An injectable polymer hydrogel and a preparation process thereof

Abstract

This invention relates to the field of medical polymer materials technology, and discloses an injectable polymer hydrogel and its preparation process. The hydrogel is prepared by in-situ crosslinking of solution A and solution B at a volume ratio of 1:1. Solution A includes oxidized dextran, zinc gluconate, anhydrous betaine, and phosphate buffer, while solution B includes carboxymethyl chitosan and phosphate buffer. This invention utilizes oxidized dextran and carboxymethyl chitosan to construct Schiff base covalent crosslinks, and introduces zinc gluconate to provide delayed metal coordination crosslinks, forming a double crosslinked network; simultaneously, anhydrous betaine is used to construct a hydration shield within the network. This invention effectively solves the problems of needle blockage caused by excessively rapid gelation time in traditional in-situ hydrogels and excessive swelling after implantation, achieving controllable gelation time, stable osmotic pressure, and uniform internal network, exhibiting excellent clinical operability and structural stability.

Description

Technical Field

[0001] This invention relates to the field of medical polymer materials technology, specifically to an injectable polymer hydrogel and its preparation process. Background Technology

[0002] Hydrogels, due to their excellent biocompatibility and high-water-content three-dimensional network structure similar to the extracellular matrix, are widely used in medical fields such as tissue defect repair, drug delivery, and wound dressings. Among them, injectable in-situ gelling hydrogels, which can remain in a fluid state before injection and precisely fill physiological defects of any irregular shape through minimally invasive methods, have become a research focus in the field of biomaterials in recent years. This is because they can maintain a fluid state before injection and can precisely fill physiological defects of any irregular shape through minimally invasive methods, avoiding the disadvantage of traditional pre-formed hydrogels requiring open surgery. Currently, most common injectable hydrogels are based on natural high-molecular-weight polysaccharide materials, such as those that achieve dynamic cross-linking and curing through the Schiff base reaction between modified chitosan and oxidized polysaccharides.

[0003] However, in actual clinical operations and in vivo applications, traditional in-situ injectable hydrogel systems face a bottleneck where cross-linking kinetics and application performance are difficult to balance. Hydrogel systems that rely solely on chemical covalent bonds for cross-linking typically have a fast reaction rate, with extensive cross-linking and solidification occurring in the initial contact phase between the two precursor solutions. This uncontrollable gelation time can easily lead to premature coagulation of the solution in a dual-lumen syringe or static mixing syringe, causing needle blockage and severely interfering with the normal injection operation of medical personnel. If the concentration of reactive groups or solid content is artificially reduced to obtain sufficient operation time, the final hydrogel will have poor mechanical support properties, making it prone to structural loosening and loss under tissue fluid flushing.

[0004] To compensate for mechanical property deficiencies, conventional technical improvements typically involve further increasing the crosslinking density of the polymer backbone. However, in such highly crosslinked systems, changes in the internal polarity of the polymer network and the confined accumulation of numerous hydrophilic groups often disrupt the osmotic pressure balance between the hydrogel system and the surrounding physiological environment. When such hydrogels are implanted into the human body, driven by the osmotic pressure difference, surrounding bodily fluids continuously infiltrate the three-dimensional network of the gel, causing severe overswelling. This uncontrolled volume after implantation not only directly compresses surrounding normal tissues and nerves but also generates internal stress during hydration and expansion, leading to tearing of the gel's own network structure and ultimately accelerating material failure. Therefore, developing a hydrogel material that can control the initial crosslinking rate to meet clinical injection needs while ensuring post-gelation mechanical properties and preventing excessive swelling is a pressing technical problem in this field. Summary of the Invention

[0005] The technical problem solved by this invention is that traditional in-situ injection hydrogels have problems such as difficulty in controlling gelation time leading to needle blockage or loss, insufficient mechanical properties, and excessive swelling caused by osmotic pressure imbalance.

[0006] To address the above problems, the present invention provides the following technical solution:

[0007] In a first aspect, the present invention provides an injectable polymeric hydrogel, which adopts the following technical solution:

[0008] An injectable polymer hydrogel is prepared by in-situ cross-linking of liquid A and liquid B at a volume ratio of 1:1.

[0009] By weight, solution A comprises the following raw materials: 2-6 parts by weight of oxidized dextran; 2.28-11.39 parts by weight of zinc gluconate; 1.17-11.72 parts by weight of anhydrous betaine; and 100 parts by weight of 0.1 mol / L phosphate buffer solution with a pH of 7.4. Solution B comprises the following raw materials: 3-8 parts by weight of carboxymethyl chitosan; and 100 parts by weight of 0.1 mol / L phosphate buffer solution with a pH of 7.4.

[0010] By employing the above technical solution, this invention utilizes oxidized dextran and carboxymethyl chitosan to construct the main chain framework, and introduces zinc gluconate and anhydrous betaine into the system to construct a covalent and coordinated dual crosslinking network. Specifically, after the two liquids are mixed, the free aldehyde groups on the oxidized dextran molecular chain are subjected to nucleophilic attack by the primary amino groups of the carboxymethyl chitosan main chain, forming imine double bonds through dehydration, thereby establishing the preliminary three-dimensional framework of the hydrogel.

[0011] Based on this, zinc ions dissociated from zinc gluconate in the aqueous phase utilize their empty orbitals to coordinate with the carboxyl oxygen atoms and primary amino nitrogen atoms in the chitosan structure, further forming metal complex nodes. Because zinc gluconate itself has a specific complexation constant, the release of zinc ions exhibits a certain delay. This avoids clogging of the injection needle due to excessively rapid cross-linking in the initial stage of component mixing, thus providing ample time for clinical procedures.

[0012] Furthermore, anhydrous betaine, as an amphoteric inner salt, contains trimethyl quaternary ammonium cations and carboxyl anions. Due to its strong molecular polarity, this substance can bind with water molecules in the inner layer of the gel network to form a dense hydration layer. This structure not only creates steric hindrance and electrostatic shielding, slowing down the disordered aggregation of macromolecular chains in the early stages of the reaction, but also effectively regulates the osmotic pressure difference between the gel's interior and the external physiological environment, thereby inhibiting excessive swelling after implantation. In summary, the formulation system of this application enables the final hydrogel to possess advantages such as controllable gelation time, uniform internal structure, and stable osmotic pressure.

[0013] Preferably, in solution A, the molar ratio of zinc gluconate to anhydrous betaine is 1:2 to 1:4.

[0014] By employing the above technical solution and limiting the ratio of zinc gluconate to anhydrous betaine within the specified range, the main purpose is to balance the zinc ion release rate and the hydration shielding effect of betaine. On the one hand, if the proportion of betaine is too high (molar ratio below 1:4), the excessive polarity within the system will hinder the effective approach between polymer chains, resulting in lower mechanical strength of the gel after molding. On the other hand, if the zinc ion concentration is relatively high (molar ratio above 1:2), it will accelerate local cross-linking, shorten the fluid solidification time, and thus affect the smoothness of clinical injection.

