Biocompatible polymer device

Description: TECHNICAL FIELD [Detailed description of the invention]

【0001】

(Field of invention) The present invention generally relates to synthetic polymer compositions that form an interpenetrating polymer network when administered to tissues. Such compositions are particularly well suited for use in various medical applications where the physical strength of the administered material is important. For example, the compositions of the invention designed to be used to bond tissues together are sufficient to provide effective alternatives to other surgical adhesive means such as sutures and medical staples. Strength. [0002]

(Background of invention) The use of polymer compositions in tissue engineering, especially compositions consisting of synthetic polymers, is now widely recognized. In contrast to many naturally derived compositions, synthetic polymer compositions can be formulated to exhibit certain physical properties, such as gel strength, and biological properties, such as degradability. [0003]

In various tissue engineering applications, it can be administered as a liquid, but it is desirable to use a composition that later forms a hydrogel at the site of administration. Such in situ hydrogel-forming compositions are more convenient to use. Because they can be administered as liquids from a variety of different devices. And they are more suitable for administration to any site. Because they are not preformed. Many different mechanisms that can be used to promote hydrogel formation in situ have been described. For example, a photoactive mixture of water-soluble copolyester prepolymer and polyethylene glycol has been described to form a hydrogel barrier and form a drug release matrix. In another approach, block copolymers of Pluronic and Poloxamer that are soluble in cold water but form insoluble hydrogels that adhere to tissues at body temperature have been designed (Leach, et). al., Am.J.Obstet.Gynecol.162: 1317-1319 (1990)). Polymerizable cyanoacrylates have also been described for use as tissue adhesives (Ellis, et al., J. Otolaryngol. 19: 68-72 (1990)). In yet another approach, a two-part synthetic polymer composition has been described. It forms covalent bonds with each other and with exposed tissue surfaces when mixed together (PCT WO 97/22371). In a similar approach involving a two-part composition, a mixture of protein and bifunctional crosslinker has been described for use as a tissue adhesive (US Pat. No. 5,583,114). [0004]

In other tissue engineering applications, it is neither necessary nor desired that the composition be in liquid form when administered. In fact, it may be advantageous in some situations to apply a composition that is gelatinous and is a paste-like or even preformed solid implant. This is because these forms tend to stay in place after administration more easily than the liquid form. [0005]

For many tissue engineering, this composition may form after administration a matrix strong enough to resist the biological and physical forces exerted over a long period of time to achieve its intended purpose. is important. Intensity is especially important when the composition is used as a partial or total t-alternative to the suture. In addition, for applications where adhesion to tissue is important, it is necessary for the formulation to have adequate adhesive strength. Cyanoacrylates are highly effective adhesives that form strong matrices, but are too toxic for internal use and are therefore not approved for such applications. Therefore, a task of the present invention is to provide a highly potent medical sealant rather than a toxic one. [0006]

(Gist of the invention) The present invention relates to polyfunctional polymer-containing compositions. It is specifically designed to show outstanding strength with respect to certain embodiments. The composition contains a first polyfunctional synthetic polymer having m polyfunctional groups, a polyfunctional cross-linking agent having X; n functional groups, Y; and a tensile strength enhancer, and as a result, all three. When the components of are mixed together, X and Y react to form a covalently bonded three-dimensional interpenetrating polymer network. In addition, the tensile strength enhancer can become covalent within its network, or it can become physically trapped therein. [0007]

In a preferred embodiment of the invention, the polyfunctional crosslinker is a second polyfunctional synthetic polymer. [0008]

One aspect of the invention is the inclusion of a tensile strength enhancer, which can consist of various compounds capable of forming structures with suitable properties as described elsewhere in the specification. Such compounds include, among other things, Vicryl, glass wool, plastics, resins and fibers. [0009]

Functional groups X and Y can be any pair of reactive moieties that form a covalent Z under appropriate conditions (eg, sulfhydryl, succinimidyl, acrylate and amino, regardless of the particular order. , Ethers, esters or disulfides). [0010]

To enhance or inhibit the degradation of the interpenetrating polymer network, the synthetic polymer core may further comprise a chain extender. For example, α-hydroxy acids, poly (lactones), poly (orthocarbonates) and poly (phosphoesters) moieties can act to enhance biodegradability. In addition, the chain extender can be an enzyme substrate. [0011]

In another aspect of the invention, the composition further comprises any component (eg, protein, antibiotic, growth factor and hemostatic agent). [0012]

In a preferred embodiment of the invention, any component thereof is present in the form of rigid nanofibers (eg, those formed by methylated collagen). [0013]

In yet another aspect of the invention, X and Y can each be 4 or greater to provide more efficient cross-linking. [0014]

In one embodiment, the composition of the invention comprises a pair of reactive polyalkylene oxides, rigid nanofibers (eg, methylated collagen and tensile strength enhancer). [0015]

Other aspects of the invention are described elsewhere herein. [0016]

(Detailed description of the invention) The present invention relates to synthetic polymer compositions that form an interpenetrating polymer network when administered to tissue sites according to one embodiment. Within a short time (less than 3 minutes) after administration, the composition forms a network that develops with a high degree of binding (tension) strength. It is intended that the matrix will have a tensile strength of at least 10%, and more preferably 20%, of the tensile strength of the cyanoacrylate. In a preferred embodiment, such a mesh exhibits higher tensile strength than 60 N / cm2. Preferred compositions are designed to be anchored in place by mechanical and / or chemical means that together encapsulate tissues separated as a result of various disease states or surgical procedures. Thus, a composition that forms a mesh with an adhesive strength greater than> 50 mmHg (ie, the breaking pressure required to desorb the adhesive and create a leak) is when strong tissue adhesion is desired. Preferred for application. [0017]

In another embodiment, the invention relates to a matrix preformed in dry sheet form that can be applied more quickly than sprayed liquid (laparotomy). Adhesion with very high burst strength in a short time (30 seconds) can also allow the use of active sheets in areas that must be challenged immediately after application (3 minutes for 30 seconds). [0018]

In yet another embodiment, the invention includes activated collagen sponges and sheets as further described in Examples 4 and 5 below. [0019]

(Definition) The following definitions are provided to further describe the various aspects of the preferred embodiments of the present invention. [0020]

The term "gel" refers to the state of matter between a liquid and a solid. The "gel" itself has some liquid properties (ie its shape is elastic and deformable), and some solid properties (ie its shape is 3 on a two-dimensional surface). It is sufficiently discrete to maintain its dimensions). Thus, the "gelling time", also referred to herein as the "geling time", refers to the time it takes for the composition to become non-fluid under the most mild stress. This is generally shown to achieve gel strength and is expressed as a modulus G'of 102 dynes / cm2 or greater in less than 1 minute. [0021] [0021]

The term "cohesive strength" means that the compositions of the invention are intact (that is, they may rupture, rupture, or rupture) when exposed to physical stress or environmental conditions. The ability to remain. Bond strength is measured herein as a function of "tensile strength". [0022]

The term "adhesive strength" refers to the ability of the compositions of the invention to remain adherent to the tissue at the site of administration when exposed to physical stress or environmental conditions. [0023]

The term "polymer" refers to a molecule consisting of individual chemical moieties. This individual chemical moiety may be the same or different, but is preferably the same and is bound together. As used herein, the term "polymer" refers to the structure of individual chemical moieties, as well as branches (eg, "multi-arms" or "stars"), that combine end-to-end to form a linear molecule. Refers to individual chemical moieties that are combined together in the form of. [0024]

The term "biocompatibility" refers to the ability of a composition of the invention to be applied to a tissue without causing significant inflammation or fibrosis or other harmful tissue responses. [0025]

