zn are uniformly distributed throughout the film. In particular, the presence of S further confirms the successful incorporation of DFO into the fiber system.

[0086] Figure 4 The FTIR spectra shown further validated the chemical composition of the films. All samples retained the characteristic absorption peaks of PLLA at 1085, 1183, and 1747 cm⁻¹, corresponding to the C–O–C and C=O vibrations of the polymer backbone, indicating that the PLLA backbone remained stable during complex formation and post-functionalization. In the DFO-modified group, an additional absorption peak associated with isohydroxamic acid esters appeared at approximately 1630 cm⁻¹, and a peak associated with sulfonates appeared at approximately 1030 cm⁻¹, confirming the successful immobilization of DFO on the fibrous membrane. These results are consistent with… Figure 3 The sulfur element mapping shown in e is consistent.

[0087] Figure 5 The XRD patterns further confirmed the successful incorporation of ZnO. With increasing ZnO content, characteristic diffraction peaks gradually appeared at 31.7°, 34.4°, and 36.2°, corresponding to the (100), (002), and (101) crystal planes of hexagonal wurtzite ZnO, respectively. Simultaneously, the α-phase diffraction peak of PLLA at approximately 16.5° became more pronounced, indicating that ZnO, acting as a heterogeneous nucleating agent, promoted the crystallization of PLLA. These data suggest that the incorporation of ZnO not only introduced a functional inorganic phase but also improved the crystal structure of the fibrous membrane.

[0088] The piezoelectric properties of DFO functionalized thin films, such as Figure 3 As shown in f–h. Figure 3 f shows that the PZ1D film retained a clear PFM phase diagram after DFO modification. Figure 3 g shows that the amplitude response remains strong. Figure 3hj further proved that the PZ1D film maintained a stable voltage output.

[0089] In summary, these results demonstrate that covalent DFO functionalization did not substantially impair the piezoelectric properties of the film. Therefore, the upper fibrous layer successfully combined pro-angiogenic function with piezoelectric activity, laying the foundation for constructing an ultrasound-activated biomimetic periosteum.

[0090] 2.2 Physicochemical properties of Ce-MOF@Arg drug-loaded nanoparticles To endow hydrogel phases with immunomodulatory capabilities, we synthesized and systematically characterized Ce-MOF@Arg nanoparticles. Figure 6 Image a shows SEM images of Ce-MOF and Ce-MOF@Arg. Both groups exhibit relatively uniform micro / nano structures, while Ce-MOF@Arg shows a more regular short rod-like or blocky morphology, indicating that arginine modification affects particle formation.

[0091] Figure 6 b shows the corresponding TEM images, which further confirm the structural integrity and good dispersibility of the two nanoparticle formulations. Figure 6 c shows the elemental mapping of Ce-MOF@Arg. C, N, and Ce are clearly and uniformly distributed throughout the particle. In particular, the presence of nitrogen directly confirms the successful introduction of Arg into the Ce-MOF system.

[0092] Figure 7 The XRD patterns shown further validate the crystal structure of the nanoparticles. The pristine Ce-MOF exhibits characteristic diffraction peaks consistent with the UiO-66 type framework, confirming the successful construction of the cerium-based MOF framework. After Arg modification, Ce-MOF@Arg largely retains the diffraction characteristics of Ce-MOF, with no obvious impurity peaks, indicating that Arg incorporation did not disrupt the original MOF framework. Notably, the diffraction peaks of Ce-MOF@Arg become sharper and more intense, suggesting that Arg participates in the regulation of crystal nucleation and growth, thereby inducing anisotropic crystal growth rather than forming a new crystal phase.

[0093] Figure 8 and Figure 9 The XPS spectra shown further confirm the successful functionalization of Arg. Ce-MOF@Arg exhibits a distinct N1s peak at approximately 400.2 eV, corresponding to the amino and guanidinyl groups of arginine. Furthermore, the high-resolution C1s spectrum displays a characteristic C–N peak at approximately 286.2 eV, further supporting the successful conjugation of Arg. High-resolution Ce3d spectra reveal Ce³⁺ and Ce⁻ in the material. 4The coexistence of ⁺ and ⁺ valence states demonstrates that Arg-functionalized nanoparticles retain redox-active cerium centers associated with oxygen vacancies. This mixed valence state characteristic facilitates reversible redox cycling and may enhance ROS scavenging capabilities, thus providing a structural basis for subsequent regulation of the inflammatory microenvironment.