[0015] Preferably, the carboxymethyl chitosan has a weight-average molecular weight (Mw) of 50,000 to 150,000 Da and a degree of substitution of 80% to 95%; the aldehyde oxidation degree of the oxidized dextran is 30% to 50%.

[0016] By adopting the above technical solutions, this application has achieved reasonable control over the molecular weight and reaction site density of the crosslinked backbone. Specifically, a substitution degree of 80%–95% ensures the water solubility of carboxymethyl chitosan under neutral physiological conditions while also exposing sufficient metal coordination sites. For oxidized dextran, an oxidation degree of 30%–50% provides covalent crosslinking points sufficient for gel network strength, while avoiding excessive damage to the polysaccharide ring structure, preserving the inherent flexibility of the main chain, and preventing brittle fracture of the hydrogel due to excessively dense crosslinking points.

[0017] Preferably, the oxidized dextran is a solid powder obtained by the oxidation reaction of dextran and sodium periodate, and the molar ratio of sodium periodate to glucose residues in the dextran structure is 0.4 to 0.8:1.

[0018] By adopting the above technical solution, and utilizing the characteristic that sodium periodate can specifically break adjacent hydroxyl groups of dextran to generate a dialdehyde structure, combined with strict control of the molar ratio of the two, the degree of aldehyde oxidation of the final product can be precisely controlled to meet the requirements of subsequent crosslinking sites.

[0019] Preferably, the preparation method of the oxidized dextran includes: dissolving dextran in deionized water to prepare a homogeneous solution, protecting it from light, and adding sodium periodate in batches at 20-30°C for 6-12 hours; after the reaction is completed, adding ethylene glycol to quench the unreacted sodium periodate, and then sequentially dialysis and vacuum freeze-drying to obtain the final product.

[0020] By adopting the above technical solution, ethylene glycol is introduced in the later stage of the reaction to consume the residual sodium periodate in the system, which can terminate the oxidation process in time. This is crucial for ensuring the consistency of oxidation degree of different batches of products. Subsequent dialysis removes small molecule impurities generated in the reaction, while vacuum freeze-drying gives oxidized dextran a porous solid form, making it exhibit excellent resolubility when preparing precursor solutions.

[0021] Secondly, the present invention provides a preparation process for an injectable polymer hydrogel, employing the following technical solution:

[0022] A process for preparing an injectable polymer hydrogel includes the following steps: adding carboxymethyl chitosan in batches to a phosphate buffer solution, stirring continuously until the material forms a homogeneous viscous liquid, and then statically degassing under vacuum to obtain solution B; adding oxidized dextran to the phosphate buffer solution and stirring until homogeneous, then sequentially adding zinc gluconate and anhydrous betaine; adjusting the temperature and continuously stirring the reaction, and after the reaction is completed, cooling to room temperature and degassing under vacuum to obtain solution A; filling solution A and solution B separately, connecting a static mixing tube, and injecting at a constant speed to mix and extrude the two component fluids, thus completing the in-situ crosslinking and gel formation process.

[0023] By employing the above technical solution, this process combines batch feeding with vacuum degassing to prepare two homogeneous precursor fluids. Particularly in the preparation stage of solution A, suitable temperature and continuous stirring allow the oxidized dextran molecular chains to fully expand in the aqueous phase, subsequently fusing with the betaine and zinc ion system to form a stable homogeneous system. In the final application stage, thanks to the helical geometry within the static mixing tube, solutions A and B undergo repeated cutting and laminar flow superposition during extrusion. This purely physical mixing method enables uniform dispersion of the two components in a short time, ensuring that covalent cross-linking and metal coordination are simultaneously initiated throughout the fluid, ultimately solidifying into a homogeneous hydrogel network.

[0024] Preferably, in the step of obtaining liquid B, the stirring temperature is 25-30°C and the stirring time is 2-4 hours; the vacuum static degassing is performed under a vacuum of -0.08MPa to -0.1MPa for 30-60 minutes.

[0025] By employing the above technical solution, maintaining a mild stirring condition of 25–30°C helps the macromolecular chains hydrate and de-entangle in the buffer solution. The static vacuum degassing operation performed after solution preparation can use the internal and external pressure difference to force the tiny bubbles mixed in the viscous solution to float and break, eliminating the risk of internal stress concentration caused by residual bubbles, thereby ensuring the continuity of the overall mechanical strength of the hydrogel.

[0026] Preferably, in the step of obtaining liquid A, the temperature adjustment is to adjust the temperature of the reactor jacket to 30-40°C, and the reaction is continuously stirred for 1-2 hours.

[0027] By adopting the above technical solution, the temperature of the reactor jacket is appropriately raised to 30-40°C. The increased thermal kinetic energy promotes the diffusion of anhydrous betaine molecules more rapidly around the oxidized dextran backbone, thus accelerating the construction of a stable hydrated layer.

[0028] Preferably, in the step of obtaining liquid A, the cooling rate after the reaction is completed is controlled at 0.5 to 1.0 °C / min; the vacuum degassing is performed at a vacuum degree of -0.08 MPa to -0.1 MPa for 25 to 35 minutes.

[0029] By employing the above technical solution, the cooling rate in the later stages of the reaction is controlled primarily to prevent localized aggregation of the polymer due to sudden temperature drops, thus maintaining the good fluidity of solution A. The subsequent vacuum degassing is also to remove air bubbles entrained during solution preparation, which helps improve the accuracy of the dual-chamber component volume ratio during actual clinical injection.

[0030] Preferably, during the in-situ crosslinking and gelation process, liquid A and liquid B are respectively filled into the two independent chambers of a medical dual-chamber syringe, connected to a static mixing tube with a spiral inner core, and extruded at a constant speed.

[0031] By employing the above technical solution, the two precursors are physically isolated using a medical dual-chamber syringe, effectively avoiding premature cross-linking during storage. During injection, the spiral core inside the static mixing tube, through hydrodynamic segmentation, causes the two liquids with a certain viscosity to alternately overlap. The operator only needs to perform a routine injection action to complete the mixing, without the need for additional electric mixing equipment, which highly conforms to the clinical in-situ injection operation habits.

[0032] This invention provides an injectable polymer hydrogel and its preparation process. It has the following beneficial effects:

[0033] 1. This invention utilizes oxidized dextran and carboxymethyl chitosan to construct Schiff base covalent crosslinks, and then introduces zinc gluconate to provide metal coordination crosslinks, forming a dual crosslinked network. Due to the specific complexation constant of zinc gluconate in the aqueous phase, the release of zinc ions exhibits a delayed effect, preventing rapid solidification of solutions A and B in the initial mixing stage. This formulation design preserves a sufficient injection time window for clinical operation, while the dual network structure also ensures the overall mechanical support performance of the hydrogel after molding.

