TR2022013706T2 - HEMOSTATIC DRESSING AND METHOD FOR PRODUCING THEM - Google Patents

Abstract

A hemostatic dressing is provided. The hemostatic dressing comprises: a porous matrix layer incorporating a biocompatible polymer; a hemostatic layer comprising a polymer loaded on the porous matrix layer and incorporating polyhydric phenol containing groups; and a bonding layer arranged between said porous matrix layer and said hemostatic layer to prevent separation of the porous matrix layer from the hemostatic layer.

Description

DESCRIPTION HEMOSTATIC DRESSING AND METHOD FOR PRODUCING THE Same Technical Field The present specification relates to a hemostatic dressing and a method for producing the same. More specifically, the current specification; It relates to a hemostatic dressing having improved hemostatic performance, adhesion performance, anti-leakage performance and absorption performance, and a method for producing such hemostatic dressing. The Technique That Forms the Infrastructure Although great advances have been made in today's surgical techniques, hemostatic techniques for stopping bleeding during surgery are still in the early stages of their development. The main purpose of hemostatic agents is to stop bleeding. Hemostatic agents must have many functions in order to be used appropriately. More specifically, hemostatic agents must have the function of stopping bleeding. Hemostatic agents should be fixed to the desired tissue areas where hemostasis is needed, in a way that they can fulfill their hemostatic functions. Hemostatic agents must also have additional functions of absorbing blood and reducing blood loss. Many hemostatic agents are widely known and currently available on the market. Such hemostatic agents include fibrin-based hemostatic agents that utilize fibrin in the final stage of the body's blood clotting mechanism, cyanoacrylate-based hemostatic agents that use the same ingredients and the same principle as instant adhesives to produce hemostasis (see, for example, Korean patent publication 2009-0077759) and hemostatic agents for hemostasis. There are hemostatic substances consisting of natural polymer materials such as chitosan, gelatin and collagen that will function as physical barriers. However, fibrin-based hemostatic agents carry the risk of disease infection with ingredients derived from the living body and are difficult to administer to patients receiving anticoagulants for hemostasis. Cyanoacrylate-based hemostatic substances produce toxic substances such as formaldehyde when degraded in vivo. Hemostatic substances consisting of natural polymer materials do not have better hemostatic performance and their adhesion to living tissues is lower than expected. Moreover, some of the polymers have poor qualities in terms of anti-leakage and absorbency. Sponge-like structures that do not dissolve in blood are known as medical products for hemostasis. These medical products absorb blood from bleeding areas and stop bleeding based on the in vivo blood clotting mechanism. These medical products are easy to use because they do not stick to surgical instruments due to their lack of adhesive properties. These medical products can prevent blood from leaking in bleeding areas, thanks to their ability to absorb blood. However, since these medical products do not tend to adhere to living tissues, they require a separate means to ensure close contact with the tissues in the bleeding areas. Another disadvantage is the low hemostatic performance of these medical products. Another known medical product for hemostasis is in the form of a blood-soluble membrane structure. When they come into contact with the blood in the bleeding area, these medical products dissolve and form a physical film in the bleeding area and adhere to the bleeding area. The medical product performs hemostasis based on this mechanism. Since it has an adhesive function, this medical product eliminates the hassle of using a separate tool to ensure adhesion to living tissue. Although this medical product has the advantage of excellent hemostatic performance, it tends to stick to surgical instruments, causing discomfort during use and cannot prevent blood from leaking in the bleeding area. Detailed Description of the Invention Means for Solving Problems According to one aspect of the present disclosure, a porous matrix layer incorporating a biocompatible polymer; A hemostatic layer comprising a polymer containing groups including a benzene ring having one or more hydroxyl groups, loaded on the porous matrix layer and a bonding arranged between said porous matrix layer and said hemostatic layer to prevent separation of the porous matrix layer from the hemostatic layer. A hemostatic dressing is provided that includes a layer of According to another aspect of the present specification, a method for producing a hemostatic dressing is provided. The method in one application; (a) preparing a hemostatic layer formation solution by dissolving a polymer having phenol or polyhydric phenol containing groups incorporated into it in water, (b) providing a porous matrix consisting of a biocompatible polymer, (c) applying the hemostatic layer formation solution to the porous matrix surface. contacting the hemostatic layer forming solution so that it is impregnated into the pores of the porous matrix, and (d) drying the hemostatic layer forming solution to simultaneously form a hemostatic layer on the surface of the porous matrix and a bonding layer at the interface between the porous matrix and the hemostatic layer. Contains. Brief Description of the Drawings FIG. 1 is a schematic sectional view showing one application of a hemostatic dressing. FIG. 2 is a schematic diagram illustrating a process for achieving hemostasis and preventing blood from leaking when a hemostatic dressing suitable for one application is applied to a bleeding site. FIG. 3 shows the structures of a hemostatic multilayer sponge produced by impregnating a low viscosity chitosan-catechol solution into a gelatin layer and a hemostatic multilayer sponge produced by impregnating a high viscosity chitosan-catechol solution into a gelatin layer. FIGURE 4 presents scanning electron microscopy cross-sectional images of a multilayered sponge. FIGURE 5 shows photographs demonstrating the results of analyzing the anti-leak properties of hemostatic sponges. Mode of Carrying Out the Invention Unless defined otherwise, all technical and scientific terms used herein have the same meanings as commonly understood by one of ordinary experience in the field to which this invention pertains. In general, the terminology used here and below is well known and widely used in the field. The various embodiments disclosed herein will now be explained in more detail with reference to the accompanying drawings. One aspect of the present invention provides a hemostatic dressing. FIG. 1 is a schematic cross-sectional view showing an application of the hemostatic dressing. Referring to FIG. 1, hemostatic dressing 100 includes: a porous matrix layer 110 incorporating a biocompatible polymer; A hemostatic layer (130) containing a polymer loaded on the porous matrix layer (110) and containing polyhydric phenol groups, and said porous matrix layer (110) to prevent the separation of the porous matrix layer (110) from the hemostatic layer (130). A binding layer (120) arranged between said hemostatic layer (130). The porous matrix layer (110) can be in the form of woven fabric, non-woven fabric, foam polymer or sponge. The porous matrix layer 110 may be in