A double-layer hydrogel scaffold and a preparation method and application thereof

Abstract

This invention discloses a bilayer hydrogel scaffold, its preparation method, and its applications, relating to the field of medical repair materials technology. The bilayer hydrogel scaffold comprises upper and lower hydrogel layers made of materials with different degradation rates and mechanical properties, both layers being loaded with nano-fat emulsions. The thickness of the upper hydrogel layer is greater than or equal to the thickness of the lower hydrogel layer. This invention utilizes the differences in degradation rates and mechanical properties of different materials to prepare a bilayer hydrogel scaffold, integrating functions such as promoting nerve repair, promoting vascular activity, and promoting bone activity. It achieves the effect of simultaneously adapting to two different pathological environments within the same implant, thereby avoiding secondary surgery for patients. This invention utilizes photocrosslinking technology to achieve integrated molding of the bilayer structure, ensuring a strong interface bond between the upper and lower layers and eliminating the risk of delamination. Simultaneously, the preparation process is controllable and highly customizable, allowing for flexible adjustment of the scaffold's size and pore structure according to the clinical defect morphology.

Description

Technical Field

[0001] This invention relates to the field of medical repair materials technology, and in particular to a double-layer hydrogel scaffold, its preparation method and application. Background Technology

[0002] Traumatic brain injury (TBI) has become the leading cause of death and disability worldwide due to trauma, and its incidence rate ranks first among various neurological diseases. The pathological process of TBI includes two types of damage: primary and secondary. Primary damage is caused directly by external mechanical forces (such as direct impact or acceleration / deceleration), manifesting as focal contusions, hematomas, and other structural damage. Secondary damage is a series of molecular damage mechanisms triggered after trauma, involving excitotoxicity, electrolyte imbalance, and inflammatory responses, which further aggravate brain injury. Although more than 90% of TBI patients are classified as "mild" according to the Glasgow Coma Scale, more than half of them still fail to fully recover within six months after injury, suggesting that even functional recovery in mild TBI has significant clinical implications.

[0003] To address the increased intracranial pressure and cerebral edema caused by total brain injury (TBI), craniotomy for decompression is commonly performed clinically. However, this procedure often results in skull defects. Currently, cranioplasty is an effective complementary method for repairing such defects, typically performed 3 to 6 months after decompression surgery. However, this staged surgical approach not only prolongs the treatment period but also increases patient suffering and the financial burden on healthcare providers.

[0004] Therefore, there is an urgent need for a repair material that can be implanted simultaneously during craniotomy decompression surgery to meet the needs of both brain tissue repair and skull defect repair after decompression, thereby avoiding secondary surgery. However, there are significant differences in the repair time requirements, mechanical environment, and bioactivity requirements between brain tissue and skull after TBI craniotomy decompression: brain tissue has low mechanical strength, requiring materials to quickly exert anti-inflammatory, reactive oxygen species scavenging, and blood-brain barrier function to reduce secondary damage; while skull defect repair has a longer cycle and higher mechanical requirements, requiring materials to continuously provide anti-inflammatory, angiogenesis-promoting, and bone-promoting effects. This presents differentiated requirements for the time of stratified degradation and release of active ingredients in implanted repair materials, mechanical environment adaptation, and their respective bioactivity. Currently, there are no active implantable materials in clinical practice or existing literature that can simultaneously meet the above-mentioned differentiated requirements. Summary of the Invention

[0005] The technical problem to be solved by the present invention is to provide a repair material that can be implanted simultaneously during craniotomy decompression surgery, and can simultaneously meet the needs of brain tissue repair and skull defect repair after decompression, thus avoiding secondary surgery.

[0006] To address the above problems, the present invention proposes the following technical solution: In a first aspect, the present invention provides a double-layer hydrogel scaffold, comprising: The upper hydrogel layer is made of a first photocrosslinked hydrogel and is loaded with nano-fat emulsion. The lower hydrogel layer is made of a second photocrosslinked hydrogel and is loaded with nano-fat emulsion. The first photocrosslinked hydrogel and the second photocrosslinked hydrogel are different materials, and the degradation rate of the upper hydrogel layer is slower than that of the lower hydrogel layer, while the mechanical strength of the upper hydrogel layer is higher than that of the lower hydrogel layer.

[0007] Furthermore, the first photocrosslinked hydrogel is a composite hydrogel of methacrylamide chitosan and methacrylamide gelatin, and the second photocrosslinked hydrogel is a methacrylamide gelatin hydrogel.

[0008] Furthermore, the concentration of the nanofat emulsion is 0.1–10 mg / mL.

[0009] Furthermore, the thickness of the upper hydrogel layer is greater than or equal to the thickness of the lower hydrogel layer, and the thickness of the upper hydrogel layer is 0.1–10 mm, while the thickness of the lower hydrogel layer is 0.1–5 mm.

[0010] Furthermore, the bilayer hydrogel scaffold has a three-dimensional porous structure.

[0011] In a second aspect, the present invention provides a method for preparing the aforementioned bilayer hydrogel scaffold, comprising the following steps: (1) Prepare upper printing ink and lower printing ink, wherein the upper printing ink comprises a first photocrosslinking hydrogel precursor, a nano-fat emulsion and a photoinitiator, and the lower printing ink comprises a second photocrosslinking hydrogel precursor, a nano-fat emulsion and a photoinitiator; (2) Using 3D printing technology, the lower layer printing ink and the upper layer printing ink are printed layer by layer according to the preset three-dimensional model, and photocrosslinking and curing are performed on each printed layer; (3) Obtain a double-layer hydrogel scaffold in which the upper hydrogel layer and the lower hydrogel layer are fused together by photocrosslinking.

[0012] Furthermore, the first photocrosslinking hydrogel precursor is a mixture of methacrylamide chitosan and methacrylamide gelatin, and the second photocrosslinking hydrogel precursor is methacrylamide gelatin.

[0013] Further, the degree of substitution of the methacrylamide gelatin is 30-90%, and the concentration is 1-30 w / v; the degree of substitution of the methacrylamide chitosan is 30-90%, and the concentration is 1-25 w / v; the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonite, and the concentration is 0.1-1 w / v.

