Preparation of an in-situ gelling anti-swelling enhanced tissue adhesive material

Abstract

The present invention relates to the preparation of an in-situ gelling anti-swelling enhanced tissue adhesive. Specifically, the present invention uses a multi-arm polyethylene glycol solution with terminal modified succinimide active ester and a polymer solution with modified catechol groups and containing multiple primary amino groups as the main components of the gel. By mixing the solutions for application or injection, the terminal active ester groups of the polyethylene glycol and the amino groups on the polymer spontaneously undergo nucleophilic substitution reactions to form a gel in situ that adheres to the tissue surface. After adhering to the tissue surface, the catechol in the gel can resist swelling due to water absorption by exerting a self-oxidative cross-linking effect, and the oxidized catechol groups can also undergo multiple bonding with the amino groups, thiol groups, etc. on the tissue protein to achieve an enhanced tissue adhesion effect.

Description

Technical Field

[0001] The invention belongs to the technical field of biomaterials, and in particular, relates to the preparation of an in-situ gelling anti-swelling enhanced tissue adhesive. Background Art

[0002] Tissue adhesives have high application value in clinical medicine. They are usually used in many fields such as mucosal tissue adhesion, sealing, drug delivery and release. Although various adhesive formulations with high biosafety for soft tissues have been developed, their clinical use is still limited by insufficient adhesion and large swelling. For example, the fibrin adhesive commonly used in clinical practice has an adhesion strength of only about 15 KPa to tissues, and there is also a risk of viral transmission (such as parvovirus B19, hepatitis and HIV).

[0003] As a highly biosafe polymer, polyethylene glycol (PEG) has been developed into a variety of derivative molecules for the preparation of tissue adhesives, such as multi-arm PEG active ester and multi-arm PEG thiol two-component gel (CoSeal), multi-arm PEG active ester and trilysine two-component gel (DuraSeal) and PEG acrylic acid-based light-activated gel (FocalSeal).

[0004] Although the tissue adhesion strength of this type of gel is improved compared to fibrin glue, it is still insufficient (20-30KPa); in addition, since the PEG molecules in these synthetic adhesives have strong water absorption, they absorb the surrounding tissue fluid after implantation, causing the volume to increase, which may compress the surrounding nerves and other tissues. At the same time, water absorption will cause the mechanical properties and bonding strength of the gel to drop sharply, affecting its effectiveness. Therefore, it is very worthwhile to develop an anti-swelling gel adhesive with strong adhesion in complex chemical and mechanical environments.

[0005] Marine organisms mussels can secrete adhesive proteins that allow them to firmly adhere to different surfaces in a humid environment. This is because they contain an unusual amino acid, L-3,4-dihydroxy-L-phenylalanine (DOPA). The catechol groups contained in them can form non-covalent interactions such as hydrogen bonds, cation-π, π-π stacking, etc. with hydroxyl groups, cationic groups, benzene rings, etc., and at the same time, they are converted into quinone structures under oxidation. They are highly reactive and can undergo Michael addition, Schiff base, disproportionate coupling and other reactions with various groups including thiols, amines and quinones. In recent years, there have been many reports on hydrogel adhesives inspired by mussels, such as Lee et al. reported a tissue adhesive formed by catechol functionalized chitosan (Adv.Funct.Mater.2019,29,1900495) and a catechol functionalized four-arm polyethylene glycol tissue adhesive (Langmuir 2010,26,3790-3793); Neto et al. reported a hydrogel adhesive prepared by catechol functionalized hyaluronic acid (Small 2014,10, 2459-2469). However, the self-oxidation rate of these gels is very slow, and it is impossible to quickly form gels through cross-linking between catechols, which takes several days. Therefore, such adhesives usually need to add a large amount of toxic oxidants, such as sodium periodate, hydrogen peroxide / horseradish peroxidase, etc., which greatly limits its clinical application.