The term "synthetic polymer" refers to a polymer that does not exist in nature and is produced by chemical or recombinant synthesis. Therefore, naturally occurring proteins (eg collagen and naturally occurring polysaccharides (eg hyaluronic acid)) are particularly excluded. Proteins (eg, synthetic collagen) and carbohydrates (eg, synthetic hyaluronic acid) and their derivatives are included. [0026]

The term "activated synthetic polymer" has at least one functional group (eg, sulfhydryl group) that can react with a corresponding reaction partner (eg, sulfhydryl reactive group) to form a covalent bond, or A synthetic polymer that has been chemically modified in this way. The term "polyfunctional activated" or "polyfunctional" refers to a synthetic polymer having two or more functional (usually nucleophilic or electrophilic) groups. Polyfunctional activated forms of synthetic polymers include bifunctional activated polymers, trifunctional activated polymers, tetrafunctional activated polymers, and star-activated polymers (which are four). (Has the above functional groups). [0027]

The term "medical sealant" is a composition that is separated as a result of various disease states or surgical procedures and becomes anchored in place by mechanical and / or chemical means for encapsulating the tissue together. Say something. For example, medical sealants can be used to fill voids in hard tissue to bind blood vessels and other soft tissues together, provide a mechanical barrier to promote hemostasis, and one tissue surface can be another tissue. Tissue adhesion can be prevented by preventing contact and adhesion to the surface. [0028]

The term "interpenetrating polymer network (" IPN ") is a matrix formed from a reactive polymer or polymer, such as PEG, or a polymerized monomer, exhibiting an infinite molecular weight, and of different types. It is intended to describe a matrix that comprises a copolymer of (eg, high tensile strength fibers) and interpenetrates into a PEG polymer network to form two consecutive networks in the network. See: Sperling, L Het al., "Morphology and Mechanical Behavior of Interpenetrating Polymer Networks", in Polymer Networks, AJ Chompff and S. Newman, ed., Plenum Press, New York, New York (1971), pages 435- 449. The term "filler" is a bulk agent, low organic compound, polymer, fiber, resin, etc., which is a chemical composition for various reasons (most relevant to the present invention is to enhance tensile strength). Used to refer to what is added to. [0029]

(Composition component) The compositions of the present invention contain at least one polyfunctional synthetic polymer. The polyfunctional synthetic polymer, along with a tensile strength enhancer and other optional compositional components, can form a high strength matrix at the site of administration. When this polymer is crosslinked with other polymers and / or other composition components, the polymer components provide a "continuous gel phase" through the interconnection of flexible polymers. [0030]

If the composition contains a single polyfunctional synthetic polymer, the composition is designed so that the polymer remains non-reactive until administered. When the composition contains two or more polyfunctional synthetic polymers, the composition is generally designed as a two-part composition that reacts with each other when mixed at the site of administration. Alternatively, such binary compositions can also be prepared in a non-reactive state and subsequently activated before, during, or after administration. In addition, two-part compositions containing a single synthetic polymer species and a small amount (ie, less than half the size of the polymer) of cross-linking agent can be utilized. The forms of these and other compositions are described in more detail below. [0031]

In a preferred embodiment, the composition is a two-part composition, each containing a different polyfunctional activated synthetic polymer, so that the polymers react with each other at the site of administration. Form an IPN. As such, the compositions can be easily administered separately. [0032]

Such compositions can generally be represented by Equation I as follows: Compound 1-Xm + Compound 2-Yn Compound 1-Z-Compound 2 (I) Compound 1 has a plurality of functional groups (X) that react with compound 2 (m is 2 or more), and compound 2 has a plurality of functional groups (Y) that react with compound 2 (n is 2 or more), resulting in functionality. When the groups X and Y come into contact under the appropriate conditions, a covalent bond Z is formed. Thus, the functional groups X and Y can also be referred to as "reactive partners" with each other and collectively as "reactive pairs". Only one bond was formed between Compound 1 and Compound 2, as shown in Formula I for illustrative purposes only. However, if m + n is greater than or equal to 5, and the appropriate ratio of the two components is used as described elsewhere herein, the two compounds will have multiple bonds to each other. It forms and yields a three-dimensional polymer matrix. Preferably, both compounds contain four or more functional groups. This is because such polyfunctionality results in a gel matrix with greater binding strength as a whole. In certain preferred embodiments, each of the compounds is activated into four functionalities. [0033]

In another preferred embodiment, each of the compounds has 12 functional groups. Such compounds are formed by reacting a first tetrafunctional activating polymer with a second tetrafunctional activating polymer, where the functional groups of each of the two compounds are reactive pairs. It reacts with each other to form a functionally active polymer of "12-arms". An example of such a "12 arm" compound is dodeca-sulfhydryl-PEG (molecular weight 50,000), which is the core bound to four (external) four functional sulfhydryl-PEG molecules. Constructed from functional succinimide ester PEG. Such polymers vary in size from greater than 10,000 to greater than 100,000, depending on the molecular weight of the four functional activated polymer initiators. [0034]

Other types of polyfunctional polymers can be readily synthesized using conventional synthetic methods. However, care must be taken to produce a large number of arm products with consistent arm lengths and to avoid steric hindrance of functional groups. Thus, activated polymers suitable for use in the present invention may have a variety of geometric morphologies and arrangements. [0035]

(Polymer core) As mentioned above, each of the compounds has multiple functional groups X and Y. The rest of the non-reactive compound is considered to be its "core". To form a strong gel matrix, at least one of the two compounds has a polymer core. If one of the compounds contains a polymer core, the other compound can be a low organic molecule with multiple functional groups. However, for most applications it is preferred that both compounds have the same or different polymer cores. [0036]

The polymer core can be a synthetic polyamino acid, a polysaccharide, or a synthetic polymer. For some applications, the preferred polymer core material is a synthetic hydrophilic polymer. Suitable synthetic hydrophilic polymers include, among other things, polyalkylene oxides such as polyethylene oxide ((CH2CH2O) n), polypropylene oxide ((CH (CH3) CH2O) n) or polyethylene / polypropylene oxide mixtures ((CH2CH2O) n). -(CH (CH3) CH2O) n) can be mentioned. A particularly preferred synthetic hydrophilic polymer for a particular application is polyethylene glycol (PEG) having a molecular weight in the range of about 100 to about 100,000 (more preferably, about 1,000 to about 20,000). When the polymer core is polyethylene glycol, the polymer core even more preferably has a molecular weight in the range of about 7,500 to about 20,000 in general. Most preferably, polyethylene glycol has a molecular weight of about 10,000. [0037]

Polyfunctional activated polyalkylene oxides (eg, polyethylene glycol) are commercially available and are also readily prepared using known methods. For example, Chapter 22 of Poly (ethylene Glycol) Chemistry: Biotechnical and Biomedical Applications, edited by J. Milton Harris, Plenum Press, NY (1992); and Shearwater Polymers, Inc. Catalog, Polyethylene Glycol Derivatives, Huntsville, Alabama. See (1997-1998). Illustrative combinations of activated polymers for use as tissue sealants are: 4 functional PEG, pentaerythritol poly (ethylene glycol) ether tetrasuccinimidyl glutarate (molecular weight 10,000); And 4-functional PEG, pentaerythritol poly (ethylene glycol) ether tetrasulfhydryl (molecular weight 10,000). In both cases, these "four arm" PEGs are formed by ethoxylation of pentaerythritol. Here, each of the four chains has a molecular weight of about 2,500 and is then derivatized to introduce a functional group into each of the four arms. Also preferred are similar poly (ethylene glycol) -like compounds polymerized from diglycerol instead of pentaerythritol. [0038]

