Preparation method of bio-functionalized nano hydrogel for skull regeneration

Abstract

The invention relates to the technical field of skull regeneration, and particularly discloses a preparation method of bio-functionalized nano hydrogel for skull regeneration, which comprises the following steps: dissolving deferoxamine in deionized water, and adding magnesium lithium silicate nanosheets into the solution to form magnesium lithium silicate / deferoxamine nanoparticles; the preparation method comprises the following steps: mixing pure polyvinyl alcohol and deionized water in a flask to form a pure polyvinyl alcohol aqueous solution, and preparing polyvinyl alcohol hydrogel; mixing a polyvinyl alcohol aqueous solution, tannic acid and the magnesium lithium silicate / deferoxamine nanoparticles in a flask to form a polyvinyl alcohol / magnesium lithium silicate / deferoxamine mixture, and preparing the nano hydrogel. According to the preparation method of the bio-functionalized nano hydrogel for skull regeneration, the polyvinyl alcohol hydrogel is used as a skull defect repair base material, active ions and drugs in the deferoxamine modified magnesium lithium silicate nanosheets are delivered through the hydrogel, the long-term requirements of cellular reaction and skeleton remodeling are met, and the skull regeneration effect is improved. Vascularization is promoted; and osseointegration is accelerated.

Description

Technical Field

[0001] The invention relates to the technical field of bone regeneration, and in particular to a method for preparing a biofunctionalized nano hydrogel for skull regeneration. Background Art

[0002] The repair of craniomaxillofacial bone defects caused by traumatic brain injury remains a major challenge. Despite the positive results, the application of autologous / allogeneic transplantation in reconstructive surgery is limited due to the huge costs associated with bone transplantation, limited bone sources, and potential complications. To address these issues, various substitutes including titanium mesh, polyetheretherketone (PEEK), calcium phosphate cement, and polymethylmethacrylate (PMMA) have been developed. However, these substitutes have disadvantages such as limited flexibility, insufficient bone integration, significant tissue friction, and high Young's modulus. Therefore, their application is affected by their inability to adapt to the dynamic changes of intracranial tissues and their unsuitability for immediate implantation after craniectomy. In addition, the formation of skull bones is closely related to the development of capillaries, and the formation of highly vascularized tissue is essential for providing the necessary oxygen and nutrients for bone regeneration and bone integration. This emphasizes the critical interaction between bone formation and vascularization in skull development. Although the currently studied skull regeneration strategies have achieved certain results, they are still insufficient in generating highly vascularized new bone formation. Therefore, there is an urgent need for a functional implant with appropriate mechanical strength and bioactivity that can adapt to the intracranial microenvironment for immediate implantation and fixation in the defect site after craniectomy. At the same time, it should also have the potential to mobilize human umbilical vein endothelial cells (HUVECs) and mesenchymal stem cells (MSCs) to promote skull tissue regeneration.

[0003] Pure polyvinyl alcohol (PVA) solutions can be converted into polymer hydrogels by continuous freeze-thaw cycles to form microcrystalline structures without the use of chemical crosslinkers that may pose a risk of cytotoxicity. In recent years, as a new type of high-performance polymer system, PVA hydrogels have attracted widespread attention for their applications in biopharmaceuticals and biomedicine. Due to its excellent biocompatibility, remarkable mechanical properties, low cost, and ease of processing, it has become an important material for tissue engineering and drug delivery systems. However, unmodified PVA hydrogels are associated with reduced cell adhesion and limited biological functions in wound repair due to their ultra-smooth surface. These factors have severely hindered their potential application in bone regeneration. Recent studies have shown that modifying original PVA hydrogels with bioactive molecules can effectively improve their adhesion and achieve sustained release of bioactive molecules, thereby promoting tissue repair.

[0004] Lithium magnesium silicate nanosheets (LAP), as an emerging two-dimensional biomaterial, present exciting possibilities in biomedical engineering. LAP is known for its excellent biocompatibility, dual charge properties, unique shape, and high surface area-to-volume ratio. These characteristics make LAP suitable for use in a variety of fields, including the delivery of guest molecules such as drugs, proteins, and genes, as well as tissue repair. The degradation products of LAP are non-cytotoxic silicon (Si), magnesium (Mg), and lithium (Li) ions, which act as bioactive factors in the physiological microenvironment. Modification of PVA hydrogels with LAP as an osteogenic bioactive factor will produce a variety of beneficial effects.

[0005] Furthermore, vascularization of bones is crucial for their health and regenerative capacity. Inadequate blood supply often limits new bone formation and delays the healing process. In light of this, the use of deferoxamine (DFO) in combination with osteogenic bioactive molecules is an attractive strategy. It has been shown to increase the levels of hypoxia-inducible factor 1-α (HIF-1α) and activate downstream angiogenesis-related genes. Among the different delivery systems that can be used to carry positively charged DFO in biomedical applications, the use of LAP as a film-forming polymer is a suitable choice. The edges of LAP are positively charged, while the surface of the nanosheets is negatively charged. LAP can be used as a smart carrier for the delivery of bioactive factors and antibodies, which provides a minimalist approach to deliver positively charged DFO through the negatively charged surface of LAP and can cleverly achieve the combination of DFO with bioactive factors. The bioactive LAP and DFO materials were integrated into a biocompatible PVA hydrogel with remarkable mechanical properties and provide a new design idea for comprehensive bone regeneration management. Summary of the invention

[0006] The purpose of the present invention is to provide a method for preparing a biofunctionalized nanohydrogel for skull regeneration, using a polyvinyl alcohol hydrogel system as a substrate for skull defect repair, and delivering active ions and drugs in a deferoxamine-bound lithium magnesium silicate nanomaterial through biocompatibility and adhesive tannic acid, thereby meeting the long-term needs of cell response and bone remodeling, promoting rapid vascular invasion and accelerating bone integration.