[0094] 2.3 Compositional characteristics and physical properties of GelMA-MXene-Ce-MOF@Arg hydrogel As a defect-filling component, hydrogels are designed not only to provide structural support but also to create a local microenvironment conducive to cell growth and electrical signal transmission. Therefore, their microstructure, swelling behavior, mechanical properties, electrical properties, degradation, and release behavior were systematically evaluated.

[0095] Figure 6 Image d shows SEM images of lyophilized GelMA, GelMA-MXene, and GelMA-MXene-Ce-MOF@Arg hydrogels. All three groups exhibit a continuous three-dimensional porous network structure. With the addition of MXene and Ce-MOF@Arg, the pore walls gradually thicken and the pore size distribution becomes more uniform, indicating that the functional additives modulate the internal microstructure of the hydrogels. Figure 10 The elemental mapping further confirms that Ce-MOF@Arg is uniformly distributed throughout the hydrogel network, supporting its subsequent local bioactivity.

[0096] The swelling behavior of hydrogels, such as Figure 6 As shown in e. All groups exhibited rapid water absorption in the initial soaking stage, subsequently gradually reaching equilibrium. Compared to pure GelMA, the equilibrium swelling ratios of the GM and GMCe groups were significantly reduced, indicating that the incorporation of MXene and Ce-MOF@Arg increased the effective crosslinking density and made the hydrogel network more compact through physical interactions, space occupancy, and pore modulation. These findings are consistent with... Figure 6 The microstructural changes observed in d are consistent.

[0097] Figure 6 f shows the compressive stress-strain curves of the hydrogel. The GMCe group exhibits higher compressive strength than the G and GM groups, indicating that Ce-MOF@Arg significantly improves the mechanical stability of the hydrogel. This strengthening effect can be attributed to the reinforcing effect of Ce-MOF@Arg within the network and its ability to restrict polymer chain slippage, thereby improving the overall load-bearing capacity of the structure.

[0098] Figure 11The cyclic compression results shown further demonstrate that the composite hydrogel maintains good shape recovery and stable viscoelastic behavior during repeated load-unload cycles. These properties support its structural integrity under dynamic mechanical conditions, which is beneficial for its long-term application in complex physiological environments.

[0099] Figure 6 g shows the impedance spectra of the three hydrogels. Compared with pure GelMA, both the GM and GMC groups showed significantly lower impedance, indicating that MXene established a continuous conductive network within the hydrogel. This conductive network is expected to facilitate the transmission of local electrical signals generated by the upper piezoelectric membrane under ultrasound stimulation to deeper defect areas, thereby creating a more favorable electrophysiological microenvironment for cellular response.

[0100] These results indicate that the bilayer system possesses not only an upper layer with dynamogenic properties but also a lower layer with conductive characteristics. Degradation curves under different conditions are shown below. Figure 6 As shown in h.

[0101] In PBS, the hydrogel underwent gradual degradation, with approximately 20% of its mass remaining at day 28. Degradation was accelerated under the combined effects of ultrasound stimulation, suggesting that the local electrical effects generated by stimulation may promote the disruption of the hydrogel network. This tunable degradation behavior may facilitate time-controlled disappearance of the scaffold in vivo according to treatment needs.

[0102] Figure 6 i shows the cumulative release curve of Arg. Arg exhibits sustained release in PBS, while ultrasound stimulation further accelerates its release, especially in the early stages, indicating that the system has good stimulus-response release behavior. Figure 6 j shows the cumulative release curve of DFO. DFO also exhibits sustained release, with a faster release rate under ultrasound stimulation.