If only one of the polyfunctional activated compounds contains a polymer core, the other compound is a polyfunctional active low organic molecule. Such compounds include bifunctional discicin imidazole and maleimidyl compounds, as well as other well-known commercially available compounds (Pierce Chemical). Co., Rockford, Illinois). In addition, those skilled in the art can readily synthesize low molecular weight polyfunctional reaction compounds using conventional organic chemistry techniques. One such compound is shown in Figure 9a. This compound is pentaerythritol bound to four glutaric acids, each arm covered with N-hydroxy-succinimidyl ester (NHS). Similar compounds can be synthesized from inositol (six radial arms), lactitol (nine arms), or sorbitol (six linear arms). The end-covered reactive group can easily be sulfhydryl, maleimidyl, vinyl-sulfone, etc., instead of NHS. Polymers or small molecules can have any of the reactive end groups as long as they are reactive pairs in the composition (eg, NHS and SH, maleimide and SH, etc.). [0039]

In addition to polyalkylene oxides (eg, polymethylene, polyethylene and polypropylene), other polymers are also useful as core materials in the practice of the present invention. For example, polyester, polymethacrylate, polycaprolactone, polyalkene (eg, polybutadiene) may also be useful. Many of these materials have widespread utility in the pharmaceutical industry. For example, polycaprolactone is a component of nylon-6; polypropylene is a component of medical implants; polymethacrylate is found in polymethyl-methacrylate and poly-hydroxy-methyl-methacrylate, and of medical implants. It is a component; and polybutadiene is present in commercially available rubber. [0040]

An exemplary non-polyalkylene oxide-based composition consists of two different components, each containing a different penta-erythritol-based compound, so that the two compounds, when mixed together, will each other. To form a strong matrix. The first compound is penta-erythritol tetrakis (3-mercapto-propionate) (PESH-P) and the second compound is penta-erythritol tetraacrylate (PETA). A wide variety of similar molecular structures (four arms, radially symmetrical) can be synthesized based on pentaerythritol. The length of the molecular chain can preferably be extended using segments that are not alkyoxyl (eg, polyester, polymethylene, polyamide, or other material that is a component of known biocompatible polymers). [0041]

Other radially branched molecules (eg, glycerol or lactitol) can be utilized to construct the gel-forming material. The desired structure is a low molecular weight (molecular weight 350 to about 12,000) that is immiscible with water so that it remains liquid. Polymer gel-forming structures are also intended. It is preferred that such compositions are miscible with or dispersible in water, such as being compatible with the use of water as a liquid medium for delivery. [0042]

In addition to the branched molecules described above, the compositions of the invention can be formed from the linear molecules shown in FIG. Such a linear molecule can have a molecular weight of about 100,000 as long as it has a biodegradable element (O) and sufficient functional groups (R), as shown in FIG. .. [0043]

(Tensile strength enhancer) It is generally desirable to add a "tensile strength enhancer" to the composition in order to enhance the matrix strength. Such tensile strength enhancers are preferably glass with high tensile strength fibers (usually well above 37 ° C) of micron size (preferably 5-40 microns in diameter and 20-5000 microns in length). Has a transition temperature). [0044]

Suitable tensile strength enhancers for use according to the present invention include, among others, collagen fibers, polyglycolides and polylactide fibers, as well as other organic tensile strength enhancers and inorganic tensile strength enhancers. A particularly useful tensile strength enhancer is Vicryl (polyglycolide: polylactide, 90:10). The use of tensile strength enhancers, which are part of the broader "filler" category, is well known. For example, when crosslinked with peroxide, the "silicone gum" is weak and cheesy and has a tensile strength on the order of only 34 N / cm2. When properly mixed with the reinforcing filler, the tensile strength of these gums can be increased by as much as 50 times. Lichtenwalner, HK and Sprung, MN (Mark, HF, Gaylord, NG, and Bikales, NM, Encyclopedia of Polymer Science and Technology, Vol.12, p.535, John Wiley, New York, 1970). [0045]

Suitable tensile strength enhancers are enhancers that have inherently high tensile strength and can interact with the polymerized gel network by covalent or non-covalent bond. Tensile strength enhancers bind to the gel either mechanically or covalently to provide strength aids. The tensile strength of the reabsorbable suture of polyglycolide is approximately 89,000 N / cm2; the tensile strength of collagen fibers is 5000-10,000 N / cm2 (Hayashi, T. (Biomedical Applic. Of Polym. Mater). ., Tsuruta, T. et al., CRC Press, Boca Raton, Fla., 1993)). [0046]

(Reactive group and matrix bond) In the present invention, the bond Z includes a covalent bond formed when the functional groups X and Y react with each other. The functional group can be sulfhydryl, succinimidyl, acrylate, amino and the like. Thus, the linkage can be an ester, ether, disulfide, or the like. Other functional groups, their reactivity, and the bonds formed by them are well known in the scientific literature. For example, Bodanszky, M., Principles of Peptide Synthesis, 2nd Edition, pp. 21-37, Springer-Verlog, Berlin (1993); and Lundbland, RL, Chemical Reagents for Protein Modification, 2nd Edition, Chapter 6, CRC. See Press, Boca Raton, Florida (1991). Further examples of functional group pairs include, among others, sulfhydryl / acrylate, sulfhydryl / succinimidyl, amino / succinimidyl and amino / acrylate. [0047]

In addition to sulfhydryl-reactive compounds that form thioester linkages, various other compounds that form other types of linkages may be utilized. For example, a compound containing a methylimidate derivative forms an imide-thioester bond with a sulfhydryl group. Alternatively, sulfhydryl reactive groups that form disulfide bonds with sulfhydryl groups can be used (eg, orthopyridyl disulfide, 3-nitro-2-pyridenesulfenyl, 2-nitro-5-. Thiocyanobenzoic acid, 5,5'-dithiobis (2-nitrobenzoic acid), derivative of methane-thiosulfate, 2,4-dinitrophenyl cystenyl disulfide). In such cases, auxiliary reagents (eg, hydrogen peroxide or diter-butyl ester or azodicarboxylic acid) can be used to promote disulfide bond formation. [0048]

Yet another class of sulfhydryl reactive groups forms a thioether bond with the sulfhydryl group. Such groups include, among others, iodoacetamide, N-ethylmaleimide and other maleimides (dextranmaleimide, mono-bromo-biman and related compounds, vinylsulfonate, epoxides, derivatives of O-methyl-isourea, ethyleneimine, etc. Aziridine and 4- (aminosulfonyl-) 7-fluoro-2,1,3-benzoxadiazole) can be mentioned. [0049]

(Chain Extender) The functional group can be attached directly to the compound core or indirectly via a chain extender. Such chain extenders are well known to those of skill in the art. For example, a PCT that describes a "linking group" suitable for use as a chain extender in the compositions of the invention. See WO 97/22371. Chain extenders are useful for avoiding the sometimes problematic steric hindrance problem associated with the formation of direct bonds between molecules. Alternatively, a chain extender can be used to link several polyfunctional activating compounds together to make longer molecules. In certain preferred embodiments, chain extenders can also be used to alter the degradation properties of the composition after administration and the resulting gel formation. For example, the chain extender can be incorporated into one or both of the polyfunctional activated polymers to promote hydrolysis, prevent hydrolysis, or provide an enzymatic degradation site. The chain extender can also activate or suppress the activity of sulfhydryl groups or sulfhydryl reactive groups. For example, electron-withdrawing groups in one or more carbons of sulfhydryl groups are expected to reduce binding efficiency due to reduced nucleophilicity. Double bond carbons and carbonyl carbons are expected to have this effect. The bulky, adjacent groups of either partner are expected to reduce binding rate due to steric hindrance. The electron-withdrawing group adjacent to the reactive carbonyl of glutaryl-N-hydroxysuccinimidyl is expected to make this carbonyl carbon even more reactive with the sulfhydryl partner. [0050]