[0006] The present invention uses a water-soluble polymer modified with catechol functional groups and containing multiple primary amino groups as a multifunctional crosslinking agent to perform an in-situ gelation reaction with multi-arm polyethylene glycol modified with active esters at the end to form a polymer network and seal the tissue. Figure 1 As shown, the catechol groups in the polymer network are slowly oxidized in a neutral pH environment to form highly reactive quinones. The self-crosslinking reaction between quinones and the reaction between quinones and primary amino groups make the gel network tighter, thereby continuously resisting the entry of water molecules to avoid swelling of the gel. At the same time, it can also form multiple bonds with amino groups, thiol groups, etc. on tissue proteins to enhance tissue adhesion.

[0007] In summary, the present invention designs a high-adhesion hydrogel adhesive that can quickly gel and resist swelling by applying or injecting, utilizing the rapid amidation reaction of active esters and primary amino groups, and the cross-linking reaction based on catechol groups. It is expected to improve the problems of poor tissue adhesion and high swelling of existing gel adhesives in clinical practice. Summary of the invention

[0008] The purpose of the present invention is to provide a medical tissue adhesive material which is convenient for clinical use, has strong adhesion, does not swell, has strong tissue adhesion ability and excellent biocompatibility.

[0009] The first aspect of the present invention provides an in-situ gelling anti-swelling enhanced tissue adhesive, the hydrogel comprising:

[0010] (1) Multi-arm polyethylene glycol with terminally modified active esters;

[0011] (2) polymers grafted with catechol groups and containing multiple primary amino groups;

[0012] (3) gelling buffer solution;

[0013] Wherein, the four-arm polyethylene glycol with terminal modification of succinimide active ester group has one or more of the following characteristics:

[0014] (1) The number of arms of polyethylene glycol is ≥2;

[0015] (2) Weight average molecular weight of 5000 to 40000;

[0016] (3) The polydispersity index (PDI) of a single polyethylene glycol polymer (PEG) is 1 to 1.1.

[0017] In another preferred embodiment, the number of arms of the four-arm polyethylene glycol terminally modified with a succinimide active ester group is 2 to 8, with 4 being the most preferred.

[0018] In another preferred embodiment, the molecular weight of the four-arm polyethylene glycol with terminal modified succinimide active ester group is 5000-20000, with the best being 10000.

[0019] In another preferred embodiment, the four-arm polyethylene glycol with a terminal modified succinimide active ester group is four-arm polyethylene glycol succinimidyl carbonate (SC), four-arm polyethylene glycol succinimidyl acetate (SCM), four-arm polyethylene glycol succinimidyl glutarate (SG), four-arm polyethylene glycol succinimidyl succinate (SS), etc.; preferably, it is four-arm polyethylene glycol succinimidyl carbonate (SC) and four-arm polyethylene glycol succinimidyl glutarate (SG), and the best is four-arm polyethylene glycol succinimidyl carbonate (SC).

[0020] Wherein, the polymer grafted with catechol groups and containing multiple primary amino groups has one or more of the following characteristics:

[0021] (1) The polymer has a weight average molecular weight of 1,000 to 100,000, and each repeating unit contains a primary amino group, which may be one or more of polyallylamine, polyethyleneimine, polylysine, polyarginine, and chitosan;

[0022] (2) Usually catechol derivatives or amino-containing polymers are used for grafting modification, and the grafting rate is between 5% and 40%, preferably ≥15%;

[0023] (3) The polymer derivative has good water solubility and can be degraded in vivo, and the products before and after degradation are non-toxic.

[0024] In another preferred embodiment, the polymer containing multiple primary amino groups can be polylysine, polyarginine, polyethyleneimine, carboxymethyl chitosan, etc., among which polylysine and polyethyleneimine are preferred, and polylysine is the most preferred.