[0014] Furthermore, the conditions for photocrosslinking are an ultraviolet light wavelength of 365 nm and an irradiation time of 10–60 seconds.

[0015] Furthermore, the 3D printing process parameters are as follows: printing speed 5–20 mm / s, extrusion air pressure 1.5–3.0 bar, needle diameter 300–800 μm, row spacing 300–800 μm, and layer spacing 100–300 μm.

[0016] The present invention also provides the application of the aforementioned double-layer hydrogel scaffold or the double-layer hydrogel scaffold prepared by the method in the preparation of repair materials for treating brain tissue damage and skull defects after craniotomy decompression for traumatic brain injury.

[0017] Compared with the prior art, the technical effects achieved by the present invention include: The bilayer hydrogel scaffold provided by this invention employs an upper hydrogel layer and a lower hydrogel layer composed of different materials, both layers being loaded with nano-fat emulsions. The upper hydrogel layer degrades more slowly and has higher mechanical strength than the lower hydrogel layer. Utilizing this differential design of degradation rate and mechanical properties, the lower layer rapidly degrades and releases the nano-fat emulsion, quickly acting on damaged brain tissue to inhibit neuroinflammation and reduce secondary damage, while its softness prevents secondary injury to the brain tissue. The upper layer, on the other hand, slowly degrades and continuously releases the nano-fat emulsion, maintaining a microenvironment conducive to angiogenesis and bone regeneration over a long period, and providing sufficient mechanical support to meet the repair needs of skull defects. Therefore, this invention solves the mismatch in time, mechanical, and activity requirements between brain tissue repair and skull defect repair after craniotomy decompression, achieving the effect of simultaneously adapting to two different pathological environments within the same implant, thereby avoiding secondary surgery for patients.

[0018] The present invention provides a method for fabricating a bilayer hydrogel scaffold, which involves printing an upper and lower hydrogel layer using 3D printing technology. The two layers are then fused together via photocrosslinking to obtain the bilayer hydrogel scaffold. This invention utilizes photocrosslinking technology to achieve integrated molding of the bilayer structure, ensuring a strong interface bond between the upper and lower layers and eliminating the risk of delamination. Furthermore, the fabrication process is controllable and highly customizable, allowing for flexible adjustment of the scaffold's size and pore structure according to the clinical defect morphology.

[0019] This invention applies a double-layered scaffold with differentiated degradation and mechanical properties to the clinical treatment of brain tissue damage and skull defects after craniotomy decompression for traumatic brain injury. It can simultaneously meet the dual needs of rapid anti-inflammatory repair of brain tissue and long-term regenerative repair of skull, realizing integrated brain-bone repair and providing an effective medical material solution for the simultaneous treatment of defects after craniotomy decompression for traumatic brain injury. Attached Figure Description

[0020] To more clearly illustrate the technical solutions of the embodiments of the present invention, the drawings used in the following description of the embodiments will be briefly introduced. Obviously, the drawings described below are some embodiments of the present invention. For those skilled in the art, other drawings can be obtained based on these drawings without creative effort.

[0021] Appendix Figure 1 Transmission electron microscopy image of nanofat emulsion (FE). The scale bar represents 200 nm, showing that FE consists of spherical nanoparticles with a diameter of 100–200 nm.

[0022] Appendix Figure 2 Rheological test curves of the lower and upper hydrogels with different FE concentrations.

[0023] Appendix Figure 3 Scanning electron microscope images (left) and swelling rate curves (right) of different hydrogel scaffolds.

[0024] Appendix Figure 4 Curves showing the cumulative DHA release from the upper hydrogel (FE / C / G) and lower hydrogel (FE / G) with different FE concentrations.

[0025] Appendix Figure 5 : In vitro toxicity test results of FE on BV2 cells (left) and hBMSC cells (right) using CCK-8 assay.

[0026] Appendix Figure 6 : Characterization of the anti-inflammatory effect of FE.

[0027] Appendix Figure 7 Characterization of the angiogenesis-promoting effect of FE.

[0028] Appendix Figure 8 Results of ALP staining (top) and ARS staining (bottom) for FE-promoted osteogenicity.

[0029] Appendix Figure 9 mNSS score curve of mice treated with bilayer hydrogel scaffold. Detailed Implementation

[0030] The technical solutions in the embodiments of the present invention will be clearly and completely described below with reference to the accompanying drawings. Obviously, the embodiments described below are only some embodiments of the present invention, and not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without creative effort are within the scope of protection of the present invention.

[0031] It should be understood that, when used in this specification and the appended claims, the terms "comprising" and "including" indicate the presence of the described features, integrals, steps, operations, elements and / or components, but do not exclude the presence or addition of one or more other features, integrals, steps, operations, elements, components and / or collections thereof.

[0032] It should also be understood that the terminology used in the specification of embodiments of the present invention is for the purpose of describing particular embodiments only and is not intended to limit the embodiments of the present invention. As used in the specification of embodiments of the present invention and the appended claims, unless the context clearly indicates otherwise, the singular forms "a," "an," and "the" are intended to include the plural forms.

[0033] This invention provides a bilayer hydrogel scaffold, comprising: The upper hydrogel layer is made of a first photocrosslinked hydrogel and is loaded with nano-fat emulsion. The lower hydrogel layer is made of a second photocrosslinked hydrogel and is loaded with nano-fat emulsion. The first photocrosslinked hydrogel is a composite hydrogel of methacrylamide chitosan and methacrylamide gelatin (CSMA@GelMA), and the second photocrosslinked hydrogel is a methacrylamide gelatin hydrogel (GelMA); the thickness of the upper hydrogel layer is greater than or equal to the thickness of the lower hydrogel layer.

[0034] In specific implementation, both layers of the bilayer hydrogel scaffold of this invention are loaded with nano-fat emulsion (FE). The upper layer uses CSMA@GelMA, which has high mechanical strength, and the loaded FE is released slowly, enabling long-term continuous release to promote angiogenesis and bone regeneration, meeting the needs of long cranial defect repair cycles and the requirement for continuous bioactivity. The lower layer uses GelMA, which has lower mechanical strength and a softer texture, avoiding secondary damage to damaged brain tissue. The loaded FE can be released rapidly, inhibiting neuroinflammation and reducing secondary damage in the early stages.