[0007] To achieve the above object, the present invention provides a method for preparing a biofunctionalized nanohydrogel for skull regeneration, comprising the following steps:

[0008] Step 1: Dissolve deferoxamine in deionized water to prepare a mixed liquid, and store the mixed liquid at 4°C for use;

[0009] Step 2: adding an equal amount of lithium magnesium silicate nanosheets to the deferoxamine in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles, and washing the obtained nanoparticles with deionized water and drying them under nitrogen conditions;

[0010] Step 3: Pure PVA and deionized water were mixed in a flask and heated at 90 °C for 1 h. Pure PVA hydrogel was prepared using the mixed liquid by repeated freezing and thawing for four additional cycles.

[0011] Step 4: The polyvinyl alcohol aqueous solution, tannic acid, and the lithium magnesium silicate / deferoxamine nanoparticles obtained in step 2 are mixed in a flask, and the mixed liquid is used to prepare the polyvinyl alcohol / lithium magnesium silicate / deferoxamine nanohydrogel by repeated freezing and thawing for four additional cycles.

[0012] Preferably, in step 1, the concentration of the mixed liquid is 1 mg / mL.

[0013] Preferably, in step three, the amount of pure polyvinyl alcohol is 40 g and the amount of deionized water is 100 mL.

[0014] Preferably, in step 4, the amount of tannic acid is 0.1 mg, and the mass of lithium magnesium silicate / deferoxamine nanoparticles is 5%, 10%, or 15% of the mass of polyvinyl alcohol.

[0015] The advantages and positive effects of the method for preparing the biofunctionalized nanohydrogel for cranial bone regeneration described in the present invention are:

[0016] 1. The present invention utilizes a polyvinyl alcohol hydrogel system as a substrate for skull defect repair. The system not only strengthens damaged bone tissue, but also acts as a carrier to transport active ions and drugs in a lithium magnesium silicate nanocomposite material combined with deferoxamine through biocompatibility and adhesive tannic acid. Tannic acid can be loaded into polyvinyl alcohol through the electrostatic interaction of polyphenol groups and enhance the physical crosslinking of the hydrogel and the LAP / DFO nanocomposite material. At the same time, the hydrogel system has suitable stiffness, and due to the protective effect of the hydrogel matrix, the self-degradation of the LAP / DFO nanocomposite material will be significantly inhibited. A large amount of deferoxamine is loaded on the LAP surface through electrostatic interaction, thereby producing an obvious sustained-release effect, enhancing the biological activity of DFO, and meeting the long-term needs of cell response and bone remodeling. In the microenvironment of tissue injury, it can be observed that DFO and bioactive ions are sequentially released from the PLA hybrid hydrogel loaded with LAP / DFO, creating a pro-regenerative microenvironment.

[0017] 2. The present invention comprehensively analyzes the physicochemical structure and properties of biofunctionalized nanohydrogels, the release behavior of bioactive factors in vitro, and the cell responses including adhesion, proliferation, angiogenesis effects and osteogenic differentiation. In vivo studies have shown that when multi-component PVA hydrogels are implanted into skull defects, they promote rapid vascular invasion and accelerate bone integration.

[0018] The technical solution of the present invention is further described in detail below through the accompanying drawings and embodiments. BRIEF DESCRIPTION OF THE DRAWINGS

[0019] Figure 1 The figure is a flow chart of the preparation of the biofunctionalized nano hydrogel of the present invention;

[0020] Figure 2 Figures showing the morphology and composition characteristics of different hydrogels in the verification examples of the present invention, wherein A is a TEM image showing the dispersion of LAP and DFO / LAP in water, B is an EDS image showing the composition of LAP, C is an EDS image showing the composition of DFO / LAP, and D is the preparation and cross-linking process of the nanocomposite hydrogel construct; E is a photo of PVA, PL5, PL10 and PL15 hydrogels (left) and their freeze-dried forms (right) (scale bar: 1 cm); F is a representative FE-SEM image of PVA, PL5, PL10 and PL15 hydrogels;

[0021] Figure 3 The mechanical and physical properties characterization diagram of the PVA, PL5, PL10 and PL15 hydrogels of the present invention, wherein A is the tensile curve, B is the compressive stress-strain curve, C is the swelling rate, D is the compression modulus of various hydrogels, E is the cumulative release of DFO in saline solution, F is the cumulative release of Si ions, G is the cumulative release of Mg 2+ Cumulative release of ions, H is Li + Cumulative release of ions;

[0022] Figure 4 The figure is a characterization diagram of the cell adhesion ability of different hydrogels in the verification example of the present invention, wherein A is a three-dimensional confocal microscope image of cytoskeleton staining of human umbilical vein endothelial cells (HUVECs) on different hydrogels, B is a microscopic image of tubular structures formed by human umbilical vein endothelial cells (scale: 100 microns); C is the quantification result of the total microtubule area; D is the cell viability of human umbilical vein endothelial cells after culturing on different hydrogels for 1 day and 3 days;