[0025] In another preferred case, the modified catechol method can be obtained by utilizing the terminal carboxyl groups of 3,4-dihydroxybenzoic acid, 3,4-dihydroxyphenylacetic acid, 3,4-dihydroxyphenylpropionic acid, and 3,4-dihydroxyphenylbutyric acid and the amino groups on the polymer, and reacting them under the action of a condensing agent carbodiimide (EDC) / N-hydroxysuccinimide (NHS) or 4-(4,6-dimethoxytriazine-2-yl)-4-methylmorpholine hydrochloride (DMTMM), wherein 3,4-dihydroxyphenylpropionic acid is preferably synthesized under the action of EDC / NHS, the molar ratio of the catechol derivative to the polymer repeating unit is between 2:1 and 1:2, and the molar ratio of EDC and NHS or DMTMM to the polymer dissolving unit is between 2:1 and 1:2.

[0026] In another preferred case, the modified catechol modification method can be obtained by spontaneously reacting the aldehyde group on 3,4-dihydroxybenzaldehyde or 3,4,5-trihydroxybenzaldehyde with the amino group on the polymer to form a carbon-nitrogen double bond using sodium borohydride reduction after the reaction. The molar ratio of the catechol derivative to the polymer repeating unit is between 2:1 and 1:2, and the molar ratio of sodium borohydride to the polymer repeating unit is 1:1.

[0027] In another preferred case, the polymer grafted with catechol groups and containing multiple primary amino groups has good water solubility in water, the dissolution concentration can be greater than or equal to 10%, and the dissolved solution is clear and transparent with good fluidity.

[0028] The gelling buffer comprises two solutions, namely, a weakly acidic buffer A with a pH of 4 to 6 and a weakly alkaline buffer B with a pH of 7 to 10.

[0029] In another preferred embodiment, the weakly acidic buffer solution A with a pH of 4 to 6 can be water, hydrochloric acid dilution, acetic acid dilution, sulfuric acid dilution, etc.; or an acidic salt solution such as sodium dihydrogen phosphate and sodium bisulfate.

[0030] In another preferred embodiment, the weakly alkaline buffer solution B with a pH of 8 to 10 is an alkaline salt solution such as disodium hydrogen phosphate, sodium bicarbonate, sodium tetraborate, etc., among which the salt solution with a pH of about 9 is the best, and the salt concentration is 0.05 to 0.2M, and the best is 0.15M.

[0031] In another preferred embodiment, the tissue adhesion strength of the hydrogel may be ≥40 KPa.

[0032] In another preferred embodiment, the hydrogel has a good adhesion effect on the tissue surface, and the tissue adhesion strength of the hydrogel should not be significantly reduced or even enhanced after the gel is swollen and balanced in simulated body fluid (PBS) in vitro.

[0033] In another preferred embodiment, the water absorption rate of the hydrogel after swelling equilibrium in vitro is ≤5%.

[0034] In another preferred embodiment, the microstructure of the hydrogel does not undergo significant changes after the hydrogel reaches swelling equilibrium in simulated body fluid (PBS) in vitro.

[0035] The second aspect of the present invention provides a method for preparing the in-situ gelling anti-swelling enhanced tissue adhesive according to the first aspect of the present invention, comprising the steps of:

[0036] (a) modifying and grafting a polymer containing multiple primary amino groups with catechol groups to obtain a multifunctional crosslinking agent having both amino groups and catechol groups;

[0037] (b) using gelation buffer A to dissolve multi-arm polyethylene glycol powder with terminal modified active esters as gelation precursor solution A;

[0038] (c) using a gelation buffer B to dissolve a polymer modified with catechol groups and containing multiple primary amino groups to serve as a gelation precursor solution B;

[0039] (d) The gelling precursor solution A and the gelling precursor solution B are mixed evenly, and injected / applied to the wound tissue for bonding / sealing.

[0040] In another preferred embodiment, the order of steps (b) and (c) can be interchanged.