[0103] These data indicate that DFO fixed within the upper fiber membrane remains stable within the construct while still being able to release effectively in response to external stimuli.

[0104] In summary, this bilayer construct exhibits a stable microstructure, good mechanical and electrical properties, and tunable degradation and stimulus-responsive release behavior. The spatially layered incorporation of DFO in the upper fibrous membrane and Arg in the lower hydrogel enables coordinated time-release, providing a material basis for subsequent synergistic pro-angiogenic, anti-inflammatory, and osteogenic therapies.

[0105] To verify the biological function of the piezoelectrically driven layered biomimetic scaffold, we grouped them as follows: GM group: GelMa hydrogel + Mxene GMC group: GelMa hydrogel + Mxene + Ce-MOF@Arg GMCP group: GelMa hydrogel + Mxene + Ce-MOF@Arg + ZnO-containing PLLA spun membrane GMCPUI group: GelMa hydrogel + Mxene + Ce-MOF@Arg + PLLA spun membrane containing ZnO + DFO + ultrasonic treatment 2.4 Excellent biocompatibility Excellent biocompatibility is an indispensable prerequisite for in vivo implantation. This study comprehensively evaluated the cytotoxicity of the scaffold to rat bone marrow mesenchymal stem cells (BMSCs) and RAW264.7 macrophages. Live / dead fluorescence staining showed that after 48 hours of co-culture, all groups exhibited a dense distribution of surviving (green) cells, with negligible dead cells (red), and cell viability remained well above 90%. Figure 12 a,b& Figure 13 a, b). CCK-8 proliferation kinetics further confirmed that all scaffolds supported robust cell proliferation at 1, 3, and 5 days, with no significant differences between groups. Figure 12 c& Figure 13 c).

[0106] 2.5GMCP UI Driving macrophages to polarize towards the pro-repair M2 phenotype Diabetic patients suffer from long-term oxidative stress and chronic inflammation caused by hyperglycemia, among which persistently activated M1 macrophages are a core pathological driver of delayed bone healing. To evaluate GMCP... UI To investigate the immunomodulatory effects of the scaffold, we simulated the harsh diabetic microenvironment in vitro using a high-glucose culture medium supplemented with LPS.

[0107] Immunofluorescence staining clearly confirmed significant phenotypic reversal: compared with the GM group, which primarily expressed the M1 marker CD86, GMCP... UI The green CD86 fluorescence in the group was greatly suppressed ( Figure 14 a,c). Conversely, the red fluorescence representing the pro-repair M2 marker CD206 showed strong upregulation ( Figure 14 b,d).

[0108] Absolute quantification by flow cytometry strongly reinforced this observation, indicating that GMCP UI The group achieved the most significant reversal of the M2 / M1 cell ratio ( Figure 14 gi). RT-qPCR and ELISA confirmed that GMCP UI This leads to a sharp downregulation of the transcriptional levels and secretion concentrations of pro-inflammatory cytokines (Tnfa, Il1b), while the expression of anti-inflammatory and pro-remission factors (Tgfb, Il10) is strongly activated. Figure 14 e,f,j).

[0109] 2.6GMCP UI Promoting osteogenic-angiogenic coupling through immunomodulation The high-glucose microenvironment of diabetes not only causes immune dysregulation but also triggers intracellular reactive oxygen species (ROS) overload in BMSCs, leading to severe mitochondrial dysfunction and osteogenic impairment. Flow cytometry combined with DCFH-DA staining showed that GMCP... UI The scaffold can effectively remove high glucose-induced intracellular ROS in BMSCs. Figure 15 a, c, e, g). Importantly, JC-1 staining confirmed GMCP. UI Intervention rescued BMSCs fluorescence signals from the green monomer state (mitochondrial damage) to the red J-polymer state (healthy, high membrane potential). Figure 15 (b, d, f, h) demonstrates that the scaffold can actively protect mitochondrial integrity.

[0110] Bone regeneration is a deeply coupled systemic process involving osteogenic and angiogenesis. To investigate the effect of immune-regulated macrophages on this coupling, a Transwell co-culture system was employed.