[0035] Fe, a nanolipid formulation rich in polyunsaturated fatty acids (PUFAs), possesses multiple biological activities, including anti-inflammatory, angiogenic, and osteogenic effects. However, liquid FE is difficult to stably remain at the site of skull defects and achieve effective phased release. This invention loads FE into a bilayer hydrogel with differentiated degradation rates, utilizing the characteristics of sustained release in the upper layer and rapid release in the lower layer. Within the same scaffold, it simultaneously meets the needs of rapid anti-inflammatory repair of brain tissue and long-term regenerative repair of the skull, achieving integrated brain-bone repair. Furthermore, this scaffold has high biocompatibility, readily available raw materials, and a wide range of sources.

[0036] In this invention, the concentration range of FE is 0.1–10 mg / mL, preferably 0.2–5 mg / mL, for example 0.1 mg / mL, 0.2 mg / mL, 0.5 mg / mL, 1 mg / mL, 2 mg / mL, 5 mg / mL, 8 mg / mL, or 10 mg / mL. If the FE concentration is below 0.1 mg / mL, it is difficult to achieve an effective therapeutic dose; if it is above 10 mg / mL, it may inhibit cell proliferation.

[0037] In specific implementation, the upper hydrogel layer of this invention uses a composite hydrogel of methacrylated chitosan and methacrylated gelatin (CSMA@GelMA). Methacrylated gelatin (GelMA) is a photocrosslinked modified gelatin prepared by reacting gelatin with methacrylic anhydride. It retains the cell adhesion sites and enzyme degradation sites of gelatin while introducing photocrosslinking groups, allowing it to form a three-dimensional network structure under ultraviolet light irradiation in the presence of a photoinitiator. Methacrylated chitosan (CSMA) is a photocrosslinked modified polysaccharide prepared by reacting chitosan with methacrylic anhydride, possessing anti-inflammatory and cartilage / bone repair-promoting bioactivity.

[0038] CSMA and GelMA are mixed in a certain proportion and cross-linked under the action of a photoinitiator and ultraviolet light to form a CSMA@GelMA composite hydrogel. The introduction of CSMA increases the cross-linking density and mechanical strength of the hydrogel, while slowing down its degradation rate. In this invention, the upper hydrogel layer utilizes the high mechanical strength of CSMA@GelMA to provide support for skull defects, and its slow degradation rate enables the long-term slow release of the loaded nano-fat emulsion (FE), continuously exerting its angiogenesis and osteogenic effects.

[0039] The degree of substitution of GelMA ranges from 30% to 90%, for example, 30%, 50%, 80%, and 90%. The degree of substitution affects the photocrosslinking density: the higher the degree of substitution, the more methacryloyl groups can be crosslinked, resulting in higher mechanical strength and slower degradation rate after crosslinking; if the degree of substitution is too low (<30%), the crosslinking is insufficient and the scaffold strength is inadequate; if the degree of substitution is too high (>90%), it may lead to over-crosslinking, making the material brittle and degrading too slowly.

[0040] The GelMA concentration range is 1–30 w / v%, for example, 10 w / v%, 15 w / v%, 20 w / v%, 25 w / v%, or 30 w / v can be used; if the GelMA concentration is less than 1 w / v%, the mechanical strength after cross-linking is too low to support the skull defect; if it is higher than 30 w / v%, the ink viscosity is too high and it is difficult to be extruded through the 3D printing nozzle.

[0041] The degree of substitution of CSMA ranges from 30% to 90%, for example, 30%, 50%, 80%, and 90%. Similar to GelMA, the degree of substitution affects the crosslinking density and degradation rate.

[0042] The CSMA concentration range is 1–25 w / v%, for example, 5 w / v%, 10 w / v%, 15 w / v%, or 25 w / v can be used. If the CSMA concentration is below 1 w / v%, the degradation rate and mechanical strength of the upper hydrogel are not significantly different from those of the lower layer, and differential repair cannot be achieved; if it is above 25 w / v%, the cross-linking is too high, the degradation is too slow, the FE is difficult to release effectively, and it may cause local inflammatory reactions.

[0043] In specific implementation, the lower hydrogel layer of this invention uses methacrylamide gelatin hydrogel (GelMA). GelMA is a photocrosslinked modified gelatin prepared by reacting gelatin with methacrylic anhydride. Its synthesis method is as follows: gelatin is dissolved in phosphate buffer, an appropriate amount of methacrylic anhydride is added for reaction, and the product is obtained by dialysis and lyophilization. GelMA retains the inherent arginine-glycine-aspartic acid (RGD) cell adhesion sequence and matrix metalloproteinase (MMP) sensitive degradation sites of gelatin, exhibiting good cell compatibility and degradability; at the same time, the introduced methacrylamide group enables it to rapidly crosslink into a gel under ultraviolet light irradiation in the presence of a photoinitiator.

[0044] In this invention, the lower hydrogel layer utilizes the low mechanical strength and fast degradation rate of GelMA hydrogel. On the one hand, its soft properties prevent secondary compression damage to the damaged brain tissue. On the other hand, it degrades rapidly after implantation, allowing the loaded FE to be released quickly, penetrate the dura mater in time, act on microglia, inhibit the expression of pro-inflammatory factors, and reduce neuroinflammation and secondary damage.

[0045] The degree of substitution of GelMA ranges from 30% to 90%, for example, 30%, 50%, 80%, and 90%. The degree of substitution affects the crosslinking density and degradation rate: a moderate degree of substitution (e.g., 50-70%) yields good mechanical properties and degradation rate; too low a degree of substitution results in insufficient crosslinking, and the scaffold is prone to collapse; too high a degree of substitution results in slow degradation, and FE cannot be released quickly.