[0041] In another preferred embodiment, the steps (b) and (c) further include the following steps: sterilizing the gelling precursor solution A and the gelling precursor solution B using a sterile filter (0.22 μm sterile filter membrane), or sterilizing all major components by irradiation (effective sterilization dose ≥ 20 KGy) before preparation.

[0042] In another preferred embodiment, the temperature required for gelation is 25°C±10°C; preferably, 25°C.

[0043] In another preferred embodiment, the volumes of the gelling precursor solution A and the gelling precursor solution B may be the same or different, preferably 1:1.

[0044] In another preferred embodiment, the mass concentration of the gel precursor solution A after dissolution is 160-300 mg / mL, preferably 200±10 mg / mL.

[0045] In another preferred embodiment, the mass concentration of the gel precursor solution B after dissolution is 50-400 mg / mL, preferably 200±100 mg / mL.

[0046] In another preferred embodiment, in step (d), the gelling precursor A and the gelling precursor liquid B are pre-mixed and then loaded into a sterile syringe (single barrel) for injection and gelling, thereby forming the hydrogel material.

[0047] In another preferred embodiment, in step (d), the gelling precursor A and the gelling precursor liquid B are pre-mixed and then transferred to the desired wound tissue site by a smear brush to form a hydrogel in situ.

[0048] In another preferred embodiment, in step (d), the gelling precursor A and the gelling precursor liquid B are respectively loaded into a double-barreled syringe (mixer) at a volume ratio of 1:1, and injected and gelled to form the hydrogel material.

[0049] In another preferred embodiment, the time required for gelation is 5s to 180s.

[0050] The third aspect of the present invention provides a use of the in-situ gelling anti-swelling enhanced tissue adhesive described in the first aspect of the present invention for preparing medical materials.

[0051] In another preferred embodiment, the medical material is used in the fields of tissue adhesion, tissue sealing, tissue occlusion, drug sustained release, tissue repair, etc.

[0052] In another preferred embodiment, the medical material is one or more of the following:

[0053] (1) As a skin dressing (it can be applied to wounded skin to protect the injured skin and promote wound repair and healing);

[0054] (2) As a tissue adhesive (can replace surgical sutures to connect injured tissues);

[0055] (3) As a hemostatic material (can be applied to the human vascular system as an acute hemostatic occlusion for open bleeding wounds);

[0056] It should be understood that within the scope of the present invention, the above-mentioned technical features of the present invention and the technical features specifically described below (such as embodiments) can be combined with each other to form a new or preferred technical solution. Due to space limitations, they will not be described one by one here. BRIEF DESCRIPTION OF THE DRAWINGS

[0057] Figure 1 The interactions formed by the catechol groups within the gel adhesive are shown.

[0058] Figure 2 Three methods for preparing and using the in situ gelling anti-swelling enhanced tissue adhesive are shown.

[0059] Figure 3 The UV absorption graph of the synthesized PLL-DOHA and the corresponding grafting rate calculation are shown.

[0060] Figure 4 The in situ gelling anti-swelling reinforced tissue adhesive showed good injection moldability and excellent mechanical properties.

[0061] Figure 5 The gelation time of the in situ gelling anti-swelling reinforced tissue adhesive is shown.

[0062] Figure 6 The swelling properties of the in situ gelled anti-swelling enhanced tissue adhesive were shown.

[0063] Figure 7 The rheological properties of the in situ gelling anti-swelling reinforced tissue adhesive were shown.

[0064] Figure 8 The stability of the anti-swelling enhanced tissue adhesive adhered to the porcine skin surface by in situ gelation in water was shown.

[0065] Fig. 9 Shown is the tissue adhesion strength of the anti-swelling enhanced tissue adhesive before and after swelling equilibrium after in situ gelation in water.