[0034] 2. This invention utilizes anhydrous betaine added to solution A, which contains trimethyl quaternary ammonium cations and carboxyl anions, classifying it as an amphoteric internal salt. This highly polar substance can bind with water molecules within the three-dimensional network of the hydrogel, forming a dense hydration layer. This internal structure not only creates steric hindrance to slow the local aggregation of macromolecular chains in the early stages of the reaction but also directly regulates the osmotic pressure difference between the internal and external physiological environments of the hydrogel. This reduces the risk of pressure on surrounding tissues caused by the hydrogel's water absorption and swelling, thus improving post-implantation morphological stability.

[0035] 3. This invention achieves physical isolation of the reacting components by separating and preparing the precursor liquid and filling them separately into a medical dual-chamber syringe, thus avoiding premature cross-linking during storage. During the injection phase, a static mixing tube with a spiral inner core connected to the front end enables simultaneous mixing of the fluid during injection through mechanical cutting and laminar flow superposition of the internal geometry. This design allows medical personnel to achieve uniform dispersion of the two components with only routine injection procedures, ensuring that covalent and coordination reactions are initiated synchronously throughout the fluid, avoiding gel structure defects caused by uneven mixing. Attached Figure Description

[0036] Figure 1 This is a graph showing the change in the initial viscosity of liquid A over storage time in the embodiments and comparative examples of the present invention;

[0037] Figure 2 This is a graph showing the change in absorbance of liquid A at 400 nm over storage time in the embodiments and comparative examples of the present invention.

[0038] Figure 3 This is a comparative curve of the cumulative release behavior of different components of the hydrogel in simulated body fluid according to an embodiment of the present invention;

[0039] Figure 4 This is a comparison chart of initial setting time and dynamic extrusion thrust between embodiments and comparative examples of the present invention;

[0040] Figure 5 This is a comparison chart of the mechanical properties of the embodiments and comparative examples of the present invention at different gelation time points. Detailed Implementation

[0041] To enable those skilled in the art to better understand the present invention, the present invention will be further described in detail below with reference to specific embodiments. Obviously, the described embodiments are merely some embodiments of the present invention, and not all embodiments. Based on the embodiments of the present invention, all other embodiments obtained by those skilled in the art without creative effort are within the scope of protection of the present invention.

[0042] The main raw materials and reagents used in the following examples and comparative examples have the following sources and specifications. Reagents not specifically mentioned are all commercially available analytical grade or higher grade products.

[0043] Carboxymethyl chitosan, CAS No. 83512-85-0, has a weight-average molecular weight (Mw) of 50,000–150,000 Da and a degree of substitution of 80%–95%. It is a white or off-white powder, readily soluble in water. Its basic structural unit is a linear copolymer of glucosamine substituted with carboxymethyl groups and N-acetylglucosamine. The primary amino groups retained on the main chain and the newly introduced carboxyl groups provide reaction sites for Schiff base covalent crosslinking and metal coordination crosslinking, respectively, in this invention.

[0044] Dextran, CAS No. 9004-54-0, with a weight-average molecular weight (Mw) of 40,000 to 70,000 Da, is a water-soluble polysaccharide polymer composed of a large number of glucose molecules mainly linked by α-1,6-glycosidic bonds. In this invention, it serves as the basic scaffold material for preparing oxidized polysaccharide precursors with free aldehyde groups.

[0045] Zinc gluconate, CAS No. 4468-02-4, molecular formula is C 12 H 22 O 14 Zn, with a purity of not less than 99%, is a white or off-white crystalline powder. This molecule forms a complex structure through coordination interactions between gluconate anions and zinc ions, serving as a metal ion source for the delayed crosslinking network required in the system of this invention.

[0046] Anhydrous betaine, chemical name N,N,N-trimethylglycine, CAS number 107-43-7, molecular formula C5H 11 NO2, with a purity of not less than 99%, is a white crystalline powder. Its molecular structure contains both a trimethyl quaternary ammonium cation and a carboxyl anion, belonging to an amphoteric inner salt structure. It has high molecular polarity and good hydration capacity, and is used as a hydration shield and osmotic pressure regulator in the composition of this invention.

[0047] Preparation Examples 1-4:

[0048] Preparation Example 1:

[0049] This preparation example provides a process for preparing an injectable polymer hydrogel, including the following steps:

[0050] Dissolve 12 g of dextran in 100 ml of deionized water to prepare a homogeneous solution, transfer it to a light-protected reaction vessel and control the temperature at 25°C;

[0051] According to the molar ratio of sodium periodate to glucose residues in the dextran structure of 0.6:1, the corresponding mass of sodium periodate powder was weighed and slowly added to the above homogeneous solution in 4 batches. The mechanical stirring speed was controlled at 200 rpm, and the reaction was carried out continuously for 9 hours under the dark.

[0052] After the reaction was complete, 3 ml of ethylene glycol was added to the reaction solution to quench any unreacted sodium periodate, and stirring was continued for 1 hour.

[0053] The reaction solution was then placed into a dialysis bag with a molecular weight cutoff of 8000-14000 Da and dialyzed continuously in deionized water for 4 days, with the water changed 3 times a day, until no white precipitate was produced when silver nitrate was added to the dialysate.

[0054] The dialyzed solution was pre-frozen at -80°C and then treated in a vacuum freeze dryer for 60 hours to obtain a white, spongy oxidized dextran solid powder. The degree of aldehyde oxidation of the oxidized dextran was determined to be 40% by hydroxylamine hydrochloride titration. It was then sealed and protected from light as a precursor material for preparing hydrogels.

[0055] Preparation Example 2:

[0056] This preparation example provides a process for preparing an injectable polymer hydrogel, including the following steps:

[0057] Dissolve 10 g of dextran in 100 ml of deionized water to prepare a homogeneous solution, transfer it to a light-protected reaction vessel and control the temperature at 20°C;

[0058] Weigh out the corresponding mass of sodium periodate powder according to the molar ratio of sodium periodate to glucose residues in the dextran structure of 0.4:1, and slowly add it to the above homogeneous solution in 3 batches. Control the mechanical stirring speed at 150 rpm and react continuously for 6 hours under the dark.

[0059] After the reaction is complete, add 2 ml of ethylene glycol to the reaction solution to quench any unreacted sodium periodate, and continue stirring for 1 hour.