[0111] On day 7 of osteogenic induction, alkaline phosphatase (ALP) staining was performed on GMCP. UI The co-culture group showed a strong blue-purple precipitate, and quantitative detection confirmed that it reached the peak of ALP activity. Figure 16 a, b). By day 21, Alizarin Red S (ARS) staining showed dense, compact red calcium nodules, indicating a high degree of extracellular matrix mineralization. Figure 16 c,d).

[0112] RT-qPCR results showed that in GMCP UI In this group, the mRNA levels of key osteogenic genes (Runx2, Col1a1, Alpl, Spp1, Bglap) were upregulated several times. Figure 16 gk), which was further confirmed by the strong immunofluorescence expression of COL1 ( Figure 16 e,f).

[0113] Rapid recruitment of endogenous vascular endothelial cells (GMCPs) is crucial in the early stages of bone defect repair. This study evaluated GMCP's potential for pro-angiogenic cell recruitment using in vitro scratch healing and Transwell migration assays. The scratch healing assay showed that GMCPs, at 6 and 12 hours, significantly increased GMCP recruitment. UI HUVECs in the co-culture group exhibited the highest centripetal migration rate and achieved a significantly better scratch closure rate than other groups. Figure 17 ).

[0114] Transwell experiments yielded consistent results at GMCP. UI In this group, the number of cells migrating through the porous polycarbonate membrane reached its maximum value. Figure 18 These findings confirm that the release of active ingredients, such as DFO and Ce ions, exerts a powerful chemotactic effect, driving directed cell migration and laying the cellular foundation for rapid reconstruction of vascular networks.

[0115] Simultaneously, we evaluated angiogenesis in HUVECs under the same co-culture conditions using in vitro tube formation experiments. Under the synergistic effect of continuously released DFO and beneficial M2 macrophage-derived cytokines, HUVECs rapidly assembled a mature, highly complex capillary-like network. Figure 16 In addition, to assess the regulatory effects of the intervention on vascular endothelial cells at the molecular level, Western blot analysis was performed on key proteins related to angiogenesis in HUVECs.

[0116] The results showed that, compared with the control group, the protein expression level of HIF-1α in the experimental group of HUVECs was significantly upregulated. Figure 16 Meanwhile, the expression of vascular endothelial growth factor (VEGF), a core downstream pro-angiogenic factor regulated by HIF-1α, was also significantly increased. Figure 16 Furthermore, CD31, a hallmark protein reflecting endothelial cell functionalization and the potential for angiogenesis, also showed significant upregulation. Figure 16 l).

[0117] In summary, these changes in protein levels confirm that the material system can effectively activate HIF-1α and its associated angiogenesis signaling pathways within HUVECs. These results highlight the effectiveness of GMCP. UI The scaffold can effectively coordinate the bone-immune-vascular axis, perfectly mimicking the physiological cascade of natural bone repair.

[0118] 2.8GMCP UI Composite scaffolds significantly accelerated the regeneration of critical-sized skull defects in diabetic rats. To verify GMCP UI The clinical translation potential of the scaffold was demonstrated in a critical-sized skull defect model in diabetic rats. Figure 19 a). Micro-CT scans and 3D reconstructions at 4 and 8 weeks post-implantation showed that the defect group exhibited almost no inward bone growth at the defect margin. Figure 19 b). Conversely, implanting GMCP. UI The defect was covered by extensive, continuous, and dense new bone tissue. Quantitative analysis showed that at each time point, GMCP... UIThe bone mineral density (BMD) and bone volume fraction (BV / TV) of the group were significantly higher than those of all other groups. Figure 19 c,d).

[0119] Using H&E ( Figure 19 d) and Masson's trichrome staining ( Figure 20 a) The organizational morphology assessment further reveals that GMCP UI The group not only formed a continuous bone bridge at the defect site, but also developed a highly ordered collagen fiber network and mature osteocyte lacunae, without any severe fibrous encapsulation or adverse foreign body reaction.