[0046] The concentration range is 1–30 w / v%, for example, 10 w / v%, 15 w / v%, 20 w / v%, 25 w / v%, or 30 w / v can be used. If the GelMA concentration is below 1 w / v%, the cross-linked hydrogel strength is too low, making it difficult to maintain the three-dimensional structure, and the rapid degradation leads to the explosive release of FE, which may cause excessively high local concentrations and toxicity. If it is above 30 w / v%, the hydrogel is too dense, the degradation is too slow, and FE cannot be released rapidly in the early stages, affecting the anti-inflammatory effect. At the same time, the increased mechanical strength may cause damage to brain tissue.

[0047] In specific implementations, the thickness of the upper hydrogel layer of this invention is 0.1–10 mm, and the thickness of the lower hydrogel layer is 0.1–5 mm, with the thickness of the upper hydrogel layer being greater than or equal to the thickness of the lower hydrogel layer. For example, the thickness of the upper layer can be 0.5 mm, 1 mm, 2 mm, 5 mm, or 8 mm, and the thickness of the lower layer can be 0.2 mm, 0.5 mm, 1 mm, 2 mm, or 3 mm. If the thickness of the upper layer is less than 0.1 mm, it is difficult to provide sufficient mechanical support and long-term release of active ingredients; if the thickness of the upper layer is greater than 10 mm, it may exceed the conventional range of skull defects, which is not conducive to implantation and healing. If the thickness of the lower layer is less than 0.1 mm, the FE release is insufficient, making it difficult to quickly inhibit neuroinflammation; if the thickness of the lower layer is greater than 5 mm, it may produce excessive space-occupying effect on brain tissue, increasing the risk of damage. Ensuring that the thickness of the upper layer is greater than or equal to the thickness of the lower layer allows the skull side, which has higher mechanical requirements, to obtain sufficient material support, while the soft lower layer maintains a thinner shape to conform to the surface of brain tissue.

[0048] In specific implementation, the bilayer hydrogel scaffold of the present invention has a three-dimensional porous structure. This porous structure is conducive to cell adhesion, migration, and proliferation, and promotes angiogenesis and the formation of new tissue. When fabricated using 3D printing, the porous structure is achieved by setting the diameter of the printing lines to a range of 0.1–5 mm and the pore size between the lines to a range of 0–1 mm. For example, the line diameter can be 0.2 mm, 0.5 mm, 1 mm, 2 mm, or 4 mm, and the pore size can be 0 mm (i.e., dense structure), 0.2 mm, 0.5 mm, or 0.8 mm. If the line diameter is less than 0.1 mm, the scaffold's mechanical strength is too low, making it difficult to maintain its three-dimensional morphology; if the line diameter is greater than 5 mm, the specific surface area is too small, which is not conducive to cell adhesion and substance exchange. If the pore size is greater than 1 mm, the scaffold structure is too sparse, resulting in insufficient mechanical strength, and cells cannot effectively spread and migrate within the excessively large pores; if the pore size is 0 mm, the scaffold has a dense structure, and although it can still achieve repair functions through material degradation and the release of active ingredients, the cell infiltration and tissue ingrowth capabilities will be reduced. Therefore, the pore size can be adjusted within the range of 0 to 1 mm according to specific repair needs.

[0049] This invention also provides a method for preparing a bilayer hydrogel scaffold, comprising the following steps: (1) Prepare the upper layer printing ink and the lower layer printing ink.

[0050] In practice, the lower-layer printing ink comprises a second photocrosslinking hydrogel precursor, a nanofat emulsion (FE), and a photoinitiator. The second photocrosslinking hydrogel precursor is methacrylamide gelatin (GelMA), and the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonite (LAP). GelMA is dissolved in phosphate buffer or cell culture medium, and LAP is added to a final concentration of 0.1–1 w / v% (e.g., 0.2 w / v%, 0.5 w / v%, 0.8 w / v%, 1 w / v%). Then, the FE solution is added, and the mixture is stirred until homogeneous to obtain the FE / GelMA printing ink.

[0051] If the LAP concentration is below 0.1 w / v%, the photocrosslinking speed is too slow, the molding efficiency is low, and the crosslinking is insufficient; if it is above 1 w / v%, it may cause cytotoxicity.

[0052] The upper printing ink comprises a first photocrosslinking hydrogel precursor, FE, and a photoinitiator. The first photocrosslinking hydrogel precursor is a mixture of methacrylamide chitosan (CSMA) and GelMA. CSMA and GelMA are dissolved in phosphate buffer or cell culture medium, and LAP is added to a final concentration of 0.1–1 w / v% (e.g., 0.1 w / v%, 0.2 w / v%, 0.5 w / v%, 0.8 w / v%, 1 w / v%). Then, FE solution is added, and the mixture is stirred until homogeneous to obtain the FE / CSMA / GelMA printing ink.

[0053] (2) Using 3D printing technology, the lower and upper printing inks are printed layer by layer according to the preset three-dimensional model, and photocrosslinking and curing are performed on each printed layer.

[0054] Using a 3D bioprinter, the prepared lower and upper printing inks are loaded into printing cartridges respectively. Based on the size and shape of the target defect, a three-dimensional model of the double-layer hydrogel scaffold is created using computer-aided design software, setting parameters such as the thickness and porous structure (line diameter, pore size) of the upper and lower layers.

[0055] During printing, print the lower layer first: Install the ink cartridge for the lower layer onto the printer, and set the printing process parameters, such as printing speed 5–20 mm / s (e.g., 5 mm / s, 8 mm / s, 12 mm / s, 15 mm / s, 20 mm / s), extrusion air pressure 1.5–3.0 bar (e.g., 1.5 bar, 2.0 bar, 2.5 bar, 3.0 bar), nozzle diameter 300–800 μm (e.g., 300 μm, 500 μm, 600 μm, 800 μm), line spacing 300–800 μm (e.g., 300 μm, 500 μm, 600 μm, 800 μm), and layer spacing 100–300 μm (e.g., 100 μm, 200 μm, 300 μm). Print the lower layers layer by layer according to the slicing data of the 3D model. After printing each layer or slice, perform photocrosslinking and curing.