[0060] The reaction solution was then placed into a dialysis bag with a molecular weight cutoff of 8000-14000 Da and dialyzed continuously in deionized water for 3 days, with the water changed twice a day, until no white precipitate was produced when silver nitrate was added to the dialysate.

[0061] The dialyzed solution was pre-frozen at -80°C and then treated in a vacuum freeze dryer for 48 hours to obtain a white, spongy oxidized dextran solid powder. The degree of aldehyde oxidation of the oxidized dextran was determined to be 30% by hydroxylamine hydrochloride titration. It was then sealed and protected from light as a precursor material for preparing hydrogels.

[0062] Preparation Example 3:

[0063] This preparation example provides a process for preparing an injectable polymer hydrogel, including the following steps:

[0064] Dissolve 15 g of dextran in 100 ml of deionized water to prepare a homogeneous solution, transfer it to a light-proof reaction vessel and control the temperature at 30°C;

[0065] According to the molar ratio of sodium periodate to glucose residues in the dextran structure of 0.8:1, the corresponding mass of sodium periodate powder was weighed and slowly added to the above homogeneous solution in 5 batches. The mechanical stirring speed was controlled at 250 rpm, and the reaction was carried out continuously for 12 hours under the dark.

[0066] After the reaction is complete, add 5 ml of ethylene glycol to the reaction solution to quench any unreacted sodium periodate, and continue stirring for 1 hour.

[0067] The reaction solution was then placed into a dialysis bag with a molecular weight cutoff of 8000-14000 Da and dialyzed continuously in deionized water for 5 days, with the water changed 3 times a day, until no white precipitate was produced when silver nitrate was added to the dialysate.

[0068] The dialyzed solution was pre-frozen at -80°C and then treated in a vacuum freeze dryer for 72 hours to obtain a white, spongy oxidized dextran solid powder. The degree of aldehyde oxidation of the oxidized dextran was determined to be 50% by hydroxylamine hydrochloride titration. It was then sealed and protected from light as a precursor material for preparing hydrogels.

[0069] Examples 1-6:

[0070] Example 1:

[0071] This embodiment provides a preparation process for an injectable polymer hydrogel, including the following steps:

[0072] Add 100 mL of 0.1 mol / L phosphate buffer solution with pH 7.4 to a mixing tank with a constant temperature jacket. Turn on the stirrer and control the speed at 250 rpm. Add 5 g of carboxymethyl chitosan in batches at a uniform speed at 25°C. Continue stirring for 3 hours until the material forms a homogeneous viscous liquid. Then turn on the vacuum system and statically degas at -0.09 MPa for 45 minutes to obtain a transparent and bubble-free solution B. Store it in a sealed container away from light for later use.

[0073] Add 100 mL of 0.1 mol / L phosphate buffer solution with a pH of 7.4 to a clean mixing vessel, turn on the stirrer and control the speed at 200 rpm, add 4 g of oxidized dextran (aldehyde oxidation degree of 40%) obtained in Preparation Example 1, and stir to dissolve until homogeneous at room temperature; then add 6.84 g of zinc gluconate and 5.27 g of anhydrous betaine (molar ratio of zinc gluconate to anhydrous betaine is 1:3) to the vessel in sequence; adjust the jacket temperature of the reaction vessel to 35°C, and continue stirring at this temperature for 1.5 hours; after the reaction is completed, slowly cool to room temperature at a rate of 0.8°C / min, and degas under vacuum at -0.09 MPa for 25-35 minutes to obtain solution A, which exhibits low viscosity Newtonian fluid characteristics, and seal it for later use.

[0074] In-situ mixing and gel formation: The above-mentioned degassed liquid A and liquid B are filled into the two independent chambers of a medical dual-chamber syringe at a volume ratio of 1:1. By connecting a static mixing tube with a spiral inner core, the two component fluids are mixed and extruded as they pass through the mixing tube at a constant speed, thus completing the in-situ cross-linking and gel formation process.

[0075] Example 2:

[0076] This embodiment provides a preparation process for an injectable polymer hydrogel, including the following steps:

[0077] Preparation of solution B: Add 100 ml of 0.1 mol / L phosphate buffer solution with pH 7.4 to a mixing tank with a constant temperature jacket. Turn on the stirrer and control the speed at 200 rpm. Add 3 g of carboxymethyl chitosan in batches at a uniform speed at 25°C. Continue stirring for 2 hours until the material forms a homogeneous viscous liquid. Then turn on the vacuum system and statically degas at -0.08 MPa for 30 minutes to obtain a transparent and bubble-free solution B. Store it in a sealed container away from light for later use.

[0078] Add 100 mL of 0.1 mol / L phosphate buffer solution with a pH of 7.4 to a clean mixing vessel, turn on the stirrer and control the speed at 150 rpm, add 2 g of oxidized dextran (aldehyde oxidation degree of 30%) obtained in Preparation Example 2, and stir to dissolve until homogeneous at room temperature; then add 2.28 g of zinc gluconate and 1.17 g of anhydrous betaine (molar ratio of zinc gluconate to anhydrous betaine is 1:2) to the vessel in sequence; adjust the jacket temperature of the reaction vessel to 30°C, and continue stirring at this temperature for 1 hour; after the reaction is completed, slowly cool to room temperature at a rate of 0.5°C / min, and degas under vacuum at -0.08 MPa for 25-35 minutes to obtain solution A, which exhibits low viscosity Newtonian fluid characteristics, and seal it for later use.

[0079] In-situ mixing and gel formation: The above-mentioned degassed liquid A and liquid B are filled into the two independent chambers of a medical dual-chamber syringe at a volume ratio of 1:1. By connecting a static mixing tube with a spiral inner core, the two component fluids are mixed and extruded as they pass through the mixing tube at a constant speed, thus completing the in-situ cross-linking and gel formation process.

[0080] Example 3:

[0081] This embodiment provides a preparation process for an injectable polymer hydrogel, including the following steps:

[0082] Add 100 mL of 0.1 mol / L phosphate buffer solution with pH 7.4 to a mixing tank with a constant temperature jacket. Turn on the stirrer and control the speed at 300 rpm. Add 7 g of carboxymethyl chitosan in batches at a uniform speed at 30°C. Continue stirring for 4 hours until the material forms a homogeneous viscous liquid. Then turn on the vacuum system and statically degas at -0.1 MPa for 60 minutes to obtain a transparent and bubble-free solution B. Store it in a sealed container away from light for later use.