The chain extender may provide a degradation site (ie, a hydrolysis site). Examples of hydrolyzable chain extenders include, among others, α-hydroxy acids (eg, lactic acid and glycolic acid); poly (lactones) (eg, caprolactone, valerolactone, γ-butyl lactone and p-dioxanone (eg,). Dioxanone); poly (amino acid); poly (anhydrous) (eg, glutarate and succinate); poly (orthoester); poly (orthocarbonate) (eg, trimethylene carbonate); and poly (phosphoester). Examples of non-degradable chain extenders include, among others, succinimide, propionic acid and carboxymethylates. See, for example, PCT WO 99/07417. Examples of enzymatically degradable chain extenders include collagen. Examples include Leu-Gly-Pro-Ala; which is degraded by and Gly-Pro-Lys which is degraded by plasmin. [0051]

Other general principles to be considered when designing the compositions of the present invention are: When using high molecular weight structures, preferably have the above biodegradable linkages, and as a result, Fragments with a molecular weight greater than 20,000 are not produced during in vivo reabsorption. In addition, it may be desirable to impart sufficient charge or hydrophilicity to promote miscibility with water and / or solubility in water. Such hydrophilic groups can be easily introduced using known chemical syntheses as long as they do not produce gels that swell more than 3-5 fold and do not produce gels with low tensile strength. In particular, the polyalkoxy moiety can weaken the gel strength. [0052]

(Optimal composition ingredients) In addition to functional group-activated polymer compounds and tensile strength enhancers, the compositions of the invention may be included in one or both components of the two-component composition, or may be administered separately. Other compounds may also be included. In two embodiments, these compounds are covalent within the IPN itself by allowing the functional groups to be crosslinked with one or both of the activated compounds after the compounds have been mixed. Can be incorporated into. In another embodiment (eg, if the compound is not reactive with any functionally activated compound), the compound will be physically or ionicly associated with the matrix-forming compound after mixing. It can be administered in a fashion and thus be part of the matrix itself. [0053]

Additional compounds that can be added are glycosaminoglycans and proteins. Suitable glycosaminoglycans include, among others, hyaluronic acid, chitin), chondroitin sulfate A, B or C, keratin sulfate, keratin sulfate and heparin, and derivatives thereof. In another embodiment, the protein can be added for a variety of purposes. For example, collagen can improve the biocompatibility of the matrix, including the potential for cell colonization, accelerated wound healing, and so on. Collagen and proteins containing any amino group also contribute to the integrity of the structure of the matrix by being crosslinked into the matrix along with other matrix components. In particular, when PEG succinimidyl ester is used, the amide bond formed with collagen is more stable to hydrolysis than the bond formed by the reaction of succinimidyl ester with sulfhydryl. [0054]

Suitable proteins include, among others, collagen, fibronectin, gelatin, and albumin, as well as peptide fragments thereof. Particularly preferred is collagen, which can be in the form of non-fibrous, microfibrillar, fibrillar collagen. Type I and type III collagen isolated from bovine dermis or human placenta, or collagen prepared by recombinant DNA methods are suitable. See PCT WO 90/05755 for a description of suitable collagens and collagen derivatives. When collagen is added to the composition, it is important to adjust the concentration of other composition components to avoid precipitation. [0055]

Additional ingredients that may be added to the composition include antibiotics, growth factors, hemostatic proteins (eg, thrombin, fibrin, fibrinogen, blood factors, etc.), cells, genes, DNA, and the like. [0056]

(Optimal Rigidity Nanofiber) The above-mentioned optimum composition components can be added to the IPN of the present invention in various forms well known to those skilled in the art. For example, DNA can be added in the form of genomic DNA, oligonucleotides, polynucleotides, dinucleotides and the like. Similarly, proteinaceous components can be added as polypeptides, natural proteins, synthetic proteins, protein fragments and the like. [0057]

A particularly preferred structure for addition to the IPN of the present invention is nanofibers. Many of the above optimal components, and especially methylated collagen, can be easily formed in such structures. The term "nanofiber" is usually intended for fibers less than micron in length. In some cases, these fibers are so small that this form of collagen is considered to be "non-fibrous" in nature. In any case, when incorporated into an IPN (eg, PEG) that consists of a flexible polymer and has a relatively low (less than 0) glass transition temperature, it has a relatively high (greater than 25 ° C) glass transition temperature. The rigid nanofibers it has provide additional strength. [0058] [0058]

Collagen is a good example of rigid nanofibers, as collagen is formed from three polypeptide chains in a triple helix that hold each other by hydrogen bonds, but other polymers may also be suitable. For example, derivatives of other rod-like biopolymers (eg, tubulin and keratin) can be produced in the form of rigid nanofibers. Rigid nanofibers vary in structure as long as they are nanometer-scale rod-like polymers, water compatible, preferably have polar groups on their surface, and more preferably amino groups on their surface. It can be made from different materials. See, for example, Liu, G. et al., "Diblock Copolymer Nanofibers", Macromolecules 29: 5508-5510 (1996). [0059]

(Composition prescription) The compositions of the present invention contain two separate parts (or "ingredients"), which may be in liquid or solid form. In a preferred embodiment, both components are liquids, so that each can be easily applied separately to the administration site. Thus, one of these components is a dry powder that is mixed with the second component (in liquid form) either when each is sprayed separately onto the tissue or by mixing at the tissue site. Can be in the form of. It is also possible to have both components delivered to the site as a powder and mixed with buffer at the site of administration. [0060]

In an alternative embodiment, both components can be mixed together in a single aqueous medium (ie, eg, low pH buffer) in which they are both non-reactive. They can then be sprayed onto the tissue site with a high pH buffer, after which they react rapidly to form a gel. [0061]

The concentration of the reactive compound in each of the components of this composition always depends on a number of factors. For example, if the composition components are each 4-arm PEG (ie, PEG-PEG composition), the concentration of 20-25% by weight in each of the two components before mixing is about the modulus of elasticity (G'). Ten⁵~10⁶Produced after mixing the gel of dyne / cm2, this modulus of elasticity is sufficient for use as a surgical sealant. When methylated collagen and 4-arm succinimidyl PEG are used, concentrations of 2-4% and 0.2-0.4% yield gels with binding strength comparable to 10-15% PEG-PEG gels, respectively. When albumin is used as one of its constituents, concentrations of 30% and above achieve similar binding. Appropriate concentrations of this compound and other optional components in each component, and thus the relative concentration of matrix components in the final gel matrix, can be determined using conventional experiments to achieve the desired gelation time and gel strength. It can be easily optimized to achieve. Using the preferred 4-arm PEG described above, the synthetic polymer is generally present at a concentration of 2-50% (w / v), more preferably 10-25%. However, for some applications that require a high-strength matrix where the fluidity of the composition is not important, it is possible to use PEG at higher concentrations (eg 50-70%, more preferably 60%). Desired. [0062]

The liquid components of the compositions of the invention are prepared separately by addition to an activated synthetic polymer (in dry form or as a concentrated solution) to a liquid medium. Suitable liquid media include aqueous buffer solutions with a concentration of 0.5-300 mM (eg, monobasic sodium phosphate / dibasic sodium phosphate buffer, sodium carbonate / sodium hydrogen carbonate buffer, glutamate buffer or acetate buffer. ). Generally, sulfhydryl-reactive PEG is prepared in water or dilution buffer at a pH of about 5-6. A buffer with a pK of about 8 to 10.5 for preparing sulfhydryl-PEG components is useful for achieving fast gelation times of compositions containing sulfhydryl-PEG / SG-PEG mixtures. These include carbonate buffer, borate buffer and AMPO (3- [1,1-dimethyl-2-hydroxyethyl) amino] 2-hydroxy-propane-sulfonic acid). In contrast, when using a combination of maleimidyl PEG and sulfhydryl-PEG, a pH of about 5-9 is preferred for the liquid medium used to prepare sulfhydryl PEG. Particularly preferred compositions for hemostatic application to actively bleeding tissue sites include, as a first component, a mixture of maleemidyl PEG and succinimidyl PEG, and as a second component, sulfhydryl PEG. Such compositions produce gels that are more biodegradable and have excellent gel time when compared to compositions having only maleimidyl PEG or succinimidyl PEG alone. [0063]