[0066] Fig.10 The infrared (FTIR) spectra of the in situ gelling anti-swelling enhanced tissue adhesive before and after self-oxidation are shown. DETAILED DESCRIPTION

[0067] The inventors have designed and developed a novel double-crosslinked in-situ gelling tissue adhesive for the first time through extensive and in-depth research. The present invention uses a multi-arm polyethylene glycol solution with terminal modified succinimide active ester and a polymer solution containing multiple primary amino groups with side modified catechol groups as gelling precursors, and forms a double-crosslinked self-curing gel through spontaneous nucleophilic substitution reaction between the terminal active ester groups of polyethylene glycol and multiple amino groups on the polymer and multiple interactions between catechol groups on polylysine. The active ester groups and catechol groups can also form multiple bonds with amino groups, sulfhydryl groups, etc. on tissue proteins to achieve a strong tissue adhesion effect. At the same time, the slow self-crosslinking between the catechol groups can also resist the swelling caused by the entry of tissue fluid, so that the gel performance remains stable. On this basis, the present invention has been completed.

[0068] The main advantages of the present invention are:

[0069] (1) The present invention uses multi-arm polyethylene glycol (n-arm-PEG-NHS) with terminal modified active esters and a polymer modified with catechol groups and containing multiple primary amino groups as two main raw materials. Through the efficient amidation reaction between active esters and amino groups, a gel can be quickly formed to adhere to the wound without using any toxic catalyst groups.

[0070] (2) The present invention uses a polymer modified with catechol groups and containing multiple primary amino groups as a functional crosslinking agent. The catechol groups in the polymer network undergo a self-oxidative crosslinking reaction in a neutral pH environment, making the gel network tighter and effectively avoiding the large swelling caused by the entry of water molecules.

[0071] (3) The catechol on the cross-linking agent can also form multiple bonding reactions with amino groups, sulfhydryl groups, etc. on tissue proteins to achieve enhanced tissue adhesion.

[0072] (4) The present invention does not use any toxic substances and has high biological safety, thus providing a good basic guarantee for the product transformation of the project. DETAILED DESCRIPTION

[0074] The present invention will be further described below in conjunction with specific examples. It should be understood that these examples are intended only to illustrate the present invention and are not intended to limit the scope of the present invention. The experimental methods for the unrecorded specific conditions in the following examples are usually based on conventional conditions or the conditions recommended by the manufacturer. Unless otherwise stated, percentages and parts are weight percentages and weight parts.

[0075] Unless otherwise specified, the experimental materials and reagents used in the following examples can be obtained from commercial channels.

[0076] Example 1: Preparation of a novel in situ gelling anti-swelling enhanced tissue adhesive

[0077] First, under nitrogen protection conditions, dihydroxyphenylpropionic acid (DOHA) was used to graft polylysine (PLL) in 0.1M MES (morpholinoethanesulfonic acid) pH=6.0 buffer solution, and diimide carbodiimide (EDC) / N-hydroxysuccinimide (NHS) was used as a condensation agent. The group feed ratio was PLL: DOHA: EDC: NHS=1:0.5:1.5:0.5, the reaction time was 24h, and the initial concentration of PLL was 20mg / mL. After the reaction was completed, dialysis was carried out. The dialysate was 0.3mmol HCl. The dialysis was carried out for 3 days, and then dialyzed with pure water for 4h. The PLL-DOHA derivative was obtained by freeze drying, and its grafting rate was tested by ultraviolet absorption method. Figure 3 It can be seen that the grafting rate of the grafted PLL-DOHA derivative is about 16%, and the results are stable.