[0120] Immunohistochemical (IHC) staining of the defect area confirmed this pro-regenerative microenvironment: high expression of osteogenic and angiogenesis markers (COL1A1 and CD31) was observed. Figure 20 df), the core inflammatory factor TNF-α was completely inhibited, while TGF-β expression was significantly upregulated ( Figure 20 b,c).

[0121] Finally, a systemic H&E assessment of the major vital organs (heart, liver, spleen, lungs, and kidneys) showed no pathological damage or systemic inflammatory infiltration. Figure 21 Due to its superior in vivo bone immune modulation capabilities and excellent biocompatibility, GMCP... UI The biomimetic composite scaffold successfully reversed the damaged bone regeneration microenvironment in diabetic patients, highlighting its enormous potential for clinical application.

[0122] It should be noted that, in this document, the terms "comprising," "including," or any other variations thereof are intended to cover non-exclusive inclusion, such that a process, method, article, or apparatus that comprises a list of elements includes not only those elements but also other elements not expressly listed, or elements inherent to such a process, method, article, or apparatus. These undisclosed elements are all prior art known to those skilled in the art.

[0123] The specific embodiments described above further illustrate the purpose, technical solution, and beneficial effects of the present invention. It should be understood that the above description is only a specific embodiment of the present invention and is not intended to limit the scope of protection of the present invention. Any modifications, equivalent substitutions, improvements, etc., made within the spirit and principles of the present invention should be included within the scope of protection of the present invention.

Claims

1. A piezoelectrically driven layered biomimetic scaffold for repairing bone defects in diabetic patients, characterized in that: The scaffold is a double-layer biomimetic structure, comprising an upper piezoelectric biomimetic periosteum simulating the periosteum and a lower conductive filling hydrogel simulating the defect filling matrix. The upper piezoelectric biomimetic periosteum is a nanofiber membrane, and the lower conductive filling hydrogel is a composite three-dimensional porous hydrogel.

2. The piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects according to claim 1, characterized in that: The upper piezoelectric biomimetic bone membrane is a composite electrospun nanofiber membrane made of poly-L-lactic acid grafted with deferroamine and zinc oxide, and the lower conductive filling hydrogel is a composite three-dimensional porous hydrogel of methacrylamide gelatin-transition metal carbide loaded with Ce-MOF@Arg nanoparticles.

3. The piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects according to claim 2, characterized in that: The composite electrospun nanofiber membrane is composed of PLLA, ZnO, and DFO. ZnO nanoparticles have piezoelectric properties for the fiber network and also act as heterogeneous nucleating agents to promote the crystallization of PLLA. DFO molecules are activated by surface oxygen plasma and cross-linked with carbodiimide.

4. The piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects according to claim 2, characterized in that: The Ce-MOF@Arg nanoparticles contain two-dimensional MXene nanosheets that are cross-linked and assembled within a three-dimensional gel matrix to construct a continuous low-impedance electronic conduction network, which is physically coupled with the upper piezoelectric biomimetic periosteum.

5. The piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects according to claim 2, characterized in that: The Ce-MOF@Arg nanoparticles contain cerium-based metal-organic framework materials that internally load amino acid-based drugs.

6. A method for fabricating a piezoelectrically driven layered biomimetic scaffold for repairing diabetic bone defects according to any one of claims 1-5, characterized in that, Includes the following steps: Step 1: First, prepare the upper piezoelectric biomimetic periosteum; Step 2: Next, Ce-MOF@Arg nanoparticles are prepared as the lower layer material, and then GelMA is synthesized to prepare the lower conductive hydrogel. Step 3: Finally, assemble the lower conductive hydrogel with the double-layer biomimetic scaffold.