contact with a wound site, such as a bleeding site, and may have a series of pores to accommodate absorbed body fluids. The porous matrix layer (110) can hold the blood, exudates (leaking fluids) and other fluids it contains and prevent them from flowing out of the wound area, thus achieving hemostasis. The matrix preferably consists of a biocompatible polymer that is reasonably elastic and has good wettability and absorbency for liquids such as blood. The pores in the matrix may be connected to each other to form a three-dimensional structure. Biocompatible polymer preferably; It is a material that can be absorbed in vivo over time through enzymatic degradation, biodegradation or hydrolysis following in vivo application. The biocompatible polymer is preferably a material that is insoluble in blood. More specifically, the biocompatible polymer is insoluble in water at a temperature of around 37°C and a neutral pH, which are in vivo conditions. The biocompatible polymer biodegrades within 3 months, preferably within 8 weeks, following in vivo application. It is preferred that the biocompatible polymer decomposes naturally within the time specified above after achieving the desired goal with wound closure. The biocompatible polymer can be a natural or synthetic biodegradable polymer. More specifically, the natural biodegradable polymer can be selected from the group consisting of collagen, fibronectin, fibrin, gelatin, elastin, chitin, chitosan, starch, dextrose, dextran, alginic acid, hyaluronic acid, cellulose and their derivatives and copolymers. More specifically, synthetic biodegradable polymer poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(D,L-lactide-co-glycolide) (PLGA), poly(α-caprolactone). (PCL), polyhydroxyalkanoate, polyanhydride, polyorthoester, polyamine and their derivatives and copolymers. The biodegradable polymers mentioned above are presented for illustration purposes only. Any polymer containing a hydrolyzable backbone such as an amide, aliphatic ester, urea, urethane, ether or peptide backbone can be used without particular limitation. Derivatives of biocompatible polymers can be in the form of salts of biocompatible polymers. Alternatively, derivatives of biocompatible polymers, modification products or modification products formed by replacing or modifying the functional groups of biocompatible polymers (e.g. -OH, -SH, -NH2, -COOH or -NHC(O)CH3) with other hydrophilic groups or other biocompatible polymers. It may be in the form of its salts. The matrix incorporating the biocompatible polymer can, for example, be in the form of a gelatin or collagen pad, which causes less inflammatory response after in vivo transplantation and provides excellent biofunctionality and biodegradability. Gelatin and collagen pads degrade at the high rate described above due to the action of proteases abundantly expressed in the corresponding tissues after hemostasis. The matrix has the primary hemostatic effect of biocompatible polymer. The secondary hemostatic effect of the matrix results from the absorption of body fluids and the effective compression of the expanding matrix. In addition, the matrix has excellent anti-leakage performance. Biocompatible polymer can be in the form of a cross-linked or non-cross-linked polymer. The biocompatible polymer preferably has a cross-linked structure that is insoluble in water and can maintain its shape even when exposed to blood. The biocompatible polymer can form a porous structure that can absorb blood while maintaining its shape. The porous structure can provide the matrix with the ability to prevent blood leakage. Biocompatible polymer can be prepared from a fiber, powder or liquid raw material that can be processed into a porous matrix. Preparation of the biocompatible polymer may involve acidification carried out at a pH in the range of 1.5-4 and neutralization with an alkaline solution. The porous matrix using the biocompatible polymer can be prepared by freeze-drying a gel, suspension or solution of the biocompatible polymer. The matrix can be crosslinked by mixing the raw material with a crosslinking agent in an organic solvent and polymerizing the raw material. The crosslinking agent may be, for example, glutaraldehyde, formaldehyde, acetaldehyde, ethylene formaldehyde, acetaldehyde or ethylene isopropanol. Although there is no specific limitation on the thickness of the porous matrix layer (110), it is in the range of 1 to 20 millimeters (mm), preferably 1 to 15 millimeters. If the thickness is below the lower limit, the porous matrix layer cannot function as an adequate absorption barrier to blood. Meanwhile, if the thickness exceeds the upper limit, it may take a very long time for in vivo degradation to occur. The hemostatic layer (130) exists as a separate layer from the porous matrix layer (110) and, thanks to the binding layer (120), it can be stacked together with the porous matrix layer (110) and without being separated from the porous matrix layer (110). The hemostatic layer 130 includes a polymer in which groups containing a benzene ring with one or more hydroxyl groups (i.e., phenol or polyhydric phenol-containing groups) are included. Groups containing a benzene ring are polyhydric phenol-containing groups, which can be preferred in terms of hemostatic performance and adhesion performance. Polyhydric phenol refers to a benzene ring with two or more hydroxyl groups (-OH). Polyhydric phenol is preferably in the form of catechol with two hydroxyl groups (-OH) or gallol with three hydroxyl groups (-OH). In some applications, some of the hydrogen atoms of phenol or polyhydric phenol can be replaced with other hydrophilic functional groups. The polymer is preferably a biocompatible polymer. The biocompatible polymer can be a natural or synthetic biodegradable polymer. More specifically, the biocompatible polymer is the same as described above with respect to the polymer used in the porous matrix layer 110. It can be selected from the group consisting of polymergelatin, collagen, fibrin, elastin, hyaluronic acid, alginic acid, alginate, carboxymethyl cellulose, carboxyethyl cellulose, hydroxymethyl cellulose, fibrin, heparin, chitin, chitosan and their derivatives, which have hemostatic performance, biodegradability and tissue regeneration effect. is taken into account and preferred. Phenol-containing groups or polyhydric phenol-containing groups such as catechol or gallol can be attached to the side chains or ends of the polymer backbone via functional groups such as -NH2, -SH, -OH or -COOH. The phenol or polyhydric phenol-containing group may be represented by the expression "phenol-L-" or "polyhydric phenol-L-", where L is a single bond, optionally -O-, -S-, -NH-, -C(O A C1-C10 aliphatic hydrocarbon chain containing )-, - C(O)O-, -OC(O)O- or -C(O)NH- or optionally -O-, -S-, -NH- It can be in the form of a C3-C10 alicyclic hydrocarbon chain containing -C(O)-, -C(O)O-, -OC(O)O- or -C(O)NH-. Typically, phenol or polyhydric phenol-containing groups can be conjugated to the biocompatible polymer through a reaction between the functional groups of the biocompatible polymer and the end groups of the phenol or polyhydric phenol-containing compound. The reaction can be carried out using various approaches such as chemical synthesis, electrochemical synthesis and enzymatic synthesis. Phenol or polyhydric phenol-containing groups are 1 to 50 mol%, 1 to 40 mol%, 1 to 30 mol%, or if the amount of polyhydric phenol-containing groups is below the lower limit, relative to the total mole of repeating