[0078] The preparation process of the gel is to first add a certain amount of hydrochloric acid diluent to the four-arm polyethylene glycol (4-arm-PEG-NHS) powder with terminal modified active esters, dissolve it as gel precursor A, and the concentration of 4-arm-PEG-NHS is 20%; then add an equal volume of 0.15M Na2HPO4 to PLL-DOHA or PLL to dissolve it as gel precursor B, and the concentration of PLL-DOHA or PLL is 5-20%, and then mix the two phases of gel precursors in equal amounts and apply or inject them to the target area. The in situ formed hydrogel is recorded as PEG-PLL / PEG-PLL-DOHA. In order to accelerate the oxidative cross-linking reaction of the catechol group in the hydrogel as a control, 0.25% H2O2 and 0.02mg / mL can be added to the 0.15M Na2HPO4 buffer, recorded as PEG-PLL-DOHA (+H2O2 / HRP).

[0079] like Figure 4 As shown: the mixed gelling liquid is sucked into the ring mold. It can be seen that the formed PEG-PLL hydrogel is colorless and transparent, the PEG-PLL-DOHA hydrogel is light yellow and transparent, and the PEG-PLL-DOHA (+H2O2 / HRP) is brown-black. It proves that there are o-phenyl groups in the hydrogel, which can quickly undergo oxidative polymerization under the catalysis of H2O2 / HRP, causing the appearance of the hydrogel to change significantly. At the same time, two transparent glass sheets were used to press the cylindrical gel, and it can be seen that the PEG-PLL hydrogel can be easily flattened; and the elastic modulus of the PEG-PLL-DOHA gel is further increased due to the non-covalent interaction between -DOHA or between -DOHA and -NH2, and the hydrogel has a certain toughness; and the PEG-PLL-DOHA (+H2O2 / HRP) gel has become very tough due to the rapid oxidative cross-linking reaction of -DOHA. Under the glass pressing, the glass sheet breaks, and the hydrogel still maintains its original shape well.

[0080] Example 2: Determination of gelation time of a novel in situ gelling anti-swelling enhanced tissue adhesive

[0081] The gelation time of PEG-PLL / PEG-PLL-DOH hydrogels was determined by observing the flow of gelling liquid in an inverted centrifuge tube. According to the above gel preparation method, the two phases of gelling precursor liquid were mixed and the timing was started. The flow of the gelling liquid after mixing was observed, and the time recorded until the flow stopped was regarded as the gelation time of the gel.

[0082] like Figure 5As shown, under the same ratio, PEG-PLL-DOHA has a slower gelation rate than PEG-PLL. This can prove that due to the introduction of -DOHA, some amino sites are occupied by DOHA in advance, thereby reducing the cross-linking sites with active esters (NHS), thereby slowing down the gelation rate of the gel. This can allow the gel to have a longer time to mix the gelling liquid evenly before application or injection.

[0083] Example 3: Determination of swelling rate of a novel in situ gelling anti-swelling enhanced tissue adhesive

[0084] Prepare PEG-PLL and PEG-PLL-DOHA gels according to the above method and record the initial weight (W b ), the hydrogel was placed in 20 mL PBS and placed in a 37°C constant temperature shaker for water absorption equilibrium. Samples were taken out at different set time points, and excess water on the surface was wiped off with absorbent paper and weighed (W a ).

[0085] Calculate the water absorption of the hydrogel according to the following:

[0086] Water absorption (%) = W a / W b ×100%(n=3)

[0087] like Figure 6 As shown: PEG-PLL-DOHA initially swelled slightly due to the rapid entry of water molecules, but the water absorption rate of the gel decreased after about 5 hours and then remained stable. Therefore, it is proved that the polymer network of the hydrogel can be made denser through the self-oxidation effect of -DOHA, thereby playing an anti-swelling and water discharge effect. At the same time, it can be found that as the amount of PLL-DOHA added increases, the cross-linking based on -DOHA in the gel becomes tighter, thereby discharging more water and causing the gel mass to drop below the initial mass. Therefore, after screening, when the PLL-DOHA content in the gel is 50 mg / mL (that is, the PLL-DOHA content in the gel-forming solution B is 10%), when the mass of the hydrogel reaches equilibrium in water, the mass is consistent with the initial mass, and almost no swelling occurs. Therefore, PEG-PLL-DOHA (PLL-DOHA: 50 mg / mL) hydrogel can be selected as the subsequent research object. At the same time, as a comparison, the PEG-PLL-DOHA (PLL: 50 mg / mL) hydrogel swells significantly, and the swelling rate has exceeded 200% on the 5th day, and continues to increase.