[0083] Add 100 mL of 0.1 mol / L phosphate buffer solution with a pH of 7.4 to a clean mixing vessel, turn on the stirrer and control the speed at 250 rpm, add 6 g of oxidized dextran (aldehyde oxidation degree of 50%) obtained in Preparation Example 3, and stir to dissolve until homogeneous at room temperature; then add 11.39 g of zinc gluconate and 11.72 g of anhydrous betaine (molar ratio of zinc gluconate to anhydrous betaine is 1:4) to the vessel in sequence; adjust the jacket temperature of the reaction vessel to 40 °C, and continue stirring at this temperature for 2 hours; after the reaction is completed, slowly cool to room temperature at a rate of 1.0 °C / min, and degas under vacuum at -0.1 MPa for 25-35 minutes to obtain solution A, which exhibits low viscosity Newtonian fluid characteristics, and seal for later use.

[0084] In-situ mixing and gel formation: The above-mentioned degassed liquid A and liquid B are filled into the two independent chambers of a medical dual-chamber syringe at a volume ratio of 1:1. By connecting a static mixing tube with a spiral inner core, the two component fluids are mixed and extruded as they pass through the mixing tube at a constant speed, thus completing the in-situ cross-linking and gel formation process.

[0085] Example 4:

[0086] This embodiment provides a preparation process for an injectable polymer hydrogel, including the following steps:

[0087] Add 100 mL of 0.1 mol / L phosphate buffer solution with pH 7.4 to a mixing tank with a constant temperature jacket. Turn on the stirrer and control the speed at 300 rpm. Add 8 g of carboxymethyl chitosan in batches at a uniform speed at 30°C. Continue stirring for 4 hours until the material forms a homogeneous viscous liquid. Then turn on the vacuum system and statically degas at -0.1 MPa for 60 minutes to obtain a transparent and bubble-free solution B. Store it in a sealed container away from light for later use.

[0088] Add 100 mL of 0.1 mol / L phosphate buffer solution with a pH of 7.4 to a clean mixing vessel, turn on the stirrer and control the speed at 250 rpm, add 5 g of oxidized dextran (aldehyde oxidation degree of 50%) obtained in Preparation Example 3, and stir to dissolve until homogeneous at room temperature; then add 9.11 g of zinc gluconate and 7.03 g of anhydrous betaine (molar ratio of 1:3) to the vessel in sequence; adjust the jacket temperature of the reaction vessel to 35°C, and continue stirring at this temperature for 1.5 hours; after the reaction is completed, slowly cool to room temperature at a rate of 0.8°C / min, and degas under vacuum at -0.09 MPa for 25-35 minutes to obtain solution A, which exhibits low viscosity Newtonian fluid characteristics, and seal for later use.

[0089] In-situ mixing and gel formation: The above-mentioned degassed liquid A and liquid B are filled into the two independent chambers of a medical dual-chamber syringe at a volume ratio of 1:1. By connecting a static mixing tube with a spiral inner core, the two component fluids are mixed and extruded as they pass through the mixing tube at a constant speed, thus completing the in-situ cross-linking and gel formation process.

[0090] Comparative Examples 1-5:

[0091] Comparative Example 1:

[0092] Compared with Example 1, the difference is that zinc gluconate and anhydrous betaine are not added in the preparation of solution A, but only phosphate buffer solution of oxidized dextran is added, and the rest are the same.

[0093] Comparative Example 2:

[0094] Compared with Example 1, the difference is that anhydrous betaine is not added in the preparation of solution A, only 6.84 grams of zinc gluconate is added, and the rest are the same.

[0095] Comparative Example 3:

[0096] Compared with Example 1, the difference is that in the preparation of solution A, anhydrous betaine is replaced with L-arginine containing free primary amine in equimolar amounts. All other aspects are the same.

[0097] Comparative Example 4:

[0098] Compared with Example 1, the difference is that the amount of anhydrous betaine added in the preparation of solution A is adjusted to 1.76 grams, so that the molar ratio of zinc gluconate to anhydrous betaine is 1:1, and all other aspects are the same.

[0099] Comparative Example 5:

[0100] Compared with Example 1, the difference is that zinc gluconate is replaced with zinc chloride in equimolar amounts in the preparation of solution A, while the rest are the same.

[0101] Test Examples 1-4:

[0102] Test Example 1: Long-term storage stability test of liquid A.

[0103] In this test example, the long-term storage stability of the hydrogel precursor solution provided by this invention was mainly investigated. The specific operating steps are as follows:

[0104] Take 15 ml of each of the solutions A prepared in Examples 1 to 4 and Comparative Example 3, and dispense them into 20 ml brown glass sample bottles with polytetrafluoroethylene liner caps. After filling with an appropriate amount of nitrogen to purge the air from the top, tighten and seal the bottles.

[0105] Transfer the sealed sample vials to a constant temperature incubator set at 25°C and store them in the dark.

[0106] On days 0, 15, 30, 60 and 90 of the storage period, each sample vial was opened and an equal amount of test sample was drawn using a pipette for the determination of rheological and optical properties.

[0107] During the test, a constant shear rate of 10 s was used. -1 The initial viscosity of each sample was measured using a rotational rheometer at 25°C to assess whether spontaneous cross-linking of the polymer network inside the fluid occurred.

[0108] A portion of the sample was extracted and diluted 10-fold with 0.1 mol / L phosphate buffer. The absorbance at 400 nm was measured using a UV-Vis spectrophotometer to quantitatively assess the browning caused by Schiff base side reactions within the solution system.

[0109] Table 1. Changes in initial viscosity and absorbance of solution A in each example and comparative example at different storage times.

[0110] Test object Test metrics 0 days 15 days 30 days 60 days 90 days Example 1 Initial viscosity (mPa·s) 21.3 21.5 22.1 23.0 23.8 Example 1 Absorbance (A400) 0.021 0.025 0.033 0.038 0.042 Example 2 Initial viscosity (mPa·s) 18.2 18.7 19.1 19.5 20.4 Example 2 Absorbance(A400) 0.015 0.018 0.022 0.026 0.031 Example 3 Initial viscosity (mPa·s) 28.5 29.1 30.2 31.5 32.7 Example 3 Absorbance(A400) 0.032 0.037 0.045 0.051 0.059 Example 4 Initial viscosity (mPa·s) 25.1 25.4 26.0 26.8 28.1 Example 4 Absorbance(A400) 0.027 0.031 0.039 0.044 0.048 Comparative Example 3 Initial viscosity (mPa·s) 23.4 68.7 185.3 412.9 895.2 Comparative Example 3 Absorbance(A400) 0.025 0.412 0.854 1.345 1.892

[0111] Figure 1 The results show the change in the initial viscosity of liquid A over storage time. Figure 2The figure shows the test results of the absorbance of liquid A at 400 nm as a function of storage time. The horizontal axis corresponds to different test samples, namely Example 1, Example 2, Example 3, Example 4, and Comparative Example 3. Each sample area contains five bars of different gray levels, which, from lightest to darkest, represent the test data obtained on days 0, 15, 30, 60, and 90 of static storage at 25°C. Figure 1 The vertical axis reflects the initial viscosity value of the fluid sample. Figure 2 The vertical axis represents the absorbance, which characterizes the degree of browning in the system.