units in the biocompatible polymer, low solubility in blood and Low hemostatic performance and adhesion performance may occur. Meanwhile, if the amount of phenol or polyhydric phenol-containing groups exceeds the upper limit, catechol can be oxidized. Various materials are suitable for use as hemostatic layer (130). Reaction Schemes 1 and 2 illustrate synthetic methods for chitosan polymers in which catechol-containing groups are introduced by a reaction between a chitosan polymer and a catechol-containing compound. Reaksi on emasi1 CHß xiysznm j +"9 ,Judge »HEVÄLHÜJ mi 10" EDCINHS 0% 0%; x+yâz=m b Reaction diagram 2 + i... Ã NaCNsHa Hot-w ww inch zam agi/7' In the formulas here, x, y and zl are all integers and m is an integer in the range of 50 to 6,000. Reaction Scheme 1 shows the formation of amide groups via carbodiimide coupling, where catechol terminated with a carboxylic acid such as 3,4-dihydroxyhydrocinnamic acid (i.e., hydrocaffeic acid) is conjugated to the amine groups of the chitosan backbone. The reaction is carried out at a pH between 5 and 6 to prevent catechol oxidation. A protonic solvent such as ethanol is used to prevent precipitation of water-insoluble chitosan during the reaction. As a result of this amide bonding, catechol is generally incorporated into the polymer backbone with a degree of substitution in the range of 4 to 30 mol% relative to the amine groups. By carrying out the reaction over an extended period of several days, it is possible to incorporate catechol with a degree of replacement of 80 mol% or more. Reaction Scheme 2 shows a reductive amination reaction where catechol terminated with an aldehyde such as 3,4-dihydroxybenzaldehyde reacts with the amine groups of the chitosan backbone and the resulting imine bonds (-C=N-) are converted to -C-N using a reducing agent such as NaBH4. - is reduced to its ties. As a result of this reductive amination, catechol can be incorporated into the polymer backbone in a short time with a degree of substitution in the range of 18 to 80 mol%. A material that dissolves and forms a gel upon contact with blood can be used in the hemostatic layer. Such materials include chitosan-catechol, hyaluronic acid-catechol, gelatin-catechol, collagen-catechol, polyamine-catechol, chitosan-gallol, hyaluronic acid-gallol, gelatin-gallol, collagen-gallol and polyamine-gallol. Among the various polymers mentioned above, polymeric compounds such as chitosan are known to stop bleeding. More specifically, the positively charged -NH3+ in chitosan attracts the negatively charged platelets in the blood, and as a result, the platelets come together quickly and rapid hemostasis is achieved. However, since this property is limited to a highly acidic environment with a pH of 2, chitosan largely lacks hemostatic capacity in the living body in a near-neutral environment (pH 7). On the other hand, the polymer used in the hemostatic layer of the hemostatic dressing according to the present description is modified with a polyhydric phenol such as catechol or gallol in order to have adhesive properties even in a neutral aqueous medium such as blood. As a result, the modified polymer dissolves in blood and, after a while, has an electrical attraction force with erythrocytes to form a hemostatic film on the bleeding tissue area. The hemostatic film is preserved for several days and can be naturally degraded by enzymes in the body. Thanks to the presence of aromatic and -OH groups in phenol or polyhydric phenol, the solid hemostatic layer (130) can easily dissolve in blood and then become a gel. It is possible to explain the hemostasis and adhesion mechanism of the hemostatic layer (130) in various ways. Thanks to the 715-715 interaction between aromatic groups, hydrogen bonding, coordination bonding with metal ions, or cation-n-bonding between metal ions and aromatic rings, the hemostatic layer (130) is formed by the use of a separate special cross-linking agent or by ultraviolet or thermal energy. It provides fast and strong hemostasis and tissue adhesion without the need for application. In a preferred embodiment, the hemostatic layer (130) can contain both phenol or polyhydric phenol-containing groups and their oxidized forms. As a result, -OH-containing benzene rings and :O-containing benzene rings can coexist in the polymer of the hemostatic layer 130. For example, the hydroxyl groups (-OH) of the catechol groups (1 OH) included in the polymer of the hemostatic layer (130) can be oxidized artificially or naturally during the production of the hemostatic dressing (100), and some of the catechol groups can be converted into quinone groups (\` "0). Quinone The formation of covalent bonds can also increase the ability of the hemostatic layer to adhere to tissue and stop bleeding by reacting with -NHz present in organic substances such as tissue proteins and blood proteins or in the hemostatic layer polymer dissolved in blood. After the bonds are formed, the oxo groups (:0) of the quinone groups are reduced to hydroxyl groups (-OH). Hydroxyl groups can maintain hydrogen bonds with nitrogen and oxygen atoms in organic substances, such as a Michael-type addition reaction (Reaction Scheme 3). It can form a covalent bond by reacting with an amine through the Schiff base reaction (Reaction Scheme 4). Reaction reaction 3 Organic matter Organic matter li? Y 1,/ \\ O 2;› f/' 4/ H Reaction diagram 4 Organic matter __ *4 "M \ *V , e; I 11_ Organic matter In one embodiment, the amount of oxidized groups in the hemostatic layer (130), Within this range, it is 50 mol% or less, preferably 40 mol% or less, and more preferably preferably 0.1 mol% or more, relative to the total mole of phenol or polyhydric phenol-containing groups and oxidized groups. Adhesion and ability to stop bleeding can be optimized. As the amount of oxidized groups increases, the number of strong bonds such as covalent bonds increases, as explained above, and therefore, a higher amount of -OH must be maintained within an appropriate range. Although the binding mechanism via covalent bonding is slower than the cation-n-binding mechanism explained above, it can provide a very high adhesion strength to the tissue protein. Therefore, when the hemostatic dressing (100) is applied to a bleeding human tissue, the ability of said hemostatic dressing (100) containing the hemostatic layer (130) to stop bleeding and prevent blood leakage can be improved. As the number of -OH groups attached to a benzene ring in groups containing phenol or polyhydric phenol included in the biocompatible polymer increases, the possibility of oxidation to quinone forms increases over time, but the number of sites that can interact with foreign proteins increases, thus providing improved hemostatic performance and adhesion performance. The binding layer (120) is arranged between the porous matrix layer (110) and the hemostatic layer (130). Hemostatic dressing (100) has a multi-layered structure in which the porous matrix layer (110) that can absorb blood is connected to the hemostatic layer (130) that has hemostatic performance and adhesion performance. It is very important that these two layers are well connected to each other without separation during transportation and use of the hemostatic dressing (100). On the other hand, since the water-soluble material of the porous matrix layer (110) and the water-insoluble material of the hemostatic layer (130) have different physical properties, it may be difficult to bond the two layers together without using a separate adhesive. In order to prevent the