[0088] Example 4: Rheological properties of a novel in-situ gelling anti-swelling enhanced tissue adhesive

[0089] PEG-PLL / PEG-PLL-DOHA / PEG-PLL-DOHA (+H2O2 / HRP) were prepared according to the method described above, wherein the crosslinker content in the gelling solution B was 10%. In this experiment, the dynamic mechanical properties of the hydrogel, namely the storage modulus (G') and the loss modulus (G"), were determined by rheological method. According to the above method, gelling solutions with different component ratios were prepared and injected on the sample table for testing. The LVR middle amplitude value was selected and time scanned within 0 to 600 s. The network structure and dynamic mechanical response changes of different hydrogels at the beginning of gelling solution mixing were analyzed.

[0090] from Figure 7 It can be seen from the observation that in the initial stage of the hydrogel reaction (0-30S), the rate of increase of the storage modulus (G') is PEG-PLL-DOHA (+H2O2 / HRP) > PEG-PLL-DOHA > PEG-PLL. This is because in the same time, due to the cross-linking effect of DOHA, the total number of cross-linkers in the gel is greater. After reaching the platform stability, the G' of the PEG-PLL-DOHA (+H2O2 / HRP) hydrogel is significantly greater than that of PEG-PLL-DOHA and PEG-PLL. Therefore, -DOHA can form a more compact polymer network structure after oxidation and thus has better mechanical properties.

[0091] Example 5: Underwater bonding stability test of a novel in-situ gelling anti-swelling enhanced tissue adhesive

[0092] PEG-PLL and PEG-PLL-DOHA gels were prepared according to the above method and injected into the cleaned pig skin surface, wherein the cross-linking agent content in the gelling solution B was 10%. The hydrogels were then placed in 20 mL PBS and the solution was changed every day.

[0093] from Figure 8 It can be seen that as the soaking time increases, the color of PEG-PLL-DOHA gradually changes from the initial light yellow to brown and finally to dark brown, indicating that as the soaking time increases, the hydrogel continues to undergo oxidative cross-linking reactions, making the gel resist the entry of water molecules and enhancing its adhesion. It can be seen that PEG-PLL has undergone significant swelling on the 15th day, the gel height has increased significantly and then fell off, while the PEG-PLL gel can still firmly adhere to the pig skin.

[0094] Example 6: Determination of underwater bonding strength of a novel in-situ gelling anti-swelling enhanced tissue adhesive

[0095] PEG-PLL, PEG-PLL-DOHA, and PEG-PLL-DOHA (+H2O2 / HRP) were prepared according to the above method, wherein the crosslinking agent content in the gelling solution B was 10%. The gelling solutions were mixed evenly and applied between two pieces of pig skin, and then immersed in phosphate buffer solution PBS (pH=7.2) for 15 days, and the solution was changed every day. After soaking, the bonded pig skins were fixed to one end of a glass sheet with a roughened surface using cyanoacrylate glue. A universal mechanical stretching machine was used to clamp the glass sheet and stretch the bonded pig skins at a stretching rate of 5 mm / min. The maximum value measured during the stretching process corresponds to the bonding strength of the hydrogel, and the schematic diagram of the device is shown in 9-(b).