[0056] After the lower layer is printed, the ink cartridge for the upper layer is replaced, the number of printing layers corresponding to the upper layer thickness is adjusted, and the upper layer is printed using the same or similar printing parameters. After each layer is printed, photocrosslinking and curing are performed.

[0057] If the printing speed is below 5 mm / s, production efficiency is low, and prolonged extrusion may cause ink cross-linking within the needle; if it is above 20 mm / s, lines may break or have uneven width. If the extrusion air pressure is below 1.5 bar, ink extrusion is insufficient, resulting in discontinuous lines; if it is above 3.0 bar, over-extrusion occurs, resulting in excessively thick lines and reduced forming accuracy. If the needle diameter is less than 300 μm, clogging is likely, especially with high-viscosity inks containing CSMA; if it is greater than 800 μm, lines are too thick, making it difficult to achieve fine structures. If the line spacing is less than 300 μm, the pores are too small, hindering cell infiltration; if it is greater than 800 μm, the pores are too large, resulting in insufficient scaffold mechanical strength. If the interlayer spacing is less than 100 μm, excessive interlayer fusion occurs, and pores disappear; if it is greater than 300 μm, the interlayer bonding is weak, and the scaffold is prone to delamination. Those skilled in the art can adjust the printing parameters according to actual performance.

[0058] Immediately after each printed layer or slicing layer, photocrosslinking is performed using ultraviolet light. The photocrosslinking conditions are: wavelength 365 nm, irradiation time 10–60 seconds, e.g., 20, 30, or 40 seconds. The photoinitiator LAP generates free radicals under ultraviolet light irradiation, initiating crosslinking polymerization of the methacryloyl groups in GelMA and CSMA molecules, forming a three-dimensional network structure, transforming the printed lines from a liquid state to a solid hydrogel. If the irradiation time is less than 10 seconds, the crosslinking is insufficient, resulting in inadequate scaffold strength and inability to maintain the shape; if it is more than 60 seconds, over-crosslinking may occur, altering the material's degradation properties, and prolonged ultraviolet irradiation may cause oxidative damage to the loaded FE or produce cytotoxicity. Immediate crosslinking after each layer ensures that the current layer solidifies before printing the next, thereby achieving strong fusion between upper and lower layers and between lines within the same layer, ultimately obtaining a bilayer hydrogel scaffold where the upper and lower hydrogel layers are bonded together through photocrosslinking.

[0059] (3) A double-layer hydrogel scaffold is obtained by photocrosslinking and fusing the upper hydrogel layer and the lower hydrogel layer together through the above steps.

[0060] In practice, the printed double-layer hydrogel scaffold can be processed according to subsequent application needs. For example, it can be used directly in animal experiments or clinical implantation; it can also be soaked and washed with phosphate buffer to remove unreacted monomers and excess photoinitiators; or it can be freeze-dried to obtain a dried and preserved scaffold for long-term storage and transportation. The specific conditions for freeze-drying are as follows: pre-freeze the scaffold at -80°C for 2–4 hours, then transfer it to a freeze dryer and dry it at -50°C and a vacuum degree <10 Pa for 24–48 hours. Before use, the dried scaffold needs to be rehydrated with sterile phosphate buffer or cell culture medium. If cell culture or implantation experiments are required, aseptic operation should be performed in a clean bench, and all solutions and consumables must be pre-sterilized.

[0061] The preparation method of this invention can produce a bilayer hydrogel scaffold with differentiated degradation rates and mechanical properties. The upper layer of the scaffold degrades slowly and has high mechanical strength, while the lower layer degrades quickly and has low mechanical strength. Both layers are loaded with FE. Furthermore, the size, shape, and pore structure of the scaffold can be precisely controlled by 3D printing to meet the personalized repair needs of different patients with skull defects.

[0062] The following examples illustrate the invention in detail. The nano-fat emulsion (FE) used in the embodiments of the present invention is a commercially available medium / long-chain fat emulsion injection (C8-24Ve, Guangdong Jiabo Pharmaceutical Co., Ltd., National Drug Approval Number H20213040). This product is a commonly used clinical nutritional supplement preparation, rich in polyunsaturated fatty acids. It should be noted that this source is merely illustrative, and the present invention is not limited thereto. Nano-fat emulsions rich in polyunsaturated fatty acids from other sources or compositions are equally applicable. Those skilled in the art can prepare their own or select other commercially available products according to actual needs.

[0063] Example 1: Morphological observation of nano-fat emulsions Experimental Methods: Nanofat emulsions (FE) were loaded onto copper mesh and air-dried in a dry environment. After air drying, the morphology and structure of the FE were observed using transmission electron microscopy (TEM). TEM projects an accelerated and focused electron beam onto the sample. Electrons collide with atoms in the sample, producing solid-angle scattering. The scattering angle is related to the sample density and thickness, forming images of varying brightness, thus obtaining information about the sample's morphology.

[0064] Experimental results: as attached Figure 1 As shown, FE consists of spherical nanoparticles with a diameter of 100–200 nm. This result indicates that FE possesses a nanoscale structure, which facilitates its penetration of biological barriers (such as the dura mater) and entry into brain tissue to exert its effects.

[0065] Example 2: Mechanical property testing of a double-layer hydrogel scaffold Experimental method: A series of hydrogel samples with lower and upper layers were prepared and their mechanical properties were compared and tested.

[0066] Lower series (FE / G): GelMA (substitution degree 30%) concentration 15 w / v%, LAP 0.2 w / v%, FE concentrations were set to 0 (G), 0.2 mg / mL (0.2FE / G), 0.5 mg / mL (0.5FE / G), 1.0 mg / mL (1FE / G), 2.0 mg / mL (2FE / G), and 5.0 mg / mL (5FE / G).

[0067] Upper layer series (FE / C / G): GelMA (substitution degree 30%) concentration 15 w / v%, CSMA concentration 10 w / v%, LAP 0.2 w / v%, FE concentrations were set to 0 (C / G), 0.2 mg / mL (0.2FE / C / G), 0.5 mg / mL (0.5FE / C / G), 1.0 mg / mL (1FE / C / G), 2.0 mg / mL (2FE / C / G), and 5.0 mg / mL (5FE / C / G).