[0023] Figure 5This is a validation example of the present invention, showing the effects of different hydrogels on the expression of human umbilical vein endothelial cells and bone marrow mesenchymal stem cells, wherein A is the effect of different hydrogels on the expression of angiogenesis-related genes (VEGFA, eNOS, ANG-1, VEGF and SMAD5 / 8) in human umbilical vein endothelial cells (HUVECs) after 3 days of culture, B is the relative mRNA expression of genes (ALP, OPG, COLI and OCN) related to osteogenesis in bone marrow mesenchymal stem cells (MSCs) after treatment with different samples, C is a qualitative analysis of the alkaline phosphatase (ALP) activity of MSCs cultured on the surfaces of various samples at 7 days and 14 days, D is the degree of mineralization of MSCs cultured on different samples for 14 days and 21 days, and E is the effect of human umbilical vein endothelial cell conditioned medium on MSCs mineralization;

[0024] Figure 6 1 is a validation example of the present invention for different hydrogels used for in vivo bone regeneration, wherein A is a schematic diagram of a rat skull defect model and hydrogel implantation, B is a reconstructed three-dimensional micro-CT image of a skull defect area treated with circle marks (yellow: blank, red: hydrogel), C is a quantitative analysis of bone defect regeneration by measuring BV / TV, D is a quantitative analysis of bone defect regeneration by measuring Tb.Th, E is a quantitative analysis of bone defect regeneration by measuring Tb.N, F is a quantitative analysis of bone defect regeneration by measuring Tb.Sp, and G is a quantitative analysis of bone defect regeneration by measuring BS / BV;

[0025] Figure 7 These are H&E and Masson trichrome staining images of rat skull defects after implantation of different PVA hydrogels for 1 month and 2 months, wherein A is an H&E image and B is a Masson image;

[0026] Figure 8 This is a graph showing the average pore size distribution results of the verification examples PL5, PL10 and PL15 of the present invention;

[0027] Fig. 9 The figure is a characterization diagram of the cell adhesion ability of different hydrogels in the verification example of the present invention, wherein A is a three-dimensional confocal microscope image of the cytoskeleton staining of mesenchymal stem cells on different hydrogels, and B is the cell viability of mesenchymal stem cells after culturing on different hydrogels for 1 day and 3 days;

[0028] Fig.10 These are the immunohistochemical staining images of COLⅠ, OCN and VEGF after implantation of different PVA hydrogels for 1 month and 2 months in the validation example of the present invention. DETAILED DESCRIPTION

[0029] The technical solution of the present invention is further described below through the accompanying drawings and embodiments.

[0030] Unless otherwise defined, the technical terms or scientific terms used in the present invention shall have the common meanings understood by persons with ordinary skills in the field to which the present invention belongs. The reagents and equipment involved in the present invention are all commercially available unless otherwise specified.

[0031] Example 1

[0032] like Figure 1 As shown, the preparation method of biofunctionalized nanohydrogel for skull regeneration comprises the following steps:

[0033] Step 1: Dissolve deferoxamine (DFO) in deionized water to prepare a mixed liquid with a concentration of 1 mg / mL, and store the mixed liquid at 4°C for use;

[0034] Step 2: adding an equal amount of lithium magnesium silicate nanosheets (LPA) to the deferoxamine (DFO) in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO), and the obtained nanoparticles are washed with deionized water and dried under nitrogen conditions;

[0035] Step 3: 40 g of pure polyvinyl alcohol (PVA) and 100 mL of deionized water were mixed in a flask and heated at 90°C for 1 hour. The mixed liquid was used to prepare pure polyvinyl alcohol hydrogel by repeated freezing and thawing for four additional cycles. The obtained polyvinyl alcohol hydrogel was recorded as PVA.

[0036] Example 2

[0037] A method for preparing a biofunctionalized nanohydrogel for cranial bone regeneration comprises the following steps:

[0038] Step 1: Dissolve deferoxamine (DFO) in deionized water to prepare a mixed liquid with a concentration of 1 mg / mL, and store the mixed liquid at 4°C for use;

[0039] Step 2: adding an equal amount of lithium magnesium silicate nanosheets (LPA) to the deferoxamine (DFO) in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO), and the obtained nanoparticles are washed with deionized water and dried under nitrogen conditions;

[0040] Step 3: 10 mL of polyvinyl alcohol (PVA) aqueous solution (0.4 g / mL), 0.1 mg of tannic acid (TA), and 0.2 g of lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO) were mixed in a flask, and the mixed liquid was used to prepare polyvinyl alcohol / lithium magnesium silicate / deferoxamine nanohydrogel (PVA / LPA / DFO) by repeated freezing and thawing for four additional cycles. The resulting hydrogel was recorded as PL5.

[0041] Example 3

[0042] A method for preparing a biofunctionalized nanohydrogel for cranial bone regeneration comprises the following steps:

[0043] Step 1: Dissolve deferoxamine (DFO) in deionized water to prepare a mixed liquid with a concentration of 1 mg / mL, and store the mixed liquid at 4°C for use;

[0044] Step 2: adding an equal amount of lithium magnesium silicate nanosheets (LPA) to the deferoxamine (DFO) in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO), and the obtained nanoparticles are washed with deionized water and dried under nitrogen conditions;

[0045] Step 3: 10 mL of polyvinyl alcohol (PVA) aqueous solution (0.4 g / mL), 0.1 mg of tannic acid (TA), and 0.4 g of lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO) were mixed in a flask and heated at 90°C for 1 hour. The mixed liquid was used to prepare polyvinyl alcohol / lithium magnesium silicate / deferoxamine nanohydrogel (PVA / LPA / DFO) by repeated freezing and thawing for four additional cycles. The resulting hydrogel was recorded as PL10.