The pH of the aqueous buffer solution used for each of these two (or more) compositional components is a delivery process using conventional optimizations to achieve a final pH that leads to rapid gelation. It should be adjusted without causing immediate gelation that interferes. For example, both amino PEG and sulfhydryl PEG require a basic pH to enhance nucleophilicity. The effect of pH on gel time is discussed in the Examples below. [0064]

(Gel strength) The compositions of the present invention are formulated to exhibit excellent binding strength. This generally means that the compositions of the present invention exhibit at least 10% and preferably 20% of the burst strength of the cyanoacrylate "Superglue". [0065]

(Example formulation made from COH102 / 206 / methylated collagen, + glass wool or Vicryl) Addition of methylated collagen to a PEG formulation (eg, COH102 / COH206 as described in Example 2 below) greatly strengthens the gel and resists the gel from swelling in saline buffer or water. .. Since the gel does not swell, the gel continues to have an inherent tensile strength sufficient to remain bound to the fibrous filler (eg, glass wool, Vicryl thread or silk suture). [0066]

The combination of COH102 / COH206 and methylated collagen can be thought of as an "interpenetrating polymer network". This is due to the presence of this methylated collagen and PEG reagent as a uniform clear solution prior to gelation at pH 4-5. The ability of both PEG and derivatized collagen to coexist in solution was a very unexpected discovery. At the time of gelation (induced by increasing pH), COH102 and COH206 react via their respective reactive groups to form a thioester bond. This bond results in a network of infinite molecular weight, which is characteristic of polymer gels. COH102 alone and COH206 alone form gels, but such gels are relatively weak. In the presence of methylated collagen, the amino groups on this collagen can also react with COH102, forming amide bonds and forming a three-way gel of PEG and collagen, which are covalently bonded. Becomes a mesh. [0067]

This addition of methylated collagen produces a gel that is robust enough to support the fibrous filler. The gel is not strong enough to hold the fibrous filler, so the gel separates under a tensile load and high tensile strength is not achieved. Table 12 below shows that the binding performance is reduced when other polymeric or particulate molecules replace methylated collagen. Thus, this methylated collagen is a relatively rigid molecule that can impart increased strength to relatively diluted PEG hydrogels. Flexible polymers (eg, hyaluronic acid, polylysine, and chitosan) produce relatively high molecular weight and relatively viscous solutions, but apparently do not have the same stiffness and are therefore more capable of absorbing tensile stresses. It is less likely to pass off these stresses to the fibrous filler. [0068]

In polymer terms, PEG is a flexible polymer chain with a low glass transition temperature (the temperature at which this polymer is flexible instead of elastic and hard); methylated collagen has a higher glass transition temperature. (Ie, the methylated collagen loses its rigidity and becomes flexible at about 33-35 ° C, and at this temperature, this methylated collagen "melts" in a flexible chain molecule. Form some gelatin). The result is the reinforcement of the former by the latter, as it forms an interpenetrating polymer network (IPN) between the low glass transition polymer (PEG) and the high glass transition polymer (methylated collagen). If the PEG is to be reinforced with another flexible polymer (eg, hyaluronic acid), which also has a low glass transition temperature (ie, is flexible at the temperature of interest), no reinforcement is achieved. (For example, Polymer Networks, Structural and Mechanical Properties, AJChompff, S.Netman ed., Plenum Press, NY, 1971, p. 435, Sperling, LH, Huelck, V. and Thomas, DA, "Morphology and mechanical behavior of interpenetrating polymer networks"). [0069]

The increased binding strength of the final formulation clearly depends on several factors: 1) The PEG reagents COH102 and COH206 covalently bond to form a hydrogel; this gel also provides excellent binding to the tissue. Shown, this is clearly due to the succinimidyl ester of COH102, which can bind to amino groups on the tissue (see Table 14); 2) Methyled collagen is an interpenetrating polymer containing a PEG gel network. It forms a network and reinforces the PEG gel network; and 3) passes off a pull load onto a fibrous filler (eg, a polylactide / glycolide fiber or a glass filter). Therefore, the addition of this fibrous filler is important in achieving the desired final level of gel tensile strength. [0070]

It is intended that other collagen derivatives that remain soluble (molecular, non-fibrous) at pH 4-7 may also confer the same desired properties. For example, other esters of the carboxyl group of collagen (eg, ethylated collagen, propylated collagen, or benzylated collagen) can also function like methylated collagen. Fibrous collagen is less desirable; because it forms beaded, lumpy domains in the gel and does not impart increased tensile strength. Collagen derivatives (eg, succinylated collagen) that are not soluble at pH 4 or may not contain free reactive amino groups are also preferred. Because they do not provide the desired properties. Rigid synthetic polymers are almost exclusively water-insoluble crystal structures that exist as micron-scale masses and do not form a tight interpenetrating network with the required PEG hydrogels. [0071]

(Use and administration) The compositions of the invention are generally delivered to the site of administration in such a manner that the two (or more) individual components of the composition come into contact with each other for the first time at or shortly before administration. .. Therefore, the compositions of the present invention are preferably delivered to the site of administration using a device that delivers the two components separately. Such delivery systems typically include a two-component single outlet spray device or a two component double outlet spray device. Alternatively, the two components can be delivered separately using any type of controllable drainage system, or they can be delivered and administered by hand in the form of separate pastes, liquids or dry powders. Can be mixed together by hand at the site. Many devices adapted for the delivery of two-component tissue sealants / hemostatic agents are well known in the art and they can also be used in the practice of the present invention. [0072]

Yet another method of delivering the compositions of the invention is to make the two reactive components (or a single reactive component in the case of a sulfhydryl-containing component designed to form a zirspide bond) in an inert form. It is to be prepared as either a liquid or a powder. Such compositions can then be activated by application with an activator after or shortly before application to the tissue site. In one embodiment, the activator is a buffer solution having a pH that activates the composition once mixed with it. See Example 7 for a description of a sulfhydryl-containing PEG composition that is maintained at a low pH until administration and then mixed with a high pH buffer at the site of application to initiate gelation. [0073]

Yet another way to deliver this composition is to prepare a preformed sheet and apply the sheet itself to the administration site (see Examples). [0074]

The compositions of the present invention can be used in a variety of different pharmaceutical applications. In general, the compositions described herein may be adapted for use in any tissue manipulation application for which synthetic gel matrices are currently utilized. For example, the compositions of the present invention are useful as tissue sealants, in tissue enhancement, in tissue repair, as hemostatic agents, in the prevention of tissue adhesion, in the provision of surface modifications, and in drug / cell / gene delivery applications. .. Those skilled in the art will readily develop suitable dosing protocols for use using any composition with known gel strength and gelation time, based on the principles described herein and well-known scientific principles. Can be decided. A more detailed description of some specific applications is given below. [0075]

(Tissue sealant and adhesive) In a preferred application, the compositions described herein can be used in medical situations that require a coating or sealing layer to prevent leakage of gases, liquids or solids. This method applies both components to damaged tissue or organ, 1) sealing blood vessels and / or other tissues or organs to stop or minimize blood flow; 2) leakage of air. Sealing chest tissue to stop or minimize; 3) Sealing gastrointestinal or pancreatic tissue to stop or minimize leakage of fecal or tissue contents; 4) Leakage of urine Seal the bladder or urinary tract to stop or minimize the leakage of CSF; 5) Seal the hard membrane to stop or minimize the leakage of CSF; and 6) the skin or the skin to stop the leakage of serum. Accompanied by sealing the serous tissue. [0076]