[0068] After thoroughly mixing the above-mentioned printing inks, a rotational rheometer was used for testing. The testing mode was frequency scanning, with a frequency range of 0.1–10 Hz, a strain of 1%, and a temperature of 37°C. The storage modulus (G') of each sample after photocrosslinking was measured to evaluate the effect of FE concentration on the mechanical strength of the hydrogel and the changes in mechanical properties after the introduction of CSMA.

[0069] Experimental results: as attached Figure 2 As shown, for the lower layer series (FE / G), the storage modulus of the hydrogel gradually increased with the increase of FE concentration from 0 to 5 mg / mL, indicating that the addition of FE can enhance the mechanical strength of the GelMA hydrogel in a concentration-dependent manner. For the upper layer series (FE / C / G), the storage modulus of the C / G group without FE was significantly higher than that of the G group without FE, indicating that the introduction of CSMA significantly improved the mechanical strength of the hydrogel; at the same time, the storage modulus of the FE / C / G series also showed an upward trend with the increase of FE concentration. At all concentrations, the mechanical strength of the FE / C / G series was higher than that of the FE / G series with the corresponding FE concentration. These results indicate that FE can be loaded into the hydrogel as an active ingredient without impairing its gelling properties. Furthermore, by controlling the FE concentration and the addition of CSMA, a gradient adjustment of mechanical strength can be achieved, meeting the support requirements of the upper scaffold for skull defects and the soft adaptation requirements of the lower scaffold for brain tissue.

[0070] Example 3: Microstructure and release behavior of a bilayer hydrogel scaffold Experimental method: Lower layer printing ink (FE / G) and upper layer printing ink (FE / C / G) with different FE concentrations were prepared respectively. Among them: The lower layer (FE / G) contained 15 w / v GelMA (30% substitution) and 0.2 w / v LAP. The FE concentrations were set at 0.2 mg / mL (FE / G200), 0.5 mg / mL (FE / G500), and 1.0 mg / mL (FE / G1000), respectively, and a control group (G) without FE was set up.

[0071] Upper layer (FE / C / G): GelMA (substitution degree 30%) concentration 15 w / v%, CSMA concentration 10 w / v%, LAP 0.2 w / v%, FE concentrations were set at 0.2 mg / mL (FE / C / G200), 0.5 mg / mL (FE / C / G500), and 1.0 mg / mL (FE / C / G1000), respectively, and a control group without FE (C / G) was set up.

[0072] Using an Envisiontec 3D-Bioplotter bioprinter with a low-temperature printhead equipped with a 600 μm diameter print needle, the printing speed was set to 10 mm / s, the extrusion pressure to 2.2 bar, the row spacing to 500 μm, and the layer spacing to 200 μm, and the hydrogel scaffolds of the above compositions were printed respectively.

[0073] FE / G and FE / C / G scaffolds with a FE concentration of 0.5 mg / mL were lyophilized and then imaged using a scanning electron microscope (SEM) to observe the microstructure of the scaffolds. Separately, undyophilized scaffolds of the corresponding concentrations were used for swelling tests: the initial weight of the scaffolds was weighed, and they were immersed in 10 times their volume of PBS, placed in a 37°C oven, and weighed every 12 hours until the weight remained constant. The swelling rate was then calculated.

[0074] FE / G (200, 500, 1000 μg / mL) and FE / C / G (200, 500, 1000 μg / mL) scaffolds with different FE concentrations were immersed in PBS and incubated at 37°C. The supernatant was collected weekly to determine the docosahexaenoic acid (DHA) content, in order to evaluate the FE release behavior.

[0075] Experimental results: as attached Figure 3 As shown in the SEM images, the hydrogel scaffold has a regular porous structure with abundant micropores distributed on its surface. The addition of FE increases the pore size of the scaffold, which is beneficial for cell adhesion, blood vessel ingrowth, and osteogenic formation. Swelling tests show that the scaffold has good water absorption properties.

[0076] As attached Figure 4As shown, the left figure represents the cumulative DHA release curves of the lower hydrogel (FE / G) with different FE concentrations, and the right figure represents the cumulative DHA release curves of the upper hydrogel (FE / C / G) with different FE concentrations. The results show that all groups exhibit a typical release pattern of rapid initial release followed by slow release later. For the lower hydrogel (FE / G), FE / G200, FE / G500, and FE / G1000 all reached a plateau in release within approximately 3 weeks. Higher FE concentrations resulted in greater cumulative release, but the release rate showed no significant difference, indicating that GelMA hydrogels can achieve rapid FE release. For the upper hydrogel (FE / C / G), the release period of FE / C / G200, FE / C / G500, and FE / C / G1000 was significantly prolonged, continuing to release for over 8 weeks. The release rate increased slightly with increasing FE concentration, but was still much slower than that of the lower hydrogel. These results indicate that differentiated release of FE can be achieved by selecting different hydrogel materials: rapid release from the lower layer to inhibit neuroinflammation in the early stages, and slow release from the upper layer to promote angiogenesis and bone regeneration in the long term.

[0077] Example 4: Evaluation of the cell compatibility of nano-fat emulsions Experimental methods: BV2 cells (microglia) and hBMSCs (human bone marrow mesenchymal stem cells) were cultured with different concentrations of FE (0, 50, 100, 200, 500 μg / mL). After 24 hours of culture, the CCK-8 assay was performed to evaluate the toxicity of FE to the cells.

[0078] Experimental results: as attached Figure 5 As shown, FE concentrations below 200 μg / mL showed no significant toxicity to either cell type, and cell proliferation activity remained normal. However, FE concentrations above 200 μg / mL inhibited cell proliferation. Therefore, a FE concentration of 200 μg / mL was selected for subsequent cell experiments.

[0079] Example 5: Evaluation of the anti-inflammatory effect of nano-fat emulsion Experimental methods: BV2 cells were treated with 200 μg / mL FE and 1 mg / mL lipopolysaccharide (LPS) for 24 hours. Cell RNA was extracted, and the expression levels of pro-inflammatory genes (TNF-α, MCP-1) and anti-inflammatory genes (IL-10, Arg-1) were detected by real-time quantitative PCR (RT-qPCR).