7. The preparation method according to claim 6, characterized in that, Step one specifically includes the following steps: Step F11: Weigh out poly-L-lactic acid and dissolve it in a solvent in which N,N-dimethylformamide and dichloromethane are mixed in a volume ratio of 1:

2. Stir magnetically until completely dissolved to prepare a 10% poly-L-lactic acid precursor solution. Step F12: Disperse zinc oxide nanoparticles ultrasonically in a mixed solvent of N,N-dimethylformamide and acetone in equal volume ratio; then add it to a precursor solution of poly-L-lactic acid, so that the mass ratio of zinc oxide to poly-L-lactic acid is 1%, and stir magnetically overnight at room temperature to obtain a uniform spinning solution. Step F13: Electrospinning is performed under the conditions of 15kV positive voltage and -2kV negative voltage, working distance of 10cm, and collecting drum speed of 300r / min. The collected fiber membrane is vacuum dried at room temperature for 2 hours to remove residual solvent and obtain PLLA-ZnO composite membrane. Step F14: The PLLA-ZnO composite membrane is subjected to oxygen plasma surface treatment for 5 minutes, and then immersed in a crosslinking system containing EDC, NHS and MES buffer for 2 hours for activation; after removal, it is transferred to a 1% methanesulfonic acid deferoxamine solution for reaction in the dark for 5 hours; after the reaction, it is washed alternately with 0.1M disodium hydrogen phosphate solution and deionized water, and vacuum dried to finally obtain the PLLA-ZnO-DFO piezoelectric bionic bone membrane.

8. The preparation method according to claim 6, characterized in that, Step two specifically includes the following steps: Step F21: Dissolve 300-380 mg of terephthalic acid in 10-15 mL of N,N-dimethylformamide, then add 4 mL of cerium ammonium nitrate aqueous solution; place the mixture at 100°C for 60 minutes; after the reaction, centrifuge to collect the product, wash three times each with N,N-dimethylformamide and anhydrous ethanol, and dry in an oven at 55-65°C overnight to obtain Ce-MOF nanozyme carrier; Step F22: Weigh 90-110 mg of the above Ce-MOF and disperse it in 25-35 mL of deionized water. Add 15-20 mg of arginine powder, disperse it by ultrasonication, and stir continuously for 3 hours. Step F23: Add a total of 15-25 mg of carbonyl diimidazole as a coupling agent to the system in three batches; after reacting for 30 minutes, centrifuge to collect the precipitate, wash it three times with N,N-dimethylformamide and ethanol, and dry it overnight at 55-65°C to obtain arginine-functionalized Ce-MOF@Arg nanoparticles.

9. The preparation method according to claim 6, characterized in that, Step three specifically includes the following steps: Step F31, Synthesis of GelMA: Dissolve 15-25g of gelatin in 180-230mL of PBS buffer, cool to 40°C, and slowly add 10-20mL of methacrylic anhydride at a rate of 0.3-0.8mL / min; stir the mixture in a water bath at 35-45°C for 3 hours, and then add an equal volume of PBS to terminate the reaction; place the product in a dialysis bag with a molecular weight cutoff of 14000Da, dialyze at room temperature for 7 days, and freeze-dry to obtain the lyophilized GelMA polymer; Step F32, Preparation of hydrogel precursor: Dissolve the lyophilized GelMA polymer in deionized water at a mass ratio of 1:10, add MXene nanosheets at a mass ratio of MXene to GelMA of 3:10, mix well, add the Ce-MOF@Arg nanoparticles prepared above, and finally add 0.5wt% of lithium phenyl-2,4,6-trimethylbenzoyl phosphate as a photoinitiator, and mix well; Step F33, Integrated Assembly: The above-mentioned hydrogel precursor solution is injected into a custom mold and pre-crosslinked by irradiation with ultraviolet light; then, the PLLA-ZnO-DFO piezoelectric bionic scaffold prepared in step one is laid flat and attached to the surface of the hydrogel, and subjected to a second ultraviolet light irradiation to ensure that the hydrogel network is fully crosslinked and forms a tight interfacial interlock with the upper fiber membrane, thus obtaining the final piezoelectrically driven layered bionic scaffold.

10. The application of a piezoelectrically driven layered bionic scaffold for repairing diabetic bone defects according to any one of claims 1-5, characterized in that: The scaffold is used as a piezoelectric material for repairing diabetic bone defects, simulating the physical hierarchical structure of bone tissue.

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