two layers from separating from each other, semi-finished products of the layers can be prepared and bonded together with an additional adhesive. On the other hand, the use of an adhesive may be limited due to toxicity and process inefficiency. Without the use of an additional adhesive, the constructive components of the hemostatic layer (130) can be simply coated on the surface of the porous matrix layer (110). On the other hand, in this case, the two layers may not interact with each other sufficiently, which may result in a weak adhesion between them. As a result, during transportation and use of the hemostatic dressing (100), the two layers may separate or break off, which seriously affects the quality of the finished product. Therefore, the bonding layer (120) serves the purpose of preventing the porous matrix layer (110) and the hemostatic layer (130) from separating from each other. The bonding layer (120) is preferably formed using a material miscible with the porous matrix layer (110) and hemostatic layer (130). In other words, the binding layer interacts with various functional groups of the protein or polysaccharide (e.g. -OH, -NH2 and -COOH) belonging to the porous matrix layer (110) and with the functional groups of the protein belonging to the hemostatic layer (130) through hydrogen bonding or covalent bonding as a result of condensation reactions. It is created using available materials. The material of the bonding layer is preferably a material derived from the constructive components of the porous matrix layer 110 and the hemostatic layer 130. For example, the material of the bonding layer may be a mixture or copolymer of the main constructive component of the porous matrix layer (110) and the main constructive component of the hemostatic layer (130). The binding layer 120 can be incorporated in various ways. For example, the bonding layer (120) can be incorporated by laminating a film consisting of the bonding layer material between two layers (110 and 130), coating the bonding layer material on the porous matrix layer (110), or impregnating the bonding layer material into the porous matrix layer (110). . In a preferred embodiment, the bonding layer (120) can be formed by sufficiently impregnating the hemostatic layer (130) components into the porous matrix layer (110). As a result, the binding layer (120) may contain a mixed region of the biocompatible polymer and the polymer in which phenol or polyhydric phenol-containing groups are included. During impregnation, the liquid components of the hemostatic layer (130) penetrate deep into the porous matrix layer (110), thus the hydrophilic groups of the biocompatible polymer forming the porous matrix layer (110) can interact strongly with the hydroxyl groups of the component of the hemostatic layer (130) through hydrogen bonding and bind. layer (120) can be allowed to perform its role. Therefore, the bonding layer (120) can be formed by using the constructive components of the other layers (110 and 130) and without a separate adhesive, and problems such as toxicity and poor biodegradability or bioabsorbability encountered when an additional adhesive is used can be avoided. The miscibility of the bonding layer with the other layers (110 and 130) is also high enough to prevent these layers (110 and 130) from separating from each other. In one embodiment, a concentration gradient in which the concentration of polyhydric phenol-containing groups decreases in the direction from the hemostatic layer (130) to the porous matrix layer (110) consisting of biocompatible polymer can be created in the mixed region, depending on how the binding layer (120) is formed. Thanks to the concentration gradient, the porous matrix layer (110) and the hemostatic layer (130) can be bonded to each other more strongly through the bonding layer (120) arranged between them. There is no specific limitation regarding the thickness of the bonding layer (120); in the most preferred case, it is in the range of 200 to 1,000 micrometers. If the thickness of the binding layer (120) is less than the lower limit, the porous matrix layer and the hemostatic layer can be separated from each other. Meanwhile, if the thickness of the bonding layer (120) exceeds the upper limit, the porous matrix layer may not be distinguishable, which may result in a decrease in the absorption and anti-leakage performance of this layer. As explained above, the presence of the bonding layer maintains the necessary physical properties in the constructive layers of the hemostatic dressing while ensuring a sufficient bonding strength between the layers without the problems caused by the use of an additional adhesive. The presence of the bonding layer is also advantageous in terms of achieving the desired product performance. The hemostatic dressing according to the present disclosure has the following advantageous effects. The bonding of the hemostatic layer with adhesive properties and the porous matrix layer with anti-leakage properties allows the hemostatic dressing to have adhesive properties and anti-leakage properties at the same time, while ensuring that the hemostatic dressing adheres to a bleeding tissue area without any additional means. It also prevents the hemostatic dressing from sticking to the surgical instruments during the surgical procedure. Therefore, the hemostatic dressing is easy to use, has excellent hemostatic performance, and can prevent blood from leaking from the bleeding site. The use of fibrin-based hemostatic agents, which are based on the body's blood clotting mechanism, is disadvantageous because hemostasis cannot be achieved in patients taking anticoagulant drugs. On the other hand, the hemostatic dressing according to the present description utilizes polyhydric phenol-containing groups that mimic the mussel-derived 3,4-dihydroxyphenylalanine (DOPA) protein, which exhibits excellent adhesive properties in water and provides high adhesion to the bleeding site. Thanks to the presence of both aromatic rings and -OH functional groups in polyhydric phenol-containing groups, the hemostatic component binds to blood proteins through various binding mechanisms such as hydrogen bonding, strong coordination bonding with metals, covalent bonding and cation-n- bonding, instead of the in vivo hemostatic mechanism. It enables the application of hemostatic dressing to any patient and rapid hemostasis. Instead of simply connecting the hemostatic layer and the porous matrix layer to each other, they are connected through the bonding layer. The presence of the binding layer prevents the hemostatic layer and the porous matrix layer from being separated from each other during transport and use of the hemostatic dressing. In particular, the use of polymer containing groups containing a benzene ring with hydroxyl groups in the hemostatic layer provides rapid hemostasis through a mechanism such as cation-n-bonding. In addition, the simultaneous presence of quinone-like groups capable of covalent binding with foreign proteins can achieve excellent hemostasis and strong adhesion at the same time. The presence of the porous matrix layer, which is tightly connected to the hemostatic layer by the bonding layer, allows the hemostatic dressing according to the present disclosure to have the ability to absorb blood and prevent blood leakage, as well as hemostatic performance and adhesion performance. Unlike classical hemostatic products, which must be transported under refrigerated conditions to maintain their performance, the hemostatic dressing in