[0080] Experimental results: as attached Figure 6As shown, compared with the LPS-only treatment group, the FE+LPS combined treatment group significantly inhibited the expression of pro-inflammatory genes TNF-α and MCP-1, while promoting the expression of anti-inflammatory genes IL-10 and Arg-1. This result indicates that FE can effectively regulate the inflammatory phenotype of microglia, shifting them from a pro-inflammatory to an anti-inflammatory state, thereby exerting an inhibitory effect on neuroinflammation.

[0081] Example 6: Preparation of Immunomodulatory Culture Medium Experimental methods: Mouse microglia (BV2) were treated as follows, and culture media with different conditions were prepared in groups: Blank control group (black): BV2 cells were cultured routinely for 24 hours without any treatment, and the supernatant was collected.

[0082] Model control group (con): BV2 cells were treated with 1 mg / mL lipopolysaccharide (LPS) for 24 hours, and the supernatant was collected.

[0083] FE group alone (200FE): BV2 cells were treated with 200 μg / mL nanofat emulsion (FE) for 24 hours and the supernatant was collected.

[0084] Inflammation-induced group (BV2+LPS): Same as the model control group, i.e., treated with 1 mg / mL LPS for 24 hours, and the supernatant was collected (essentially the same as the con group, but used as an independent group for subsequent experiments).

[0085] Inflammation + FE intervention group (BV2 + LPS + 200FE): BV2 cells were pretreated with 200 μg / mL FE for 1 hour, and then treated with 1 mg / mL LPS for a total of 24 hours. The supernatant was collected.

[0086] The cell supernatants collected from each group were centrifuged to remove cell debris, yielding immunomodulatory culture media. These media were then used for subsequent cell experiments. The above immunomodulatory medium and DMEM medium were mixed at a ratio of 1:1 to obtain the medium for HUVEC cell (human umbilical vein endothelial cell) tube formation experiment.

[0087] The above immunomodulatory medium was mixed with twice the amount of osteogenic induction medium at a ratio of 1:1 to obtain the osteogenic induction medium for hBMSC cells (human bone marrow mesenchymal stem cells).

[0088] Example 7: Evaluation of the angiogenesis effect of nano-fat emulsions Experimental Methods: HUVEC cells were resuspended in the culture medium prepared in Example 6 for tube formation experiments and seeded onto matrix gel-coated culture plates, with three replicates per group. The following four groups were selected for testing: Blank control group (blank): Culture medium prepared using the BV2 supernatant from the blank group.

[0089] FE group alone (200FE): Culture medium prepared using the supernatant of BV2 from group 200FE.

[0090] Inflammation-inducing group (BV2+LPS): Culture medium prepared using BV2 supernatant from the BV2+LPS group.

[0091] Inflammation + FE intervention group (200FE + BV2 + LPS): Culture medium prepared using BV2 supernatant from the 200FE + BV2 + LPS group.

[0092] After 6 hours of culture, the cells were stained with calcein and observed under a fluorescence microscope to observe the tube formation.

[0093] Experimental results: as attached Figure 7 As shown, in the blank control group (blank), HUVEC cells formed basically complete tubular structures; in the FE-only group (200FE), the tubular structures were denser and more complete, indicating that FE directly promotes endothelial cell tubulation; in the inflammation-induced group (BV2+LPS), the cell tubulation ability was significantly reduced, and the tubular structures were broken and sparse; while in the inflammation + FE intervention group (200FE+BV2+LPS), the tubular structures were significantly restored, approaching the level of the blank control group. These results indicate that FE can significantly promote the tubulation ability of vascular endothelial cells, which is beneficial for angiogenesis during the repair of skull defects.

[0094] Example 8: Evaluation of the osteogenic effect of nano-fat emulsions Experimental Methods: BMSCs were seeded into 12-well plates and cultured normally for 3 days. Then, osteogenic induction culture was performed using the osteogenic induction medium prepared in Example 6. The following five groups were selected for testing: Blank control group (blank): Osteogenic induction culture medium prepared using BV2 supernatant from the blank group.

[0095] Model control group (con): osteogenic induction culture medium prepared using BV2 supernatant from the con group.

[0096] FE group alone (200FE): Osteogenic induction medium prepared using BV2 supernatant from 200FE group.

[0097] Inflammation-inducing group (BV2+LPS): Osteogenic induction medium prepared using BV2 supernatant from the BV2+LPS group.

[0098] Inflammation + FE intervention group (200FE + BV2 + LPS): Osteogenic induction culture medium prepared using BV2 supernatant from the 200FE + BV2 + LPS group.

[0099] Cells were fixed with 4% paraformaldehyde 7 and 14 days after osteogenic induction, stained with alkaline phosphatase (ALP) for 30 minutes, and observed under a microscope (to assess early osteogenic differentiation). Cells were fixed with 4% paraformaldehyde 14 and 21 days after osteogenic induction, stained with Alizarin Red (ARS) for 5 minutes, and observed under a microscope (to assess late mineralization).

[0100] Experimental results: as attached Figure 8 As shown, ALP staining results revealed that, compared to the blank control group, the group with added FE had a deeper color than the control group without FE. Furthermore, under immunomodulatory conditions, the group with added FE also had a deeper color, indicating that FE promotes early osteogenic differentiation of BMSCs. The inflammation-induced group (BV2+LPS) had a lighter color, indicating that the inflammatory environment inhibited osteogenic differentiation. The inflammation + FE intervention group (200FE+BV2+LPS) had a significantly deeper color than the inflammation-induced group, indicating that FE could reverse the inhibitory effect of the inflammatory environment on osteogenic differentiation. ARS staining results were consistent with ALP; the inflammation + FE intervention group had more mineralized nodules than the inflammation-induced group, and these nodules were also darker. These results demonstrate that FE can promote osteogenic differentiation of BMSCs under both normal and inflammatory immune environments, including enhanced early ALP activity and the formation of late-stage mineralized nodules, thus benefiting bone regeneration during skull defect repair.