[0046] Example 4

[0047] A method for preparing a biofunctionalized nanohydrogel for cranial bone regeneration comprises the following steps:

[0048] Step 1: Dissolve deferoxamine (DFO) in deionized water to prepare a mixed liquid with a concentration of 1 mg / mL, and store the mixed liquid at 4°C for use;

[0049] Step 2: adding an equal amount of lithium magnesium silicate nanosheets (LPA) to the deferoxamine (DFO) in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO), and the obtained nanoparticles are washed with deionized water and dried under nitrogen conditions;

[0050] Step 3: 10 mL of polyvinyl alcohol (PVA) aqueous solution (0.4 g / mL), 0.1 mg of tannic acid (TA), and 0.6 g of lithium magnesium silicate / deferoxamine nanoparticles (LPA / DFO) were mixed in a flask and heated at 90°C for 1 hour. The mixed liquid was used to prepare polyvinyl alcohol / lithium magnesium silicate / deferoxamine nanohydrogel (PVA / LPA / DFO) by repeated freezing and thawing for four additional cycles. The resulting hydrogel was recorded as PL15.

[0051] Verification Example

[0052] The performance of the hydrogels PVA, PL5, PL10, and PL15 prepared in Examples 1-4 was tested by the following method.

[0053] Sample characterization: The morphology of LAP and LAP / DFO nanocomposites was observed by transmission electron microscopy (TEM), the macroscopic morphology of different hydrogels was observed by a conventional camera, and the macroscopic morphology of freeze-dried samples was observed by scanning electron microscopy (SEM).

[0054] Swelling rate measurement: All samples were made into cylinders with a diameter (d): 1 cm and a height (h): 2 mm, and then freeze-dried to measure the dry weight. Figure 2 E. The lyophilized samples were then immersed in PBS at 25°C for 1 day, and the weight of the fully swollen samples was checked. The swelling ratio was calculated as the wet weight of the fully swollen hydrogel divided by the dry weight of the lyophilized hydrogel.

[0055] Mechanical properties test: For compression and tension tests, a universal mechanical testing system was used, and the tests were carried out at a deformation rate of 1 cm / min at 25 °C. All samples were synthesized in a box (length l: 1 cm, width w: 3 mm, height h: 2 mm) and a cylinder (diameter d: 1 cm, height h: 2 mm), and the Young's modulus was calculated from the linear region of the stress-strain curve.

[0056] Ion release assay: To determine the ion concentration, the hydrogels in each group were incubated in 5 mL of saline (pH 7.4). After soaking at 37°C for different times (0, 1, 3, 7, 11, 14, 21, 3, and 60 days), all the solutions were collected and new saline (5 mL) was added. After centrifugation, the collected liquid was tested for all bioactive ions (Li, Si, and Mg) using inductively coupled plasma atomic emission spectrometry (ICP-AES). In addition, after DFO was combined with a specific amount of ferric chloride, the amount of DFO released was measured by measuring the absorbance at a wavelength of 485 nm at fixed intervals using a UV-visible spectrophotometer.

[0057] Cell culture: Mesenchymal stem cells (MSCs) were isolated from SD rats and cultured in a low-glucose medium, with 10% fetal bovine serum (FBS) and 1% penicillin-streptomycin added and cultured at 37°C in a 5% carbon dioxide (CO2) environment. The culture medium was replaced every 48 hours. When the cell concentration in the cell culture plate was close to 90%, 0.25% trypsin was used to passage and purify the cells. The third generation MSCs were used in the following experiments of the present invention. Similarly, human umbilical vein endothelial cells (HUVECs) were inoculated in a high-glucose medium and the cell culture medium was replaced every day. When HUVECs reached about 90% in the culture flask, 0.25% trypsin was used for passage and purification.

[0058] Cell viability and adhesion: HUVECs or MSCs were incubated on different substrates for 3-7 days and -3 days, respectively. Subsequently, 500 μL of fresh culture medium (DMEM) containing 3-(4,5-dimethylthiazole-2)-2,5-diphenyltetrazolium bromide (MTT, 5 mg / mL) was added to each well. After 240 minutes, the formed precipitate was dissolved in dimethyl sulfoxide (DMSO) and the mixture was transferred to a 96-well plate. The absorbance of the solution was tested at a wavelength of 490 nm using a microplate system.

[0059] In vitro vascularization experiment of human umbilical vein endothelial cells (HUVECs): In vitro vascularization experiment was performed using extracellular matrix gel (ECMatrix TM Specifically, 50 μL of extracellular matrix gel (ECMatrix) was added to a pre-cooled 96-well plate. HUVECs (1.5×10 4 After 6 and 12 h of incubation, HUVECs were observed using an inverted microscope (NIB620) at each time point, and the total area of ​​the grid (tube) was quantified using ImageJ software.

[0060] Gene expression of HUVECs and MSCs: HUVECs were co-cultured with different hydrogels for 72 h, and the messenger RNA (mRNA) expression levels of target genes, including angiopoietin-1 (ANG-1), endothelial nitric oxide synthase (eNOS), vascular endothelial growth factor A (VEGFA), vascular endothelial growth factor (VEGF), and fibroblast growth factor (SMAD / 5 / 8), were quantified by reverse transcription quantitative polymerase chain reaction (RT-qPCR). In addition, MSCs were incubated with various samples for 14 days. The expression levels of genes related to osteogenesis, namely alkaline phosphatase (ALP), type I collagen (COLⅠ), osteocalcin (OCN), and osteoprotegerin (OPG), were determined using RT-qPCR.