These compositions can also be used to bond tissues such as small vessels, nerves or skin tissue together. This material either 1) applies it to the surface of one tissue and then the second tissue can be quickly pressed against the first tissue, or 2) juxtaposes those tissues nearby. It can then be used by applying this substance. In addition, this composition can be used to fill spaces in soft and hard tissues produced by disease or surgery. [0077]

(Surgical adhesion) A preferred application is a method of reducing the formation of adhesions after a surgical procedure in a patient. This method involves applying the material onto damaged tissue or organ, either by spraying both components together or by applying a premixed component or a preformed solid implant. .. This component reacts together to form a strong matrix on the tissue surface. This medical procedure includes gynecological, abdominal, neurosurgical, cardiac, and orthopedic applications. [0078]

(Drug delivery) A preferred application is a method of topically applying a biologically active substance to a patient. The active substance, together with the two components, can be delivered such that the substance can be formed in situ, or the active substance can be part of a preformed implant. The active substance can either be released via a diffusion controlled process or can be covalently attached to its components such that the active substance is released as the resulting hydrogel decomposes. [0079]

This biologically active substance can be any of a variety of organic and inorganic substances, including proteins, sugars, and nucleic acids. Specific examples include enzymes, antibiotics, antitumor agents, cytokines, local anesthetics, hormones, anti-angiogenic drugs, antibodies, neurotransmitters, psychoactive drugs, drugs that affect the reproductive organs, and therapeutics. Examples include oligonucleotides. [0080] [0080]

(Modification of graft) A preferred application is a method of applying a coating to the implant so as to affect the surface properties of the implant or help adhere the implant to the tissue surface. Ingredient coats can be applied to 1) vascular grafts, stents to minimize or stop the leakage of blood or serum from these devices; 2) to reduce or stop excessive fibrosis. It can be applied to a catheter or chest implant; 3) It can be applied to an artificial patch or mesh to minimize excessive fibrosis and help adhere the implant to the tissue surface. [0081]

(Delivery of cells or genes) A preferred application of this composition is to encapsulate cells or genes, which include substances or synthetic DNA from natural sources, RNA and their individual antisense forms, thereby delivering them to the desired site. That is. These cells may include mesenchymal stem cells, epithelial cells and neuroectoderm cells. The cells may be of allogeneic or heterologous origin. [0082]

(Example) (Example 1) (Preparation of two-component penta-erythritol-based tissue sealant composition) Base 1.08 g of penta-erythritol tetrakis (3-mercapto-propionate) (molecular weight 489, "PESH-P" in Figure 2a) and 1.0 g of penta-erythritol tetra-acrylate (molecular weight 352 ("PETA" in Figure 2b)). Mix together in the presence of 5-10 g of poly-oxypropylene tri-amine (T403 in Figure 2c, Jeffamine, Texasco Chemical Co., Houston, Texas). [0083]

All reactive species are liquids. PESH-P and PETA are not miscible in water. Therefore, PETA is heated to about 40 ° C to form a liquid before mixing PESH-P and T403. Within 1-5 minutes after mixing, gelation begins, depending on the level of T403. The bond formed between PESH-P and PETA is shown in Figure 3. The gel is cured for several hours and then hydrated at 37 ° C for 1 hour. After that, the tensile strength of this gel is 0.88 +/- 0.3MPa. When such gels are left in physiological saline (pH 6.7), they are stable for more than 40 days and swell only about 20%. The burst strength data show only moderate adhesion to conceal the grindate. This is expected. This is because there is no chemical bond of sulfhydryl or acrylate to the protein using the PETA-P / PESH mixture. A burst pressure of 20-40 mmHg was observed in the three burst strength tests. [0084]

(Example 2) (Tensile strength of various compositions) (Materials and methods) Penta-erythritol polyethylene glycol ether tetra-thiol (molecular weight 10,000) (COH206), penta-erythritol polyethylene glycol ether tetrasuccinimidyl-glutarate (molecular weight 10,000) (COH102), and penta-erythritol polyethylene glycol ether tetra Amino (molecular weight 10,000) ("COH204) was purchased from Shearwater Polymers, Inc. (Huntsville. Alabama). Cyanoacrylate" Superglue "was purchased from a retailer. Gelatin (70-100 Bloom) was purchased from Sigma (Saint Louis, Missouri). Sulfonate-ethylene glycol biscucin imidyl succinate (S-EGS), dimethylsber imidate (DMS), and dysuccini imidazole glutarate (DSG), Pierce Chemical Purchased from Company, Rockford, Illinois. Polyethylene glycol ("PEG") (molecular weight 200) di-acrylate ("PEG-di-acrylate"); PEG (molecular weight 1,000) di-methacrylate ("PEG-di-methacrylate"); and 2-hydroxy-ethyl methacrylate ("PEG-ethyl methacrylate") HEMA ") was purchased from Polysciences, Inc., Warrington, Pennsylvania. Polypropylene (PPO) (Molecular Weight 540) Diacrylate (PPO-Di-acrylate); PPO (Molecular Weight 230) Bis-2-aminopropyl Ether (PPO-Di-Amino 2,30); PPO (Molecular Weight) 2,000) Bis-2-aminopropyl ether ("PPO-di-amino 2,000"); Poly tetrahydrofuranbis (3-aminopropyl) ("PTMO") (Molecular weight 350) ("PTMO350"); PTMO (Molecular weight 1,100) ( "PTMO 1,100"); PESH-P (Molecular Weight 489); PETA (Molecular Weight 352); and Meta-Potasie Dipersite, Aldrich Chemical Purchased from Company, Milwaukee, Wisconsin. Ammonium persulfate was purchased from Biorad, Inc., Richmond California. Methylated collagen was prepared from purified bovine dermal collagen according to a modified method from US Pat. No. 4,164,559 (see Example 7). The structures of PTMO (also referred to as polytetramethylene oxide di-amine), PPO di-acrylate and PPO di-amine are shown in FIGS. 4a, 4b and 4c, respectively. [0085]

The gel was prepared as follows: (a. COH102 / COH206) 100 mg of COH102 was dissolved in 400 μl of 0.5 mM sodium phosphate (pH 6.0). 100 mg of COH206 was dissolved in 400 μl of 300 mM sodium phosphate (pH 7.5). The two solutions were mixed in a beaker and poured into a mold approximately 8 x 0.5 x 0.5 cm. Gelation occurred in 2-3 minutes. The sample was left at room temperature until dry. The dried matrix was removed from the mold and hydrated at 37 ° C for 1 hour prior to the tensile strength test. [0086]

(b. COH102 / COH204) This sample was prepared as described in a, but COH206 was replaced with COH204. [0087]

(c.PETA / PESH-P) This sample was prepared as described in Example 1. [0088]

(d. Gelatin gel) 20% gelatin in sodium phosphate / sodium carbonate buffer (pH 9.6) with different compounds as shown in a below and 10-20 mol of active amino per mol of gelatin. Assumed and mixed using the stoichiometric level of the other compound. [0089]

(e. COH102 / PPO-di-amino 2,000 / PEG-di-acrylate) Dissolve 615 mg of COH102 in 923 μl ethanol and 246 μl PPO-di-amino 2,000 and 246 μl PEG-di as described in a. -Mixed with acrylate. [0090]

(f. PETA / PPO-di-amino 230 / PPO-di-amino 2,000) 500 μl PETA, 630 μl PPO-di-amino 230 and 150 μl PPO-di-amino 2,000 were mixed together as described in a. .. [0091]

(g.COH102 / PTMO) This gel was prepared as described in e, but PPO-di-amino 2,000 was replaced with PTMO 1,100. [0092]

(h. Cyanoacrylate) The adhesive was extruded onto water and cured immediately. [0093]