[0101] Example 9: Establishment of a mouse model of traumatic brain injury Experimental Methods: Eight-week-old C57BL / 6J mice were anesthetized with gas, and their head hair was removed and the skull exposed by incision. A circular bone window was opened 2 mm to the right of the sagittal suture using a 4 mm diameter dental drill. Type II traumatic brain injury (TBI) was induced in the mice using a precision impactor with the following parameters: impact tip diameter 3 mm, impact velocity 3.5 m / s, impact depth 1 mm, and dwell time 0.5 s.

[0102] Example 10: In vivo efficacy evaluation of bilayer hydrogel scaffold Experimental Methods: After TBI modeling, mice were randomly divided into four groups: blank control group (Sham), model control group (Control), double-layer hydrogel scaffold implanted without FE loading (C / G), and double-layer hydrogel scaffold implanted with FE loading (FE / C / G). The scaffold groups had the corresponding scaffold implanted at the brain injury site. The modified neurological deficit score (mNSS) was used to assess the neurological function recovery of each group at different time points (1, 3, 5, and 7 days after implantation). A higher mNSS score indicated a more severe neurological deficit.

[0103] Experimental results: as attached Figure 9As shown, the mNSS score in the model control group (Control) remained at a high level. The unloaded FE scaffold group (C / G) showed some improvement, but the effect was limited. The FE-loaded bilayer hydrogel scaffold group (FE / C / G) significantly reduced the mNSS score in the early post-implantation period (approximately 7 days) and maintained a low level at subsequent time points, indicating that the FE / C / G scaffold can rapidly and effectively promote the recovery of neurological function in TBI mice.

[0104] The results of Examples 1-10 above collectively demonstrate that the FE-loaded bilayer hydrogel scaffold provided by the present invention has good biocompatibility. FE has triple activity of inhibiting neuroinflammation, promoting angiogenesis, and promoting osteogenic formation. Through the differentiated design of upper layer sustained release and lower layer rapid release, it can simultaneously promote brain tissue repair and skull defect regeneration after TBI craniotomy decompression, achieving integrated brain-bone repair.

[0105] In the above embodiments, the descriptions of each embodiment have different focuses. For parts that are not described in detail in a certain embodiment, please refer to the relevant descriptions in other embodiments.

[0106] The above description describes specific embodiments of the present invention, but the scope of protection of the present invention is not limited thereto. Any person skilled in the art can easily conceive of various equivalent modifications or substitutions within the technical scope disclosed in the present invention, and these modifications or substitutions should all be covered within the scope of protection of the present invention. Therefore, the scope of protection of the present invention should be determined by the scope of the claims.

Claims

1. A double-layer hydrogel scaffold, characterized in that, include: The upper hydrogel layer is made of a first photocrosslinked hydrogel and is loaded with nano-fat emulsion. The lower hydrogel layer is made of a second photocrosslinked hydrogel and is loaded with nano-fat emulsion. The first photocrosslinked hydrogel is a composite hydrogel of methacrylamide chitosan and methacrylamide gelatin, and the second photocrosslinked hydrogel is a methacrylamide gelatin hydrogel. The thickness of the upper hydrogel layer is greater than or equal to the thickness of the lower hydrogel layer.

2. The double-layer hydrogel scaffold according to claim 1, characterized in that, The degradation rate of the upper hydrogel layer is slower than that of the lower hydrogel layer, and the mechanical strength of the upper hydrogel layer is higher than that of the lower hydrogel layer.

3. The double-layer hydrogel scaffold according to claim 1 or 2, characterized in that, The concentration of the nanofat emulsion is 0.1–10 mg / mL.

4. The double-layer hydrogel scaffold according to claim 1, characterized in that, The thickness of the upper hydrogel layer is 0.1–10 mm, and the thickness of the lower hydrogel layer is 0.1–5 mm.

5. The double-layer hydrogel scaffold according to claim 1, characterized in that, The double-layer hydrogel scaffold has a three-dimensional porous structure.

6. A method for preparing the bilayer hydrogel scaffold according to any one of claims 1-5, characterized in that, Includes the following steps: (1) Prepare upper printing ink and lower printing ink, wherein the upper printing ink comprises a first photocrosslinking hydrogel precursor, a nano-fat emulsion and a photoinitiator, and the lower printing ink comprises a second photocrosslinking hydrogel precursor, a nano-fat emulsion and a photoinitiator; (2) Using 3D printing technology, the lower layer printing ink and the upper layer printing ink are printed layer by layer according to the preset three-dimensional model, and photocrosslinking and curing are performed on each printed layer; (3) Obtain a double-layer hydrogel scaffold in which the upper hydrogel layer and the lower hydrogel layer are fused together by photocrosslinking.

7. The method according to claim 6, characterized in that, The first photocrosslinking hydrogel precursor is a mixture of methacrylamide chitosan and methacrylamide gelatin, and the second photocrosslinking hydrogel precursor is methacrylamide gelatin; the degree of substitution of the methacrylamide gelatin is 30-90%, and the concentration is 1-30 w / v; the degree of substitution of the methacrylamide chitosan is 30-90%, and the concentration is 1-25 w / v; the photoinitiator is lithium phenyl-2,4,6-trimethylbenzoylphosphonite, and the concentration is 0.1-1 w / v.

8. The method according to claim 6, characterized in that, The conditions for photocrosslinking are ultraviolet light wavelength of 365 nm and irradiation time of 10 to 60 seconds.

9. The method according to claim 6, characterized in that, The 3D printing process parameters are as follows: printing speed 5-20 mm / s, extrusion air pressure 1.5-3.0 bar, needle diameter 300-800 μm, row spacing 300-800 μm, and layer spacing 100-300 μm.

10. The use of the bilayer hydrogel scaffold according to any one of claims 1-5 or the bilayer hydrogel scaffold prepared by the method according to any one of claims 6-9 in the preparation of a repair material for treating brain tissue damage and skull defects after craniotomy decompression for traumatic brain injury.

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