[0061] In vitro osteogenic differentiation of MSCs: To determine ALP activity, cells were incubated on the four hydrogels for 7 and 14 days. Subsequently, all matrices were washed with PBS and treated with 1% tritonX-100 buffer for 0.5 h. The supernatant was then transferred to a 96-well plate and the ALP activity was assessed using an ALP kit. In addition, protein content was measured using a BCA protein assay kit and absorbance was measured at 490 nm and 570 nm. For mineralization, which indicates late osteogenic differentiation, an Alizarin Red S (ARS) staining experiment was performed. After 14 and 21 days of culture, MSCs incubated with various hydrogels were treated with 4% paraformaldehyde for 10 min and then combined with ARS (40 mM) for 30 min. To quantitatively measure matrix mineralization, all samples were soaked with 10% acetic acid for another 1 h and then the absorbance was measured at 405 nm after adding 10% equal volume of ammonia water and mixing.

[0062] Establishment of skull defect model and implantation surgery: A skull defect model of critical size was established in male SD rats (120 g). Two circular defects with a diameter of 0.5 cm were made on both sides of the skull of the SD rats using a circular drill. Then, 24 male SD rats were randomly divided into four groups, 6 in each group. For each rat, the defect on the left was not filled (control group), while the defect on the right was filled with different hydrogels. After the implantation surgery, the surgical site was sutured.

[0063] In vivo angiogenesis and osteogenesis assessment: At 1 month and 2 months after implantation, the skulls of all SD rats were treated with 4% paraformaldehyde for 72 hours. The purpose of this fixation was to facilitate further investigation of bone regeneration using Micro-CT scanning. After scanning, the acquired data were subjected to 3D reconstruction and quantitative analysis. In addition to Micro-CT analysis, hematoxylin and eosin (H&E) and Masson's trichrome staining experiments were also performed. The skulls were collected at 1 month and 2 months after implantation and fixed for 72 hours. After immersion in 12wt% ethylenediaminetetraacetic acid solution for 21 days, the decalcified skulls were prepared for subsequent staining. All samples were observed by confocal laser scanning microscopy.

[0064] For immunohistochemical staining experiments, all samples were fixed and decalcified for 2 weeks as described above. Dewaxed sections were soaked with primary antibodies (COL1, OCN, VEGF), and all samples were observed by fluorescence microscopy.

[0065] (1) Synthesis and characterization of LAP / DFO nanocomposites.

[0066] In order to observe the morphology and structure of LAP and LAP / DFO, transmission electron microscopy (TEM) was used. Figure 2As shown in Figure A, the magnified LAP shows a disk-shaped geometric structure with a diameter of 20-50 nm and a thickness of 1-2 nm. In addition, at higher magnifications, the LAP / DFO nanocomposite exhibits a nanoscale cluster structure (e.g. Figure 2 This is probably caused by the electrostatic interaction between the positively charged DFO and the negatively charged LAP surface. In addition, energy dispersive spectroscopy (EDS) analysis was used to determine the elemental distribution and composition of LAP and LAP / DFO. Figure 2 Medium B and Figure 2 As shown in Figure 3C, both samples contain silicon (Si), oxygen (O), and magnesium (Mg). It is noteworthy that nitrogen (N) is only observed in the LAP / DFO nanocomposite. It is noteworthy that N is the main component of DFO, which means that DFO has been successfully bonded to the LAP surface.

[0067] (2) Preparation and characterization of various hydrogels.

[0068] To further investigate the structure and pore size of various hydrogels, scanning electron microscopy (SEM) images were used. Figure 2 Medium D and Figure 2 As shown in Figure E, except for pure PVA, PL5, PL10, and PL15 all showed a porous interpenetrating structure. Then, based on the SEM images, the average pore size of the freeze-dried hydrogel group was quantitatively analyzed. In particular, the average pore sizes of PL5, PL10, and PL15 were 16.25±3.50, 40.50±3.11, and 79.25±2.99 μm, respectively (Figure 5E). Figure 8 As the LAP / DFO content increases, the average pore size of the hydrogel gradually increases, as shown in Figure 2 As shown in F.

[0069] (3) Mechanical and physical properties of hydrogels.

[0070] To evaluate the mechanical properties of the hydrogels, various tests including compression tests, tensile measurements, and swelling tests were performed. The results showed that from PVA to PL15, as the LAP / DFO content increased, both the compressive and tensile strengths decreased (e.g. Figure 3 Middle A and Figure 3 In contrast, the expansion rate shows the opposite trend (as shown in Figure 3 Specifically, the compression modulus values ​​of the PVA, PL5, PL10, and PL15 groups were 84.9±14.6, 42.1±5.07, 18.2±1.67, and 3.4±0.12 kPa, respectively (as shown in C). Figure 3 These results indicate that the compressive strength and modulus of all experimental groups are lower than those of pure PVA.

[0071] (4) Ion release behavior.