[0112] Summary: Based on Table 1 and Figure 1 , Figure 2 The test data allows observation of the system's state evolution in an aqueous environment. In routine stability tests of hydrogel precursor solutions, free aldehyde groups on the polysaccharide backbone are prone to physicochemical changes in aqueous media, thus affecting the material's shelf life.

[0113] In this embodiment, during the 90-day isothermal monitoring period at 25°C, the viscosity of solution A from Examples 1 to 4 showed only minor fluctuations from the original baseline, and the increase in absorbance reflecting the accumulation of by-reaction products remained at a low level, maintaining homogeneous fluidity. This test result verifies the structural advantage of introducing anhydrous betaine as a steric masking agent in the formulation. Because anhydrous betaine, as a quaternary ammonium inner salt, lacks primary or secondary amine groups capable of participating in nucleophilic addition reactions, the free aldehyde groups on the oxidized dextran chain are in a relatively stable chemical environment.

[0114] Combined with the data from Comparative Example 3, the influence of free amine substances on the system stability was further confirmed. After replacing the masking molecule with an equimolar amount of L-arginine containing a free primary amine, the viscosity and absorbance of the Comparative Example 3 sample showed a continuous upward trend over time. This indicates that spontaneous Schiff base crosslinking and browning caused by the enrichment of conjugated double bonds occurred within the reaction system at room temperature, proving that conventional amino acid-based cosolvents or osmotic pressure regulators can introduce reaction sites that trigger premature gelation.

[0115] The above comparison results show that the present invention eliminates the risk of early product failure caused by spontaneous cross-linking of the basic components, and provides objective data support for maintaining the effectiveness of materials during large-scale production and logistics transportation.

[0116] Test Example 2: Permeation-diffusion triggering and ligand exchange kinetics test.

[0117] Using a medical dual-chamber syringe, 1 ml each of solutions A and B prepared in Examples 1 and 3 were simultaneously injected into a polytetrafluoroethylene cylindrical mold with a diameter of 10 mm and a height of 2 mm with constant thrust. The mold was then placed in a constant temperature oven at 37°C and left to stand for 30 minutes to obtain a solidified initial gel sample.

[0118] The formed hydrogel samples were carefully removed and placed into dialysis bags with a molecular weight cutoff of 3500 Da. 5 ml of simulated body fluid (SBF) with a pH of 7.4 was added to the bags, and then the two ends of the dialysis bags were sealed with sealing clips.

[0119] Immerse the sealed dialysis bag in a sealed wide-mouth bottle containing 100 ml of the same SBF solution, and transfer it to a constant temperature air bath shaker at 37°C. Adjust the shaking frequency to 60 rpm to simulate the dynamic fluid microenvironment in the interstitial spaces of human tissue.

[0120] At the set time points (0.5 hours, 2 hours, 6 hours, 12 hours, 24 hours, 48 ​​hours and 72 hours), 2 ml of external release medium was drawn from the wide-mouth bottle, and 2 ml of fresh SBF solution preheated to 37°C was immediately added to the bottle to maintain the total volume of the release system.

[0121] The concentration of betaine in the sampled medium at each time point was determined by high performance liquid chromatography (HPLC), while the concentration of free zinc ions in the sample was determined by inductively coupled plasma optical emission spectrometry (ICP-OES). The cumulative percentage of substances released at different time points was calculated by combining the volume of the release medium with the total amount of drug.

[0122] Table 2. Cumulative release rates of betaine and zinc ions from hydrogels in simulated body fluids for each embodiment.

[0123] Test object Monitoring materials 0.5h 2h 6h 12h 24h 48h 72h Example 1 betaine (%) 18.2 41.7 63.4 79.8 88.3 91.5 92.7 Example 1 Zinc ions (%) 1.8 3.5 5.2 6.7 9.4 11.2 12.8 Example 3 betaine (%) 22.4 46.1 67.9 83.2 90.7 93.8 95.1 Example 3 Zinc ions (%) 2.1 4.3 6.1 8.2 10.5 11.9 13.5

[0124] Figure 3 This is a comparative curve of the cumulative release behavior of different components of the hydrogel in simulated body fluid according to embodiments of the present invention. The horizontal axis represents the time after the gel sample is immersed in the release medium, and the vertical axis represents the cumulative release percentage of the target substance relative to the theoretical total. Dashed lines and hollow markers represent the release trajectory of the osmotic pressure regulator betaine, while solid lines and solid markers represent the release trajectory of the metal crosslinking agent zinc ions. Different shaped markers (circles and squares) correspond to the test samples of Examples 1 and 3, respectively.

[0125] In summary, the data in Table 2 clearly demonstrate the ability to track the component migration trajectories of the hydrogel network under simulated physiological conditions. In in vitro drug release assessment systems, the diffusion behavior of small molecules is often constrained by the crosslinking density of the polymer network and the local osmotic pressure gradient.

[0126] In this embodiment, during the initial stage after immersion in simulated body fluid, the anhydrous betaine in the gel samples of Examples 1 and 3 exhibited a continuous outward diffusion process, with its cumulative release rate exceeding 88% within 24 hours. This high release rate is attributed to the small hydrodynamic radius of betaine and the concentration difference formed inside and outside the gel, allowing it to smoothly pass through the loose primary Schiff base network. In contrast to the loss of betaine, the release of free zinc ions, which also belong to the small molecule category, remained consistently around 13% during the 72-hour monitoring period.

[0127] As betaine diffuses extensively into the surrounding medium, the hydration shielding layer initially surrounding zinc gluconate gradually disintegrates, disrupting the local chemical equilibrium within the system. The zinc ions, no longer protected by steric hindrance, are not lost with bodily fluids but are instead captured by adjacent amino and carboxyl groups on the carboxymethyl chitosan backbone. The polydentate ligands exhibit a stronger thermodynamic affinity for transition metal ions; this competitive ligand exchange process immobilizes zinc ions in situ between polymer chains, completing the delayed assembly of the secondary coordination crosslinking network.

[0128] The test results confirm the existence of the synergistic effect of osmosis-driven and ligand exchange from a kinetic perspective, indicating that the present invention achieves spontaneous enhancement of macroscopic gel strength by adjusting the spatiotemporal distribution of molecules, providing a basis for the service of injectable hydrogels in complex tissue defects.

[0129] Test Example 3: Rheological behavior and injection extrusion thrust test.

[0130] This test case aims to investigate the evolution of rheological parameters of the hydrogel precursor solution during the initial mixing stage and to evaluate its impact on the actual injection thrust of medical devices. The specific test steps are as follows:

[0131] The A and B solutions prepared in Examples 1 to 4, as well as Comparative Examples 2, 4 and 5, were selected as test samples.