accordance with the current specification can be stored at room temperature for a long time and its quality does not change even in an environment with high humidity. After the hemostatic dressing according to the present disclosure is applied in vivo to achieve the purposes of hemostasis and leakage prevention, it is degraded to substances harmless to humans and absorbed in vivo. Therefore, the hemostatic dressing according to the present disclosure is considered to be highly biocompatible. FIGURE 2 is a schematic diagram illustrating a process for achieving hemostasis and preventing blood from leaking when the hemostatic dressing is applied to a bleeding site. According to another aspect of the present specification, a method for producing a hemostatic dressing is provided. The method in one application; (a) preparing a hemostatic layer forming solution by dissolving a polymer having polyhydric phenol containing groups incorporated into it in water, (b) providing a porous matrix consisting of a biocompatible polymer, (c) applying the hemostatic layer forming solution to the porous matrix surface, hemostatic It includes contacting the layer forming solution so that it is absorbed into the pores of the porous matrix and (d) drying the hemostatic layer formation solution and simultaneously forming a hemostatic layer on the surface of the porous matrix and a bonding layer at the interface between the porous matrix and the hemostatic layer. In step (a), a hemostatic layer formation solution is prepared by dissolving a polymer having polyhydric phenol containing groups incorporated into it in water. The polymer is used in an amount ranging from 0.1 to 3 parts by weight based on 100 parts by weight of water. Hemostatic layer formation solution Having a suitable viscosity within the range defined above, as measured at 4°C using a rotational viscometer, the hemostatic layering solution can be well absorbed and impregnated into the porous matrix, the viscosity being below the lower limit. In this case, a bonding layer with extremely high thickness may be formed and the hemostatic layer formation solution may be impregnated into the top layer, resulting in the disappearance of the porous matrix layer. Meanwhile, if the viscosity exceeds the upper limit, the porous matrix layer and the hemostatic layer may separate from each other and form a layer. It may not be possible to create a bonding layer. Phenol or polyhydric phenol-containing groups can be incorporated into the biocompatible polymer chain by chemical synthesis, electrochemical synthesis or enzymatic synthesis method and by using a compound having a phenol, catechol or gallol group. In one embodiment, the biocompatible polymer is dissolved in distilled water under heating, reaction agents, including phenol or a polyhydric phenol-containing compound, are added to the solution, and the mixture is stirred at a pH in the range of 4 to 6. For example, the hemostatic layer-forming solution containing polymer containing incorporated catechol groups is prepared chemically by the following procedure. First, a polymer/NHS mixed solution is prepared by adding a 1-hydroxy- solution to the polymer solution, and a polymer/NHS/ EDC mixed solution is being prepared. Then, a dopamine solution is added to the polymer/NHS/EDC mixed solution and allowed to react with this mixed solution to prepare a hemostatic layer forming solution containing the polymer containing the incorporated catechol groups. According to the method according to the present disclosure, the reaction between the phenol or polyhydric phenol-containing compound and the biocompatible polymer proceeds in an aqueous phase. During the reaction, polyhydric phenol, such as catechol, is oxidized to quinone-like compounds as a result of reaction with oxygen in water. The coexistence of catechol and quinone in the hemostatic layer leads to further improved adhesion performance and hemostatic performance. More specifically, it can be a compound with a phenol group incorporated into the biocompatible polymer chain, such as tyramine, L-tyrosine or phloetic acid. More specifically, the compound having a catechol or gallol group incorporated into the biocompatible polymer chain can be, for example, pyrocatechol, L-dopa, dopamine, norepinephrine, caffeic acid, gallic acid, gallolamine or 2-aminoethane pyrogallol. In step (c), the hemostatic layer formation solution is contacted with the porous matrix surface. In this context, for example, the hemostatic layer formation solution is dispersed into a PET mold, the porous matrix is arranged on the upper part of the dispersed solution, and the hemostatic layer formation solution is impregnated into the porous matrix under cooled conditions. As a result of impregnation, a bonding layer can be included between the porous matrix layer and a hemostatic layer. pm and more preferably within the range of 200 to 1,000 um. Within this thickness range, the absorption and anti-leakage performance of the porous matrix layer can be maintained at a high level. In step (d), the hemostatic layer formation solution is subjected to freeze-drying to form a solid hemostatic layer on the surface of the porous matrix. A bonding layer can also be completely formed at the interface between the porous matrix and the hemostatic layer. By the method in accordance with the present specification, a hemostatic layer is formed by a A hemostatic dressing with a multilayer structure consisting of a binding layer and a porous matrix layer can be produced simply by means of an impregnation process. The hemostatic dressing produced by the method according to the present specification is harmless to humans and has the functions of hemostasis, in vivo adhesion, blood absorption and leakage prevention. Therefore, hemostatic dressing can be used to significantly improve the performance of existing polymer hemostatic agents. Moreover, hemostatic dressing is effective in stopping bleeding and can be applied to various areas where organ perforation, bile leakage and post-surgical adhesion are required. In addition, the hemostatic dressing is easy to use because it does not stick to surgical instruments during use, and the quality of the hemostatic dressing does not change during transportation because its components do not contain hydrolyzable groups. The present specification will be explained more specifically with reference to the following examples. It will be obvious to those skilled in the art that these examples are for illustrative purposes only and that the scope of the specification should not be construed as limited to these examples. EXAMPLES Preparative Example 1: Gelatin-Catechol Synthesis 300 mL of triple distilled water was placed in a reactor and 3 g of gelatin, previously divided into small pieces, was added thereto. A gelatin solution was prepared by thawing the mixture under heating with appropriate stirring for at least 3 hours. It was confirmed by visual observation that the gelatin was completely dissolved. A gelatin-NHS solution was prepared by adding a solution containing mL of triple distilled water to the stirred gelatin solution. A gelatin-NHS-EDC solution was prepared by adding a solution containing 1.87 g of 1-ethyl-3-(-3-dimethylaminopropyl)carbodiimide hydrochloride ("EDC") in 15 mL of triple distilled water to the gelatin-NHS solution. A gelatin-NHS-EDC-dopamine solution was prepared by adding a solution containing 1.23 g of dopamine in 15 mL of triple distilled water to the gelatin-NHS-EDC solution. The reaction was carried out at room temperature for at least 5 hours. Once the reaction was completed, it was subjected to dialysis. The product was subjected to freeze-drying and stored. 