[0072] In order to obtain the release curves of bioactive factors from hydrogels, all hydrogel samples were immersed in physiological saline (pH 7.4) for 2 months. Figure 3 Middle E, Figure 3 Medium F, Figure 3 Medium G, Figure 3 As shown in Figure 5H, DFO, Mg, Si, and Li ions were continuously released from various hydrogels until the 60th day. It is worth noting that the release rate of DFO was relatively fast and higher than that of Mg, Si, and Li ions during the first 2 weeks. This result was attributed to the preferential release of DFO attached to the surface of LAP nanoplates. In addition, after the 14th day, the release of Mg, Si, and Li ions showed an accelerating growth trend. During the first 60 days, the specific DFO concentrations measured for PL5, PL10, and PL15 were 94.1, 112.1, and 162.8 mg / L, respectively. In addition, the Mg release rate of PL5, PL10, and PL15 groups increased significantly within 60 days. 2+ (25.7, 32.6 and 40.8 mg / mL), Li + The cumulative release of 2-(0.7, 1.1, and 1.4 mg / mL) and Si ions (32.3, 35.2, and 42.5 mg / mL) was calculated, respectively. The data showed that the total release of the PL15 group was much higher than that of the other groups, which may be due to the large amount of LAP / DFO nanocomposites and the macroporous interpenetrating microstructure. In addition, all the results demonstrated the successful release of bioactive factors from the polymer system.

[0073] (5) In vitro biocompatibility analysis of human umbilical vein endothelial cells (HUVECs) and mesenchymal stem cells (MSCs).

[0074] The cell adhesion ability of different hydrogels was evaluated by cytoskeleton staining. Figure 4 A in the figure shows the corresponding green fluorescence photo of HUVECs taken using a confocal laser scanning microscope (CLSM). FITC-Phalloidin can specifically bind to the F-actin of HUVECs, thereby revealing the distribution of the microfilament cytoskeleton. Therefore, HUVECs adhered to the hydrogel showed green fluorescence under CLSM. Compared with pure PVA hydrogel, the PL5, PL10 and PL15 groups all showed a large number of green fluorescent spots on the 1st and 3rd days, which means that the modified hydrogel system is effective in improving biocompatibility and cell adhesion. Figure 4As shown in Figure A, more cells adhered to the surface of PL10 compared with other experimental groups on day 1. In addition, due to the unique microporous structure and appropriate matrix strength of PL10, it showed more cell proliferation in the porous interpenetrating microstructure of the hydrogel system than other groups after 3 days (as shown in Figure 2A). Figure 4 The number of adherent cells on the PL15 surface decreased after 3 days, indicating poor biocompatibility. This trend was also observed in MSCs treated with the same method (see Figure 2A). Fig. 9 In addition, the survival rate of HUVECs incubated on different samples was evaluated by MTT assay. Figure 4 As shown in Figure D, PL5 and PL10 showed significant differences in cell survival after 3 days of culture compared with the PVA and PL15 groups (p<0.01). This finding strongly suggests that PL10 is more beneficial in promoting cell proliferation. A similar trend was observed after MSCs were treated with the same conditions for 7 days (Figure Fig. 9 (shown in B).

[0075] (6) In vitro angiogenesis and osteogenesis assays.

[0076] Bone regeneration is a complex and dynamic process that requires the collaboration of multiple cells from different systems. To evaluate the level of vascularization, HUVECs were cultured on matrix gels containing conditioned medium from different hydrogels. Conditioned medium from PL5, PL10, and PL15 significantly promoted and accelerated the formation of capillary-like structures in the gels (e.g., Figure 4 The conditioned medium from PL10 was able to promote the angiogenic activity of HUVECs, and capillary networks were still visible after 12 hours. This indicates that the appropriate concentration of DFO released into the conditioned medium is beneficial to accelerate the formation of capillary-like structures. Further quantitative analysis of the total area of ​​the mesh is shown in Figure 2. Figure 4 As shown in C. In addition, compared with the PL5 and PL15 groups, especially in the PL10 group on day 3, the gene expressions of eNOS, ANG-1, Smad5 / 8VEGF, and VEGFA were upregulated in HUVECs cultured in the PL10 group (as shown in Figure 5 These findings indicate that PL10 significantly accelerated tube formation in HUVECs compared with the other groups.

[0077] like Figure 5As shown in Figure 2B, the mRNA expression of ALP, COLⅠ, OCN, and OPG in MSCs was further measured. Compared with the pure PVA group, PL5, PL15, and especially PL10 were found to have a profound effect on the expression of osteogenesis-related genes in MSCs. In addition, to evaluate the osteogenic differentiation of MSCs cultured on different hydrogels, ALP activity was measured. Figure 5 As shown in Figure C, the PL10 group showed significantly higher ALP activity compared with the other groups at both the 1st and 2nd week (p<0.01). After 2 and 3 weeks of culture, the mineralization results showed similar trends (e.g. Figure 5 (D). These combined findings highlight that the PL10 group is better than the other groups in terms of osteoinductive potential. PL10 mimics precalcified bone, has appropriate mechanical strength and release of bioactive factors, tending to stimulate osteogenic differentiation of MSCs. To explore the correlation between angiogenesis and osteogenesis, MSCs were incubated with conditioned medium from HUVECs cultured on different samples. Figure 5 As shown in E, the results were detected after 14 and 21 days of culture (n=3, *p<0.05 and **p<0.01). In the PL10 group, the mineralization level of MSCs was significantly upregulated after culture with HUVECs-mediated conditioned medium.

[0078] (7) In vivo assessment of bone formation.