[0096] from Fig. 9 -(a) It can be seen from the observation that the PEG-PLL hydrogel swelled significantly after immersion, and part of the PEG-PLL gel had been sent out from the bonding area (red arrow), while the PEG-PLL-DOHA and PEG-PLL-DOHA (+H2O2 / HRP) gels had almost no morphological changes at the bonding area. Through the tensile test, Fig. 9 As shown in Figure 2(c), the PEG-PLL-DOHA hydrogel has a high tissue adhesion strength (~52 kPa) at the beginning of solidification before immersion, and the PEG-PLL-DOHA (+H2O2 / HRP) can reach ~72 kPa after complete oxidation, which is significantly higher than most PEG-based tissue adhesives (20~40 kPa) and porcine fibrin glue (~15 kPa). After immersion for 15 days, the adhesion of PEG-PLL-DOHA gel was improved again, and the adhesion strength reached ~68 kPa, while the adhesion strength of PEG-PLL-DOHA (+H2O2 / HRP) gel remained relatively stable. This may be because the addition of the oxidant has stabilized the gel structure before immersion. In contrast, the gel strength of PEG-PLL gel decreased significantly after immersion, and has dropped to less than 1 / 2 of the original adhesion strength. Therefore, the introduction of -DOHA groups can make the hydrogel achieve the effect of enhancing tissue adhesion.

[0097] Example 7: FTIR spectrum analysis of the novel in-situ gelling anti-swelling enhanced tissue adhesive before and after oxidation

[0098] According to the above method, PEG-PLL-DOHA was prepared and immersed in phosphate buffer PBS (pH=7.2) for 15 days, and the solution was changed every day to allow it to self-oxidize continuously. The hydrogel before and after self-oxidation was ground and freeze-dried, and KBr was added to the freeze-dried powder to make a tablet. Then a Fourier total reflection infrared tester was used at 4000-400cm -1 Scan and record within the range.

[0099] like Fig.10 The infrared spectrum shown in the figure shows that before self-oxidation, the -1 The CO vibration peak of non-oxidized catechol in PEG-PLL-DOHA gel and the peak at 842 cm -1 At the same time, after catechol self-oxidation, the hydrogen bonds between the original PLL-catechol-OH┅NH2-PLL and PLL-catechol-OH┅O-PEG molecules in the gel were -1 The broad peak at 3420 cm -1 The peak of hydrogen bond vibration within the PLL-catechol-OH┅O=C-benzoquinone-PLL molecule is located at the center.

[0100] Therefore, the catechol groups in the hydrogel can undergo self-oxidation to become a more reactive benzoquinone structure, and undergo a series of subsequent reactions such as disproportionate cross-linking to form a gel network structure with a tighter network structure, thereby resisting the entry of water molecules and enhancing the tissue adhesion strength of the gel.

Claims

1. A method for preparing an in-situ gelling anti-swelling enhanced tissue adhesive material, characterized in that: Adding a hydrochloric acid diluent having a pH of 4 to 6 to the four-arm polyethylene glycol 4-arm-PEG-NHS powder with terminal modified active esters, dissolving the powder and using it as a gel precursor A, wherein the weight percentage of 4-arm-PEG-NHS is 20%; An equal volume of 0.15 M Na2HPO4 is added to PLL-DOHA to dissolve it as a gel precursor solution B, wherein the weight percentage of PLL-DOHA is 5 to 20%; Then, the gelling precursor solution A and the gelling precursor solution B are mixed in equal volumes and then applied or injected to the target area to form the hydrogel PEG-PLL-DOHA in situ; The preparation of PLL-DOHA: under nitrogen protection conditions, in 0.1M morpholinoethanesulfonic acid MES pH=6.0 buffer solution, dihydroxyphenylpropionic acid DOHA was used to graft modify polylysine PLL, carbodiimide EDC / N-hydroxysuccinimide NHS was used as a condensation agent, the group feed molar ratio was PLL: DOHA: EDC: NHS=1:0.5:1.5:0.5, the reaction time was 24h, the initial concentration of PLL was 20mg / mL, and dialysis was carried out after the reaction was completed. The dialysate was 0.3mmolHCl, and the dialysis was carried out for 3 days, followed by dialysis with pure water for 4h, and then freeze-dried.

Images