(i.HEMA) 1.3 ml HEMA and 64 μl PEG-di-acrylate were dissolved in 600 μl of 150 mM sodium phosphate buffer (pH 7.4) and mixed with 40 mg ammonium persulfate in 100 μl water. The mixture was heated to 60-80 ° C for 4 hours. [0094]

(j. COH102 / COH206 / Methylated Collagen) 25 mg of methylated collagen, 100 mg of COH102 and 100 mg of COH206 were dissolved in 1 ml of 0.5 mM sodium phosphate (pH 6.0). [0095]

(Measurement of tensile strength) The edges of the dry gel were fixed and then the central region of all samples was rehydrated in 37 ° C saline buffer (pH 6.7) for approximately 1 hour prior to testing. The matrix was then stretched to the cut point in an Instron Model 4202 test instrument (Instron, Inc., Canton, Massachusetts) fitted with a 100 N load cell. Peak loads were recorded and converted to extreme stress using the measured cross-sectional area of the sample at the cut point. Distort the data = ΔL / L₀Was also plotted as (stress) vs. (strain). Where ΔL is the stretched portion, and L₀Is the length of the original sample. [0096]

[table 1]

<img file="JP2003508564A_D0001.tif" /> (Example 3) (Comparison with high-strength adhesives based on COH102 and COH206, as well as adhesives prepared from PETA, PESH-P and GLYC-20HS) (wrap up) Several types of gels were examined as possible suture alternative formulations. Gels based on the penta-erythritol derivative showed high adhesiveness, but poor adhesive strength. The 3-arm succinimidylglycerol-PEG-based gel showed low adhesive strength but good adhesive strength. Gels based on 60% aqueous (w / v) COH102 / COH206, to which various fibrous substances (eg, fibrous insoluble collagen, polyglycolide sutures and glass wool) are added, have good adhesion. It showed both strength and good adhesive strength. [0097]

High-strength medical adhesives are of interest as suture substitutes in the closure of surgical incisions. In particular, gels formed from PETA and PESH-P have been shown to have about 10 times greater tensile strength than the tensile strength formed from a 20% (w / v) solution of COH102 and COH206. .. When the PETA-PESH-P gel was supplemented with a fibrous or fine particle polymer, gels with even higher tensile strength were observed. [0098]

This experiment showed the adhesive properties of PETA / PESH-P and related gels, as well as the adhesive properties of the formulation with 60% (w / v) COH102 and COH206, to which collagen and other polymers are added. And both tensile properties are described. The properties of gels formed from 3-arm glycerol succinimide (NOF Corp., Japan) and the above reagents are also described. [0099]

(Materials and methods) PETA, PESH-P and penta-erythritol tetrakis (3 mercaptoacetate) (PESH-A), polyethylene, surface activated 180μ particle size, and polybutadiene, epoxy functionalization, epoxy EW260, Aldrich Chemical Co. (Milwaukee, Wisconsin) I bought it from. GLYC-20HS (poly-oxyethylene glyceryl ether) succinimidyl succinate 2600mw) and DEPA-10H (poly-oxyethylene bis-amine 1040mw) were obtained from NOF Corporation (Japan). COH102 and COH206 were reagent grades of Shearwater Polymers (Huntsville, Alabama). Polyethylene-co-acrylate-succinimidate (PE-AC-S), Aldrich Chemical Synthesized from polyethylene-co-acrylate (approximately 400K molecular weight with 5% acrylate) purchased from Company (Milwaukee, Winsconsin). Kensey-Nash Insoluble Collagen (Semed F) was purchased from Kensey-Nash Corporation (Exton, Pennsylvania). Collagen Matrix, Inc. (Franklin Lakes, New Jersey) supplied insoluble type 2 collagen. The Prolene 7-0 suture was manufactured by Ethicon Corporation. Crude fibrous collagen sheets include Prior, JJ, Wallace, DG, Harner, A. and Powers, N. "A sprayable hemostat containing fibrillar collagen, bovine thrombin, and autologous. It was excised from the same crude fibrous bovine dermis collagen used for the burst test described in "plasma", Ann.Thor.Surg.68,479-485 (1999). These collagen sheets were used as a tissue model for further study. Smaller fibrous collagen was prepared from reprecipitated pepsin-digested bovine dermis collagen produced by Collagen Aesthetics, Inc. (Palo Alto, California). I bought glass wool from VWR Corporation. The poly-glycolide suture (uncoated) ("Dexon S") was made by Davis and Geck. [0100]

Gel formation for measuring tensile strength is described in Example 1 above. The equipment used for the burst test is described by Wallace et al. (Supra). Approximately 1 ml of the total formulation was sprayed onto a crude fibrous collagen sheet substrate or spread with a spatula and cured. After the formulation reached a relatively rigid elastic solid phase (no longer sticky) texture, hydraulic pressure was applied and the pressure to rupture the seal was recorded as mmHg. [0101]

A 60% gel of COH102 and COH206 was prepared as follows: COH102 was dissolved in S-buffer (0.5 mM sodium phosphate, pH 6.0) at 60% (v / w), and COH206 was 60%. Dissolved in 300 mM sodium phosphate (pH 7.5 or 8.9), or in 117 mM sodium phosphate, 183 mM sodium carbonate (pH 9.6) (PC buffer). In some cases, the above ratio of phosphate to carbonate was varied to give pH 9.44 for faster cure times. The pH used in each case was determined by the desired gelation rate. Various additives were added to such basic formulations. For example, Kensey-Nash and smaller fiber size collagen were added at 28 mg / ml final gel; glass wool was added at 25 mg / ml; and polyglycolide suture was added at 40 mg / ml. [0102]

(Results and Discussion) The results are discussed below and are shown in Tables 2, 3 and 4 below. 60N / cm²Greater tensile strength is considered "strong". Burst strength greater than 50mmHg is considered "good adhesion". [0103]

The gel of PETA and PESH-P showed good tensile strength (Example 1). However, they showed poor adhesion (&lt;50 mmHg burst pressure) when tested for adhesion to hydrated simulated tissue (crude fibrous collagen sheet) in burst tests. As shown in Table 2 below, this formulation was then applied to the water-soluble GLYC-20HS and DEPA-10H, or COH102 and COH206 pairs (these alone in an aqueous medium, which are good for collagen sheets. Modified to contain (given adhesion). These materials had good tensile strength (manual evaluation), but again had poor adhesion to the collagen sheet. Gels formed from GLYC-20HS and DEPA-10H also had poor adhesion in the absence of water in the formulation. Since these reagents are water soluble, different results can be observed when dissolved in aqueous buffer. [0104]

However, when GLYC-20HS was the major component on a mass basis, the gel was weak but showed good adhesion in burst tests. Using these particular combinations of ingredients, it seemed possible to achieve either high tensile strength or high adhesive bond (but not both). [0105]

[Table 2]

<img file="JP2003508564A_D0002.tif" /> The ability of succinimidyl derivatized polyethylene (PE-AC-S) to act as an effective tensile strength enhancer for PETA-PESH-P gels and COH102 / 206 gels was also evaluated (Table 3). This material did not increase the tensile strength of these gels, probably because it was not a stretched filament (ie, the aspect ratio was not high enough). [0106]

[Table 3]

<img file="JP2003508564A_D0003.tif" /> Table 4 also summarizes the results of COH102 and COH206 + Kense-Nash fibrillar collagen, which showed enhanced tensile strength against 20% and 60% (w / v) gels of COH102 / 206 alone. In addition, the COH102 / COH206 / collagen formulation had good adhesive binding to the collagen sheet. Other additives (eg, hide grindate and Prolene 7-0 suture) also enhanced gel strength relative to the control. Some fillers (eg, small fibrous collagen, polyethylene, and polybutadiene) did not exhibit tensile strength enhancing properties. Finally, some fillers or combinations thereof (eg, glass wool and insoluble collagen + poly-glycolide suture) have the tensile strength found in cyanoacrylates (385 N / cm).²), A significant increase in tensile strength was shown (Example 1). Although limited burst strength data were collected, the data predict that all of these COH102 / 206 (60%) formulations are highly adhesive to the collagen surface and therefore also to the tissue. It is confirmed that it will be done. [0107]