[0132] A rotational rheometer equipped with a 20 mm parallel plate test fixture was used, and the temperature control system of the test platform was set to 37°C to simulate the body temperature environment. Using a micropipette, 0.5 ml of solution A and 0.5 ml of solution B were respectively added to the rheometer test platform for instantaneous mixing, and then the upper pressure plate was lowered to a test gap of 1 mm.

[0133] The strain parameters of the rheometer were set to 1% and the angular frequency to 10 rad / s, and the time scan mode was started. The time point at which the storage modulus (G') and loss modulus (G'') curves intersected in the continuously acquired data was extracted and recorded as the initial solidification time of the system.

[0134] Fill equal volumes of the matching solution A and solution B into a 5 ml medical double-chamber syringe, and connect a standard static mixing tube containing a spiral blade and an 18G medical injection needle to the end of the syringe.

[0135] The assembled syringe was fixed to the center of the compression clamp of the universal testing machine, and the compression rate of the testing machine was set to a constant 1 mm / s. The maximum extrusion thrust value captured by the sensor within the 20 mm displacement stroke of the push rod was recorded, and the extrusion state of the fluid at the needle tip was recorded simultaneously.

[0136] Table 3. Initial setting time and dynamic extrusion thrust test data for each embodiment and comparative example.

[0137] Test object Initial setting time (s) Maximum compressive thrust (N) Mixing tube and needle extrusion state Example 1 14.3 8.2 Smooth, continuous extrusion, with no visible gel lumps. Example 2 12.1 7.6 Smooth, continuous extrusion, with no visible gel lumps. Example 3 16.5 9.1 Smooth and continuous extrusion, resulting in a uniform and dense fluid. Example 4 15.2 8.7 Smooth and continuous extrusion, resulting in a uniform and dense fluid. Comparative Example 2 1.4 86.5 The syringe locked, and a hard clot appeared at the base of the mixing tube. Comparative Example 4 3.8 45.2 Extrusion was obstructed and intermittent, containing microgel particles. Comparative Example 5 1.2 91.3 Syringe locked, tubing blocked

[0138] Figure 4 This is a dual-axis graph comparing the initial setting time and dynamic extrusion thrust test results of various embodiments and comparative examples of the present invention. The horizontal axis corresponds to each group of test samples; the left vertical axis corresponds to the initial setting time measured by the rheometer, displayed as a gray bar chart; the right vertical axis corresponds to the maximum extrusion thrust recorded by the universal testing machine, displayed as a black broken line and scatter plot.

[0139] Summary: Based on Table 3 and Figure 4 The test data allows us to observe the influence of different components in the formulation on the early crosslinking kinetics of the hydrogel. In the injection of injectable medical polymers, the time window during which the fluid undergoes a phase transition through a static mixing tube is directly related to the feasibility of actual clinical drug administration.

[0140] In this embodiment, the initial setting time of Examples 1 to 4 was maintained in the range of 12 to 17 seconds. During the instrument injection process, this crosslinking delay provided a suitable flow window for the physical homogenization of liquid A and liquid B in the mixing tube, ensuring that the resistance value remained below 10N throughout the injection stroke, and the extrudate exhibited a homogeneous and dense state. The above-mentioned rheological behavior corresponds to the steric hindrance mechanism at the molecular level. The hydrated layer of the betaine groups acted as a physical barrier in the initial contact stage, delaying the excessively rapid crosslinking of active groups between polymer backbones.

[0141] The test results from the comparative examples further illustrate the role of this shielding mechanism in the formulation. Comparative Example 2 only added zinc gluconate to the system without using anhydrous betaine, while Comparative Example 5 directly used zinc chloride as a crosslinking agent. Both samples lacking hydration shielding exhibited modulus crossover within 2 seconds of mixing, exceeding the operable range of conventional manual injection.

[0142] In compression tests on the testing machine, the thrust of these samples rapidly increased to over 85N, leading to physical blockage of the tubing and the formation of clumps inside the mixing chamber that were difficult to shear and break. Even with the addition of a small amount of anhydrous betaine in Comparative Example 4, due to its low molar ratio with zinc ions, it failed to form sufficient coating, and the initial setting time was only extended to 3.8 seconds. The injection process was still accompanied by a feeling of resistance and the extrusion of microgel particles. The comparison of the above physical quantities indicates that by introducing an appropriate amount of betaine to construct a transient shielding layer, the timing of fluid mixing homogenization and network cross-linking can be adjusted, thereby improving the early injectability of the dual-network hydrogel.

[0143] Test Example 4: Stepped gelation and final mechanical strengthening test.

[0144] This test case primarily investigates the evolution of macroscopic mechanical parameters of hydrogels at different stages of curing and evaluates the energy dissipation level of the material under dynamic stress conditions. The specific test procedures are as follows:

[0145] Examples 1 to 4 and Comparative Example 1 (without zinc ion crosslinking agent) were selected as test samples. The corresponding solutions A and B for each group were filled into a medical double-lumen syringe and simultaneously injected into a cylindrical polytetrafluoroethylene mold with an inner diameter of 10 mm and a height of 10 mm through the static mixing tube at the front end.

[0146] To prevent excessive evaporation of moisture during gel solidification, the mold containing the mixed precursor solution was placed in a petri dish with a moistened filter paper at the bottom, and then transferred to a 37°C incubator, covered, and left to incubate.

[0147] At the designated incubation time points (10 minutes and 120 minutes of gelation), the gel cylinders were completely removed using the accompanying demolding device. The actual geometric dimensions of each sample were measured using vernier calipers, and the corresponding values ​​were recorded for accurate calculation of the cross-sectional area under stress in subsequent operations.

[0148] The molded gel specimen was placed in the center of the lower platen of a universal testing machine, and an unconfined uniaxial compression test was conducted at room temperature. During the test, the crosshead compression rate was kept constant at 2 mm / min. The maximum stress value at which the specimen was compressed to the point of macroscopic rupture was recorded as the fracture stress, and the slope of the linear segment of the test curve in the 0–10% strain range was extracted to characterize the compressive modulus of the sample.

[0149] For the sample group that completed 120 minutes of incubation, a cyclic compression test procedure was further added. The maximum compressive strain was set to 50%, and 10 consecutive loading-unloading cycles were performed using the same loading rate as before. The hysteresis loop area formed by the envelopes of the loading and unloading segments in the first test curve was extracted, and the energy dissipation rate per unit volume of material was calculated based on this.

[0150] Table 4. Mechanical parameters of gels from different gelation stages in each example and comparative example.