[0079] In order to explore the feasibility of different hydrogels in clinical bone regeneration, a skull defect model of SD rats (such as Figure 6 Two holes with a diameter of 5 mm were made on both sides of each parietal bone, and four types of hydrogels were implanted in the holes on the right. Micro-CT imaging and 3D reconstruction were performed 1 month and 2 months after implantation. Figure 6 As shown in Figure 2B. At 1 month after implantation, limited new bone formation was observed in all groups. At 2 months after implantation, each group showed more substantial new bone regeneration. Notably, a large amount of newly formed bone was found around the parietal bone defect in the rats in the PL10 group compared with the other three groups. This observation was further supported by the bone volume / total volume (BV / TV, as shown in Figure 2B). Figure 6 The quantitative analysis of the PL10 group was supported by the results of the trabecular thickness analysis (shown in Figure 5C), which showed that there was a significant difference between the PL10 group and the other groups (p < 0.05). Figure 6 D), the number of trabeculae (Tb.N, as shown in Figure 6 E) showed a consistent increasing trend, while the trabecular space (Tb.Sp, as Figure 6 Figure 5(a) and bone surface area / bone volume (BS / BV, as shown in Figure 5(b)). Figure 6In summary, consistent with the in vitro findings, the in vivo data also confirmed that the osteogenic capacity of PL10 was significantly higher than that of the other groups.

[0080] like Figure 7 As shown in A, the histological evaluation results displayed clearly show that the red arrows indicate the location of the new blood vessels (N: new bone tissue; P: non-degraded PVA hydrogel; V: new blood vessels; F: fibrous tissue; O: old bone boundary; scale bars: 1 mm and 2 mm). At 2 months after implantation, only fibrous tissue existed in the defect area. When pure PVA was implanted for 1 and 2 months, only fibrous tissue and non-degraded PVA were left in the defect area. In the PL5 and PL15 groups, only a small amount of newly formed bone coexisted with connective tissue. However, in the PL10 group, obvious newly formed bone and blood vessels (red arrows) along the fibrous membrane restoring the defect could be observed 1 month after implantation, and further extended to the central area of ​​the defect. 2 months after implantation, the newly formed bone tissue was more mature and thicker. Masson staining (MT) experiments were further performed to distinguish between bone and fibrous tissue (as shown in Figure 2). Figure 7 Figure 3 (B). In the PVA, PL5, and PL15 groups, the defect sites were filled with loose fibrous tissue and limited new bone at 1 and 2 months after implantation. In contrast, in the PL10 group, new bone and immature collagen (blue areas) dominated at 1 month after implantation. In the following month, bone tissue matured, as shown by robust mature collagen (red areas) and blood vessels (red arrows) that almost completely filled the defects, demonstrating impressive osteogenic capacity. These findings are consistent with the micro-CT data and provide a comprehensive understanding of the performance of hydrogels in promoting bone regeneration.

[0081] To further reveal the osteogenic ability of various hydrogels in vivo, immunohistochemical staining of COLⅠ, OCN, and VEGF was performed 1 month and 2 months after implantation (e.g. Fig.10 Compared with the other three groups, the positive expression of COLⅠ and OCN in the PL10 group was significant, especially OCN (as shown in Fig.10 In addition, higher angiogenesis-related VEGF expression in PL10 suggested that proper vascularization could help accelerate skull bone regeneration.

[0082] A multifunctional bioactive PVA hydrogel of the present invention promotes vascularization and cranial bone regeneration by regulating the release of bioactive factors and simulating pre-calcified bone. The uniqueness of the hydrogel lies in the mixed cross-linking bonding between LAP / DFO and the carbohydrate-based hydrogel network. This bonding provides a new and effective method for improving the microstructural stability and cell adhesion ability of the PVA hydrogel network, while providing significantly higher porosity and interconnected pore structures. These properties contribute to sustained bioactive factor release, cell growth and vascularized bone formation.

[0083] Porous structure is a key factor in designing hydrogels for bone tissue repair, as interconnected pores are essential for promoting cell adhesion, proliferation, migration, and nutrient delivery. Based on SEM images, the present invention observed that the average pore size of the hydrogel gradually increased as the LAP / DFO content increased from PL5 to PL15. Specifically, the average pore sizes of PL5, PL10, and PL15 were 16.25±3.50, 40.50±3.11, and 79.25±2.99 μm, respectively (Figure 2). Figure 8 As shown). According to the experimental results of cell adhesion, the pore size affects the initial cell attachment, and cell attachment tends to increase with the increase of pore size in the lower range (20-120μm). The cell adhesion level in PL10 is higher than that in PVA and PL5 because of the larger pore size of PL10. Although PL15 has a larger pore size than PL10, its cell adhesion level is lower. This trend is also reflected in the experimental results of cell viability. In addition, the mechanical results show that the compressive strength and modulus of the hydrogel decrease as the pore size increases from PL5 to PL15.

[0084] Mesenchymal stem cells (MSCs) cultured on a rigid substrate simulating precalcified bone (20-40 kPa) tend to differentiate into osteoblasts. Figure 3 Middle A and Figure 3 The results in Figure B show that the mechanical properties of PL5 and PL10 fall within this range, which is crucial for designing and preparing hydrogels that meet the mechanical requirements for repairing bone tissue.

[0085] To relieve intracranial pressure, patients may delay cranioplasty for months after decompressive craniectomy. This invention highlights the clinical application potential of PVA hydrogels, especially PL10, whose adhesiveness and flexibility can not only maintain the integrity of the hydrogel, but also adapt well to intracranial pressure fluctuations. Therefore, it will allow instant repair of skull defects in a one-stop surgery of decompressive craniectomy and cranioplasty.