2 mg of the obtained gelatin-catechol hemostatic sponge was dissolved in 1 mL of deionized water. The binding rate of catechol was confirmed using a UV-Vis spectrophotometer. As a result, the binding rate of catechol per gelatin molecule was determined as 33%. Preparation Example 2: Hyaluronic Acid- Catechol Synthesis 300 mL of triple distilled water was placed in a reactor and 3 g of hyaluronic acid, previously divided into small pieces, was added to it. A hyaluronic acid solution was prepared by dissolving the mixture with appropriate mixing. It was confirmed by visual observation that the hyaluronic acid was completely dissolved in mL of triple distilled water. A hyaluronic acid-NHS solution was prepared by adding a solution containing 2.27 g of NHS to the mixed hyaluronic acid solution. A hyaluronic acid-NHS-EDC solution was prepared by adding a solution containing 3.79 g of EDC in 15 mL of triple distilled water to the hyaluronic acid-NHS solution. A hyaluronic acid-NHS-EDC-dopamine solution was prepared by adding a solution containing 3.00 g of dopamine in 15 mL of triple distilled water to the hyaluronic acid-NHS-EDC solution. The reaction was carried out at room temperature for at least 5 hours. Following completion of the reaction, the reaction solution was dialyzed using a dialysis membrane (MWCO: 12,000-14,000, SpectraPor). The product was subjected to freeze-drying and stored. 2 mg of the obtained hyaluronic acid-catechol hemostatic sponge was dissolved in 1 mL of deionized water. The binding rate of catechol was confirmed using a UV-Vis spectrophotometer. As a result, the binding rate of catechol per hyaluronic acid molecule was determined as 4.3%. Preparation Example 3: Chitosan -Catechol Synthesis 300 mL of triple distilled water was placed in a reactor and 3 g of chitosan, previously divided into small pieces, was added to the mixture. 3 mL of a 5 M hydrochloric acid solution was dissolved with appropriate stirring and 0.5 M was added. 8 mL of sodium hydroxide solution was added to it. A chitosan solution was prepared by dissolving the mixture with appropriate stirring. A solution containing 2.07 g of HCA in mL of triple distilled water was confirmed by visual observation to form a chitosan-HCA solution. A solution containing 1.77 g of EDC in 200 mL of ethanol was added to the chitosan-HCA solution. The reaction was carried out at room temperature for at least 1 hour. Once the reaction was completed, it was subjected to dialysis. The product was subjected to freeze-drying and stored. 2 mg of the obtained chitosan-catechol hemostatic sponge was dissolved in 1 mL of deionized water. The binding rate of catechol was confirmed using a UV-Vis spectrophotometer. As a result, the binding rate of catechol per chitosan molecule was determined as 8.9%. Preparation Example 4: Gelatin Layer/ Production of Hemostatic Multilayer Sponge with Binding Layer/Chitosan-Catechol Layer Structure In this example, a hemostatic multilayer sponge was produced as a hemostatic dressing by using gelatin as the porous matrix and chitosan-catechol synthesized in Preparation Example 3 as the hemostatic layer. The hemostatic layer of the hemostatic multilayer sponge was formed by the following procedure. First, a hemostatic layer formation solution was prepared by dissolving 1 g of the freeze-dried chitosan-catechol synthesized in Preparation Example 3 in 100 mL of triple distilled water with stirring. The viscosity of the solution was determined as 74.3 CP. After the solution was dispersed into a mold Then, gelatin (a sponge-like absorbent polymer) as a porous matrix was placed on the dispersed solution to ensure that the hemostatic layer formation solution was properly absorbed into the gelatin, and then freeze-drying was carried out. The resulting hemostatic multilayer sponge was formed from a porous matrix layer containing gelatin. bottom), a binding layer (middle) and a hemostatic layer (top) containing chitosan containing catechol groups incorporated into it. Preparation Example 5: Production of a Hemostatic Multilayer Sponge with a Gelatin Layer/Binding Layer/Chitosan-Catechol Layer Structure. The multilayer sponge was produced in the same manner as in Preparative Example 4, except that a hemostatic layer-forming solution was prepared by dissolving 0.5 g of chitosan-catechol in 100 mL of triple distilled water. The viscosity of the hemostatic layer formation solution was determined as 25.7 CP. Preparation Example 6: Production of Hemostatic Multilayer Sponge with Gelatin Layer/Binding Layer/Gelatin-Catechol Layer Structure. It was produced in the same manner as in Fig. 4. The viscosity of the hemostatic layer formation solution was determined as 14.6 CP. Preparation Example 7: Production of Hemostatic Multilayer Sponge with Gelatin Layer/Binding LayerHyaluronic Acid-Catechol Layer Structure A hemostatic multilayer sponge contains 1 g of hyaluronic acid-catechol 100. ML is dissolved in water and the preparation of a hemostatic layer is produced in the same way. Atik Production of Layered Sponge by Coating Process. A hemostatic layer formation solution was prepared by dissolving 1.5 g of the freeze-dried chitosan-catechol synthesized in Preparation Example 3 in 100 mL of triple distilled water with stirring. The viscosity of the hemostatic layer formation solution was determined as 745 CP. Chitosan-catechol solution was evenly applied to a gelatin layer and then freeze-drying was performed to produce a hemostatic multilayer sponge. Comparative Example 1: Production of Hemostatic Multilayer Sponges with Gelatin Layer/Binding Layer-Chitosan-Catechol Layer Structure Chitosan-catechol solutions having viscosity values in the range of 10 to 2,500 CP as measured at 4°C using a rotary viscometer were used as hemostatic layer forming solution. has been prepared. Hemostatic layering solutions were used to produce hemostatic multilayer sponges with a multilayer structure. In multilayer structure, a bonding layer is properly formed and the constructive layers are not separated from each other. On the other hand, when a chitosan-catechol solution with a viscosity outside the range defined above was used, it was difficult to obtain a multilayer sponge with a stable structure. FIGURE 3 shows the structures of a hemostatic multilayer sponge produced by impregnating a low viscosity chitosan-catechol solution into a gelatin layer and a hemostatic multilayer sponge produced by impregnating a high viscosity chitosan-catechol solution into a gelatin layer. Referring to FIG. 3, when a chitosan-catechol solution with a viscosity of <10 CP was absorbed into a gelatin layer, an excessively large amount of chitosan-catechol solution was absorbed into the gelatin layer, making it impossible to form a multilayer structure. Meanwhile, as shown in part B of FIGURE 3, when a chitosan-catechol solution with a viscosity of 2,500 CP was used, a bonding layer was not formed and as a result, a sufficient bonding strength was not achieved, resulting in the separation of the gelatin layer and the chitosan-catechol layer. has been concluded. The viscosity of the solution used to produce the sponge shown in part A of FIG. 3 was 8.3 CP. The viscosity of the solution used to produce the sponge shown in part B of FIGURE 3 was 2.717 CP. Experimental Example 1: Observation of Multilayer Sponge In this experimental example, the cross-section of the multilayer sponge