[0151] Test object Compression modulus -10min (kPa) Compression modulus -120 min (kPa) Fracture stress -120 min (kPa) <![CDATA[Energy dissipation rate - 120 min (kJ / m 3 )]]> Example 1 15.2 68.4 165.3 54.2 Example 2 12.8 62.1 151.8 49.7 Example 3 16.7 74.5 188.2 62.4 Example 4 14.9 70.2 172.6 58.1 Comparative Example 1 16.1 19.3 42.5 11.4

[0152] Figure 5 The graph shows the mechanical properties of the embodiments and comparative examples of the present invention at different gelation time points. The horizontal axis in the graph indicates different test sample groups; the left vertical axis corresponds to the compressive modulus value of the gel sample, and the light gray and dark gray bars inside represent the data obtained by the sample after incubation for 10 minutes and 120 minutes, respectively; the right vertical axis corresponds to the energy dissipation rate of the sample under cyclic compression test, represented by a black broken line with diamond markings and scatter dots.

[0153] Summary: Based on Table 4 and Figure 5 The experimental data reflect the crosslinking evolution of the polymer network in the gel system at different molding time points. Typically, implantable medical materials need to maintain appropriate deformation capacity in the initial injection stage to conform to irregular wound surfaces, and after curing, they need to possess corresponding load-bearing properties to cope with the physiological load of surrounding tissues.

[0154] In this embodiment, the compressive modulus of the samples from Examples 1 to 4 after 10 minutes of curing was concentrated in the range of 12.8 to 16.7 kPa, which is similar to the data level of Comparative Example 1 without the introduction of zinc ions. The mechanical support at this stage is provided by the primary covalent network, and the material as a whole exhibits a relatively soft texture. As the internal osmotic pressure regulator diffuses and is lost to the peripheral body fluid, the local coordination balance between molecules changes, and the free zinc ions are gradually captured and bound by the carboxyl and amino groups on the polysaccharide backbone.

[0155] The migration of components over time led to a secondary strengthening of the material's macroscopic mechanics. When the testing time was extended to 120 minutes, the compressive modulus of each example group increased in a stepwise manner, and the fracture stress and energy dissipation rate, representing toughness indicators, increased accordingly. This increase in mechanical parameters is related to the completion of the assembly of the secondary metal coordination network; the newly formed metal-heteroatom coordination bonds provide a pathway for dissipating mechanical energy during compressive deformation of the material.

[0156] Under continuous compression or cyclic loading, some non-covalent coordination networks in the gel sample undergo reversible dissociation and absorb the mechanical deformation energy input from the external environment, thus limiting the propagation of microcracks in the main chain. Referring to the test results of Comparative Example 1, due to the lack of this type of dynamic dissipation mechanism in the system, its compressive modulus at 120 minutes is 19.3 kPa, and its energy dissipation rate is 11.4 kJ / m³. 3The baseline level shows brittle fracture characteristics under pressure. Comparing the above dynamic test results, the composition of this invention utilizes the time difference in component release to regulate the formation rhythm of the gel network, alleviating the contradiction between the initial fluidity and later mechanical strength of medical gel injection, and providing technical support for maintaining long-term structural stability of the material in local tissues.

Claims

1. An injectable polymeric hydrogel, characterized in that, It is prepared by in-situ crosslinking of liquid A and liquid B at a volume ratio of 1:1; By weight, liquid A comprises the following raw materials: 2-6 parts by weight of oxidized dextran; Zinc gluconate 2.28–11.39 parts by weight; Anhydrous betaine 1.17–11.72 parts by weight; 100 parts by weight of 0.1 mol / L phosphate buffer solution with a pH of 7.4; The B liquid comprises the following raw materials: 3-8 parts by weight of carboxymethyl chitosan; 100 parts by weight of 0.1 mol / L phosphate buffer solution with a pH of 7.

4.

2. The injectable polymeric hydrogel according to claim 1, characterized in that, In solution A, the molar ratio of zinc gluconate to anhydrous betaine is 1:2 to 1:

4.

3. The injectable polymeric hydrogel according to claim 1, characterized in that, The carboxymethyl chitosan has a weight-average molecular weight (Mw) of 50,000–150,000 Da and a degree of substitution of 80%–95%. The degree of aldehyde oxidation of the oxidized dextran is 30% to 50%.

4. The injectable polymer hydrogel according to claim 1, characterized in that, The oxidized dextran is a solid powder obtained by the oxidation reaction of dextran and sodium periodate, wherein the molar ratio of sodium periodate to glucose residues in the dextran structure is (0.4-0.8):

1.

5. The injectable polymeric hydrogel according to claim 4, characterized in that, The method for preparing the oxidized dextran includes: Dextran was dissolved in deionized water to prepare a homogeneous solution. Sodium periodate was added in batches at 20–30°C and reacted for 6–12 hours in the dark. After the reaction is complete, ethylene glycol is added to quench any unreacted sodium periodate, followed by dialysis and vacuum freeze-drying to obtain the final product.

6. A preparation process for an injectable polymeric hydrogel as described in any one of claims 1-5, characterized in that, Includes the following steps: The carboxymethyl chitosan was added to the phosphate buffer in batches and stirred continuously until the material formed a homogeneous viscous liquid. After vacuum static degassing, solution B was obtained. The oxidized dextran was added to the phosphate buffer solution and stirred until homogeneous. The zinc gluconate and anhydrous betaine were added sequentially. The temperature was adjusted and the reaction was stirred continuously. After the reaction was completed, the temperature was lowered to room temperature and the mixture was degassed under vacuum to obtain solution A. Liquid A and liquid B are filled separately, connected to a static mixing tube, and injected at a constant speed to mix and extrude the two component fluids, thus completing the in-situ crosslinking and gelation process.

7. The preparation process according to claim 6, characterized in that, In the step of obtaining liquid B, the stirring temperature is 25-30℃ and the stirring time is 2-4 hours; the vacuum static degassing is performed under a vacuum of -0.08MPa to -0.1MPa for 30-60 minutes.

8. The preparation process according to claim 6, characterized in that, In the step of obtaining liquid A, the temperature adjustment is to adjust the temperature of the reactor jacket to 30-40°C, and the reaction is continuously stirred for 1-2 hours.

9. The preparation process according to claim 6, characterized in that, In the step of obtaining liquid A, the cooling rate after the reaction is completed is controlled at 0.5 to 1.0 °C / min; the vacuum degassing is carried out at a vacuum degree of -0.08 MPa to -0.1 MPa for 25 to 35 minutes.

10. The preparation process according to claim 6, characterized in that, In the process of completing the in-situ crosslinking and gel formation, liquid A and liquid B are respectively filled into the two independent chambers of a medical dual-chamber syringe, connected to a static mixing tube with a spiral inner core, and extruded at a constant speed.

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