[0086] In addition, the composite hydrogels exhibited excellent porosity and interconnected pore structures, which were beneficial for the sustained release of bioactive factors. Bone regeneration is a complex and dynamic process involving the collaboration of multiple cells from different systems. Early vascularization determines the outcome of osteogenesis, showing the indispensable role of HUVECs (human umbilical vein endothelial cells) in the bone regeneration process. The sustained release of Si, Mg, and Li ions from LAPs can promote bone regeneration by stimulating the osteogenic differentiation of MSCs. This sustainable and stable ion release brings an additional benefit of preventing ectopic bone that may result from the sudden release of bioactive ions. Therefore, the combination of DFO with Si, Mg, and Li ions may provide an important incentive for the spatiotemporal manipulation of vascular development and bone formation in skull defects. In this way, the sequential release of two bioactive drugs from the hydrogel was achieved, designed for the early release of DFO and the late release of bioactive ions. As a result, the release process of DFO was extended to more than 14 days ( Figure 3 Middle E to Figure 3 H). The data showed that appropriate concentrations of DFO from PL10 were released into the microenvironment to modulate the function of HUVECs. This modulation promoted the development of angiogenesis by activating subsequent signaling pathways and / or transcriptional mediators. In addition, LAP exhibited a controlled release of Mg, Li, and Si ions for more than 30 days. This sustained release directly promoted the osteogenic differentiation of MSCs, further enhancing bone regeneration (e.g. Figure 5 This may be due to the porous structure and interconnected porosity of the gel, which makes the bioactive factors easily encapsulated in the hydrogel network and gradually decomposed from the hydrogel.

[0087] The composite hydrogels, especially PL10, have excellent porosity, pore size, and interconnected pore structure, which is conducive to the sustained release of bioactive factors into the microenvironment. This design, in turn, leads to rapid angiogenesis and stimulates bone regeneration. Figure 5 As shown in Figure E, the mineralization level of MSCs on PL10 was significantly upregulated after 21 days. The present invention confirms that HUVECs express some biological factors that work together with osteogenic growth factors to enhance the osteogenic differentiation of MSCs. PL10 not only directly promotes the functions of HUVECs and MSCs, but also synergistically regulates the interaction between them, leading to accelerated neovascularization and new bone formation. The bone repair effects of various multifunctional PVA hydrogels were evaluated in a rat skull defect model. The results showed that PL10 was most effective in inducing neovascularization and new bone formation after 2 months, indicating its potential clinical importance.

[0088] The effects of multi-component PVA hydrogels on the behavior of HUVECs and MSCs were investigated with the aim of accelerating skull regeneration. The results confirmed that PL10, with appropriate mechanical strength and optimal release of bioactive factors, significantly promoted the proliferation of HUVECs, promoted lumen formation for angiogenesis, and induced osteogenic differentiation of MSCs. Furthermore, in vivo studies showed that PL10 was superior to blank PVA in enhancing bone regeneration, confirming its potential role in promoting new bone formation. These findings suggest that PL10 has good prospects for convenient application in repairing skull defects.

[0089] Therefore, the present invention adopts the above-mentioned preparation method of biofunctionalized nanohydrogel for skull regeneration to prepare a multifunctional bioactive PVA hydrogel, which promotes vascularization and skull regeneration by regulating the release of bioactive factors and simulating pre-calcified bone. The uniqueness of this hydrogel lies in the mixed cross-linking bonding between LAP / DFO and carbohydrate-based hydrogel networks, which provides a new and effective method for improving the microstructural stability and cell adhesion ability of the PVA hydrogel network, while providing significantly higher porosity and interconnected pore structure. These properties contribute to the sustained release of bioactive factors, cell growth and vascularized bone formation.

[0090] Finally, it should be noted that the above embodiments are only used to illustrate the technical solution of the present invention rather than to limit it. Although the present invention has been described in detail with reference to the preferred embodiments, those skilled in the art should understand that they can still modify or replace the technical solution of the present invention with equivalents, and these modifications or equivalent replacements cannot cause the modified technical solution to deviate from the spirit and scope of the technical solution of the present invention.

Claims

1. A method for preparing a biofunctionalized nanohydrogel for skull regeneration, characterized in that: The steps include: Step 1: Dissolve deferoxamine in deionized water to prepare a mixed liquid, and store the mixed liquid at 4°C for use; Step 2: adding an equal amount of lithium magnesium silicate nanosheets to the deferoxamine in step 1 to the mixed liquid, and continuing stirring for 120 minutes to form modified lithium magnesium silicate / deferoxamine nanoparticles, and washing the obtained nanoparticles with deionized water and drying them under nitrogen conditions; Step 3: Pure PVA and deionized water were mixed in a flask and heated at 90 °C for 1 h. Pure PVA hydrogel was prepared using the mixed liquid by repeated freezing and thawing for four additional cycles. Step 4: The polyvinyl alcohol aqueous solution, tannic acid, and the lithium magnesium silicate / deferoxamine nanoparticles obtained in step 2 are mixed in a flask, and the mixed liquid is used to prepare the polyvinyl alcohol / lithium magnesium silicate / deferoxamine nanohydrogel by repeated freezing and thawing for four additional cycles.

2. The method for preparing the biofunctionalized nanohydrogel for cranial bone regeneration according to claim 1, characterized in that: In step 1, the concentration of the mixed liquid is 1 mg / mL.

3. The method for preparing the biofunctionalized nanohydrogel for cranial bone regeneration according to claim 1, characterized in that: In step 3, pure polyvinyl alcohol is 40 g and deionized water is 100 mL.

4. The method for preparing the biofunctionalized nanohydrogel for cranial bone regeneration according to claim 1, characterized in that: In step 4, the amount of tannic acid is 0.1 mg, and the mass of lithium magnesium silicate / deferoxamine nanoparticles accounts for 5%, 10%, and 15% of the mass of polyvinyl alcohol.

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