produced in Preparative Example 4 was measured at 40X and 200X magnification using a scanning electron microscope. has been observed. FIGURE 4 presents scanning electron microscopy cross-section images of the multilayered sponge. Referring to FIG. 4, the sponge has a three-layer structure consisting of a hemostatic layer (top), a binding layer (middle), and a porous matrix layer (bottom). Experimental Example 2: Measurement of Tensile Strength of Sponges of the bonding layers of multilayer sponges produced in Preparatory Examples 4, 5 and 8. In this experimental example, the tensile strengths of a gelatin sponge, a chitosan sponge and the multilayer sponge produced in Preparatory Example 4 were measured. First, Each sponge was cut into a sample with a width of 0.6 cm and a length of 0.5 cm. Then, both ends of the sample were clamped to a universal testing machine. The sample was pulled at a speed of 5.0 mm/min. The experiments were carried out independently in five replicates. Table 1 shows the results of analyzing the strength properties of the sponges. Among the values obtained by the tests, the elasticity moduli (slopes of the stress-strain curves) of the gelatin sponge and chitosan sponge were 0.85 and 1.02, respectively. On the other hand, the elastic modulus of the multilayer sponge was 1.71, which corresponds to 2.01 times that of the gelatin sponge and 1.68 times that of the chitosan sponge. Strength Properties Gelatin Sponge Chitosan Sponge Multilayer Sponge Maximum load (N) 3.98 1.51 3.88 Experimental Example 3: Comparison of Anti-Leakage Powers of Sponges In this experimental example; The anti-leakage strengths of a gelatin sponge, a chitosan sponge, and the multilayer sponge produced in Preparative Example 4 were compared. First, a red dye solution was dropped onto the gelatin sponge, chitosan sponge and multilayer sponge. The upper and lower surfaces of each sponge were observed. FIGURE 5 shows photographs demonstrating the results of analyzing the anti-leak properties of hemostatic sponges. Referring to FIGURE 5, the red dye solution was observed only on the lower surfaces of the gelatin sponge and multilayer sponge, indicating that the sponges have anti-leakage power. On the other hand, the red dye solution percolated from the lower surface to the upper surface of the chitosan sponge, which indicates that Experimental Example 4: Comparison of Adhesion Powers to Pig Colon Tissues In this experimental example, the adhesion powers of a gelatin sponge, a chitosan sponge and the multilayer sponge produced in Preparation Example 4 were compared. Rectangular shaped colon tissues were prepared, cut to size of 1.0X1.0 cm. After dropping animal blood onto the colon tissues, the sponge (1.0X1.0 cm in size) was attached to the corresponding colon tissue. Then, both ends of the sample were clamped into a universal testing machine. The tensile-shear strength (N) when separated from the sponge column tissue was recorded. Adhesion strength (kPa) was calculated based on the tensile-shear strength (N). Experiments were performed independently in triplicate. Table 2 shows the adhesion strength of sponges to blood-stained porcine colon tissues. The adhesion strength to the tissue increased according to the order of gelatin sponge < chitosan sponge < multilayer sponge. Each experimental value is the mean ± standard deviation. Sponge Adhesion Strength (kPa) Gelatin sponge 5.23 ± 0.46 Chitosan sponge 11.80 ± 1.30 Multilayer sponge 16.93 ± 9.25 Experimental Example 5: Anti-leakage and Adhesion Properties in Pig Colon Tissues In this experimental example; The anti-leakage and adhesion properties of a gelatin sponge, a chitosan sponge, and the multilayer sponge produced in Preparative Example 4 were analyzed. A bursting test device was used to analyze the anti-leakage and adhesion properties of the sponges. Experiments were performed independently in triplicate. A hole was created in a porcine colon tissue using a 0.8 cm diameter hole punch, and gelatin sponge, chitosan sponge and two-layer sponge were each cut to 2.0X2.0 cm size and placed on the 8 mm diameter hole. The colon tissue on which the sponge was placed was placed on the diaphragm of the bursting test device, air was blown into the bursting test device, and the pressure at which the sponge separated from the colon tissue was recorded. Table 3 shows the results of analyzing anti-leakage and adhesion properties. Since its adhesion strength to the tissue was close to 0 hPa, the gelatin sponge was easily peeled off from the tissue. Since the chitosan sponge did not have the power to prevent leakage, the air blown into the test device leaked through the pores of the sponge, making it impossible to measure the pressure. On the other hand, since the multilayer sponge has both anti-leakage power and tissue adhesion power, anti-leakage-adhesion power could be analyzed. The average anti-leakage-adhesion strength of the multilayer sponge to porcine colon tissue was determined to be 37.8 hPa. Sponge Pressure (hPa) Average Standard Deviation Chitosan sponge ~ 0 - - Gelatin sponge ~ 0 - - Multi-layered 35.5 37.8 2.8 sponge 37.8 Experimental Example 6: Comparison of User Convenience In this experimental example; An evaluation of whether a gelatin sponge for use as hemostatic agents, a chitosan sponge, and the multilayer sponge produced in Preparative Example 4 is non-adhesive to surgical instruments and adhesive to living tissues was made, taking into account the in vivo aqueous environment during surgery. Each sponge was cut to a size of 1.0X1.0 cm, the forceps were exposed to physiological saline and a rectangular colon tissue was prepared. After the sponge (1.0X1.0 cm in size) was attached to the tissue with forceps, it was verified whether the sponge was separated from the forceps. After the sponge was attached to the colon tissue in the same manner as described above, it was verified whether the sponge had separated from the tissue. Table 4 shows whether each sponge, after being attached to forceps or tissue, detached from the forceps or colon tissue when exposed to physiological saline. Gelatin sponge had poor adhesion to instruments and tissue, and chitosan sponge had good adhesion to instruments and tissue. On the other hand, the surface of the multilayer sponge in contact with the tools had a poor degree of adhesion to the tools, and the surface of the multilayer sponge in contact with the tissue had a good degree of adhesion to the tissue. As a result, the multilayer sponge was superior to other sponges in terms of user convenience. Sponge Adhesion to Instruments Adhesion to Tissue Gelatin sponge Poor Poor Chitosan sponge Good Good Multilayer sponge Poor Good Experimental Example 7: Hemostasis Performance in Rat Liver Bleeding Models In this experimental example, the hemostatic multilayer sponge produced in Preparative Example 4 was used in rat liver bleeding models to evaluate its hemostatic performance. 9-week-old male Sprague Dawley rats were used as experimental animals. The animals were divided into two groups: a control group that was left untreated after bleeding and an experimental group in which gelatin sponge or multilayer sponge was applied after bleeding was achieved. Afterwards, the injection site was sterilized. After the abdomen was cut open, the liver was exposed and the center of the exposed liver was injured with a nester to remove foreign substances. In the control group, no treatment was applied after bleeding caused by injury. In the experimental group, gelatin sponge or multilayer sponge was placed over the bleeding area. 3 minutes after bleeding occurred, it was observed with the naked eye whether hemostasis was achieved. The results are shown in Table 5. In all animals in the control group and in the experimental group where gelatin sponge was used, hemostasis was not observed even after 3 minutes of bleeding. In other words, the rate of hemostasis achieved within 3 minutes of bleeding time was 0%. On the other hand, in the experimental group where multilayer sponge was used, hemostasis was not observed. Hemostasis was observed in all animals in the group. In other words, the rate of hemostasis achieved within 3 minutes of bleeding was 100%. These results revealed that the multilayer sponge was effective in stopping bleeding in the bleeding areas of the dissected rat livers. rate not provided (%) Control group X 0 Gelatin sponge applied to the models. Experiments were performed independently in triplicate. 9-week-old male Sprague Dawley rats were used as experimental animals. The animals were divided into two groups: a control group to which a bandage was applied after inducing bleeding, and an experimental group to which a multilayer sponge was applied after inducing bleeding. After full anesthesia was achieved in each rat, the injection site was sterilized. Approximately 3 cm of the middle area from under the rat's chin to its chest was cut and bleeding was caused from the jugular vein using a 31G needle. In each animal in the control group, the bleeding site was covered with a dressing for hemostasis by compression. In each animal in the experimental group, the bleeding site was covered with multilayer sponge and bandage folded together for hemostasis by compression. After a 2-minute period, the gauze or the multilayered sponge folded with the gauze was carefully removed without damaging the clots in the hemostasis area, and it was verified with the naked eye whether hemostasis was achieved. After a 2-minute period, it was observed with the naked eye whether hemostasis was achieved. The results are shown in Table 6. Hemostasis was not observed within 2 minutes following the onset of bleeding in all animals in the control group, where a dressing was used for hemostasis by compression. On the other hand, hemostasis was observed within 2 minutes following the onset of bleeding in all animals in the experimental group, where multilayer sponge was used. In other words, the rate of hemostasis achieved within 2 minutes of bleeding was 100%. These results revealed that the multilayer sponge was effective in stopping bleeding in the jugular vein bleeding areas of rats. Group Weight (g) The rate of hemostasis achieved within 2 minutes was (%). ) not provided Control 307 Cytotoxicity was tested by an indirect method. First, 2.0% agar and 2xMEM were mixed at a 1:1 ratio. The mixture was plated on mouse-derived L929 cells and a multilayer sponge cut to 1.0X1.0 cm was placed on top. Cells were cultured in a saturated steam condition in an incubator set at 5% CO2 and 37±1°C. After 48 hours of culture, the morphology of the cells was observed using an upright microscope. Approximately 2 mL of neutral red was added and incubation was carried out in the incubator for 15-30 minutes. After removal of neutral red, cells were washed with PBS. It was confirmed whether the color of the cells changed or not. The intradermal reaction was tested as follows: First, three healthy New Zealand white rabbits weighing 2.0 kg or more were selected and the hair on the back of each animal was shaved. A 0.2 mL extract from each sponge was injected into the animal intradermally. The occurrence of adverse reactions such as erythema and edema at the injection site has been observed. Table 7 shows the results of cytotoxicity and intradermal reaction tests. Cytotoxicity testing revealed that slightly deformed or degenerated cells were observed immediately beneath the gelatin sponge and multilayer sponge, and limited cell deformation and discoloration were observed immediately beneath the chitosan sponge. Based on these results, it was possible to determine that the cytotoxicity of all sponges was insignificant. Intradermal reaction testing revealed that extracts of all sponges did not cause any adverse reactions such as erythema and edema when injected intradermally. Therefore, it has been determined that sponges are substances that do not cause intradermal reactions.TR TR

Claims (13)

1. A hemostatic dressing comprising a porous matrix layer incorporating a biocompatible polymer; A hemostatic layer comprising a polymer containing groups including a benzene ring having one or more hydroxyl groups loaded on the porous matrix layer and a bonding arranged between said porous matrix layer and said hemostatic layer to prevent separation of the porous matrix layer from the hemostatic layer. Contains the layer. 2. The hemostatic dressing according to claim 1, wherein the biocompatible polymer is a natural or synthetic biodegradable polymer. 3. The hemostatic dressing according to claim 1, wherein the porous matrix layer is a gelatin or collagen pad. 4. Hemostatic dressing according to claim 1, wherein the polymer of the hemostatic layer has groups containing phenol or polyhydric phenol incorporated into it. 5. The hemostatic dressing according to claim 4, wherein the polymer is a biocompatible polymer. 6. The hemostatic dressing according to claim 4, wherein each of the phenol or polyhydric phenol containing groups is represented by the phrase “phenol-L-” or “polyhydric phenol-L-“, where L is a single bond, optionally -O a Ci-Cio aliphatic hydrocarbon chain containing -, -S-, -NH-, -C(O)-, -C(O)O-, -OC(O)O- or -C(O)NH- or optionally a C3-C10 containing -O-, -S-, -NH-, -C(O)-, -C(O)O-, -OC(O)O- or -C(O)NH- It is in the form of an acyclic hydrocarbon chain. 7. The hemostatic dressing according to claim 1, wherein the hemostatic layer contains both -OH-containing benzene rings and :O-containing benzene rings. 8. The hemostatic dressing according to claim 1, wherein the bonding layer is formed using a material miscible with the porous matrix layer and the hemostatic layer. 9. The hemostatic dressing according to claim 8, wherein the material is a mixture or copolymer of the major constituent of the porous matrix layer and the major constituent of the hemostatic layer. 10. The hemostatic dressing according to claim 4, wherein the binding layer comprises a mixed region of biocompatible polymer and polymer in which phenol or polyhydric phenol-containing groups are incorporated. 11. A method for producing a hemostatic dressing comprising (a) preparing a hemostatic layer forming solution by dissolving in water a polymer having phenol or polyhydric phenol containing groups incorporated into it, (b) providing a porous matrix consisting of a biocompatible polymer, (c) contacting the hemostatic layer formation solution to the porous matrix surface in such a way that the hemostatic layer formation solution is absorbed into the pores of the porous matrix, and (d) drying the hemostatic layer formation solution to form a hemostatic layer on the surface of the porous matrix and a hemostatic layer at the interface between the porous matrix and the hemostatic layer. It involves the simultaneous creation of the binding layer. 12. The method according to claim 11, wherein the hemostatic layering solution has a viscosity in the range of 10 to 2,500 CP as measured at 4°C using a rotational viscometer. 13. The method according to claim 11, wherein the bonding layer has a thickness in the range of 1 to 1,500 µm.