As shown in Table 4, P-HEMA hydrogels are described in Santin, M. et al., "Synthesis and characterization of a new interpenetrated poly (2-hydroxyethylmethacrylate) -gelatin composite polymer", Biomaterials 17, 1459-1467; And gelatin-PEG-di-acrylate is Nakayama, Y. and Matsuda, T. "Photocurable surgical tissue adhesive glues composed of photoreactive gelatin and poly (ethylene glycol) diacrylate", J. Biomed. Biomat. Res. (Appl. Biomater. ) 48,511-521 (1999). [0108]

[Table 4]

<img file="JP2003508564A_D0004.tif" /> Figure 5 shows a tensile test of a piece of COH 102/206 (60%) + 28 mg / ml Collagen Matrix Collagen + 40 mg / ml Dexon S uncoated polyglycolide suture (4-0). The measured tensile strength is 700 N / cm²It was higher. Measurements were interrupted when the sample began to slide from the test device (tilted downwards). [0109]

(Example 4) (Collagen sheet complex containing active PEG) (wrap up) COH102 and COH206 can be dissolved together in ethanol without reaction. Evaporating and drying this ethanol solution may leave a reactive PEG powder. When placed in PC buffer (pH 9.6), the powder immediately gels. This finding was used to develop an adhesive sheet. A collagen sponge and a crude fibrous collagen sheet coated with the sponge were impregnated with an ethanol solution of COH102 and COH206. After drying, the reactive PEG sheet was able to bind to the crude fibrous collagen sheet. After 10 minutes, a burst intensity of about 300 mmHg was observed at pH 7. [0110]

An adhesive sheet consisting of a crude fibrous collagen lining (prepared as described in Example 3) to which a collagen sponge was attached was also prepared. The sponge layer was impregnated with reactive PEG powder. When such a sheet was placed on a second crude fibrous collagen sheet moistened with pH 9.6 buffer and dried, a strong bond was formed in about 30 seconds. The bond of the adhesive sheet to the second collagen sheet was measured by a burst test. Burst intensities up to 500 mmHg were observed. [0111]

(Materials and methods) Crude fibrous collagen sheets were prepared as described herein. A collagen sponge was attached to this collagen sheet by extruding a slurry of reconstituted fine fiber collagen to the top. A reconstituted collagen slurry was prepared by diluting expired Zyderm II collagen (65 mg / ml protein, Collagen Aesthetics, Inc., Palo Alto, California) with water to achieve a dispersion of 20 mg / ml. Approximately 7 cm per 1 ml slurry²Covered the sheet. After extruding the collagen slurry onto a crude fibrous collagen sheet, the complex was frozen and then lyophilized. After lyophilization, the porous collagen sponge remained adherent to the collagen sheet. An alcohol solution of COH102 and COH206 was added dropwise to the dry sponge. Alcoholic solutions of COH102 and COH206 were prepared by adding 200 mg of COH102 and 200 mg of COH206 to 1.8 ml of absolute ethanol (dried on molecular sieves), respectively. Dissolution of both PEG compounds in ethanol was achieved by gentle heating and stirring (up to about 40 ° C). The two solutions were then mixed. Approximately 10 cm using 1 ml ethanol solution²Covered the sheet. The sponge-collagen sheet moistened with a PEG solution in ethanol was then dried briefly in a reduced pressure chamber. The bulk solvent was blown off and the sheet was transferred to a lyophilizer for several hours to remove solvent residue. The dry sheet was stored at -20 ° C under dry argon. [0112]

A burst test was performed to test the binding to the second crude fibrous collagen sheet. A collagen test sheet was moistened with PC buffer (pH 9.6) and secured in sample cells. Additional buffer (0.1-0.2 ml) was applied to the sheet. An active PEG sponge sheet (prepared as above; 1.2 x 1.2 cm) was then placed on the test collagen sheet with the sponge side down (size of defect: 2 mm). The two sheets were held in place by hand for 2-3 seconds to ensure good contact and hydration of the active PEG sheets. The two sheets were then left to incubate at room temperature until water pressure was applied from below. The pressure required to rupture the seal was recorded. [0113]

(result) Table 5 shows that very high bond strengths can be achieved at pH 9.6 compared to control sheets that do not contain active ingredients. High bond strength is observed even for a short time. The long-term stability of this bond was qualitatively tested. A pair of bonded sheets was removed from the burst tester and placed in 100 ml of water (33 ° C). The binding was still complete, and clearly strong, even after 30 hours. [0114]

[Table 5]

<img file="JP2003508564A_D0005.tif" /> Table 6 shows the effect of lowering the pH of the hydration buffer. The strength of the bond was pH two dependent, and a much weaker bond was observed at pH 7.5. However, the bond strength at lower pH was improved by allowing longer (3-10 minutes) bond formation. This may be acceptable for a particular application. [0115]

The form of failure was not always clear. In some cases, the test sheet was removed from the test clamp without damaging the bond. These results have not been reported. If the test sheet was tightly clamped, it appeared that there was a failure between the active sheet and the test sheet. In some cases this active sheet was lifted and water spouted out of the hole. In other cases, water appeared on the edge of the test cell, but the source was unclear. [0116]

[Table 6]

<img file="JP2003508564A_D0006.tif" /> * pH 8.9 buffer was 300 mM sodium phosphate; pH 7.5 buffer was 300 mM sodium phosphate. [0117]

(Example 5: Further study on adhesive sheet containing active PEG) (Summary) Adhesive sheets were prepared from COH102 / COH206, Zyderm collagen, and Kensey-Nash (Semen F) collagen. Such sheets showed good burst strength when bound to crude fibrous collagen sheets. [0118]

A shear failure binding test was developed. In this test, the bond strength of the active PEG adhesive sheet was comparable to that of cyanoacrylate (SUPER GLUE). This experiment describes the preparation of adhesive sheets from biocompatible materials, including transplant grade Kensey-Nash collagen. The present specification also reports the comparative results of various formulations in the shear failure test. [0119]

(Materials and methods) Active sheets were prepared from crude fibrous collagen sheets, reconstituted collagen and COH102 and COH206 as described in Example 4. [0120]

The active sheet from Kensey Nash collagen was prepared as follows: Collagen (4 g) was added to 200 ml of deionized water adjusted to pH 2 with HCl. Stirring with an overhead impeller was maintained for 6 days. The fibers swelled and dispersed, forming an almost binding mass. The slurry was poured into a 9 cm diameter polystyrene heavy duty boat (approximately 15 ml per boat to cover the bottom). The poured slurry was left to dry in an incubator at 33-37 ° C. The seat was formed on the surface of a heavy boat. In some cases, such sheets were coated with Zyderm collagen (20 mg / ml slurry in water), frozen and lyophilized. Then such a dry sheet coated with a thin sponge-like layer of Zyderm collagen Was activated with COH102 and COH206 (each component had a concentration of 10% w / v). These were referred to as "acidic KN sheets". An alternative procedure utilized a dry Semed F sheet (dried from an acidic slurry) coated once with a Zyderm collagen sheet and incubated in water for 3 hours to adjust to pH 9-10. Additional sodium hydroxide was added to hold the solution and sheet at pH 9-10 as needed. The sheet was removed from the solution and dried at 37 ° C overnight. The sheet was then recoated with Zyderm collagen at 12 mg / ml, frozen, lyophilized, and finally immersed in active PEG in ethanol to give a dried active sample, which was "based". It was called "sex KN sheet". [0121]

A shear failure test on the active sheet was developed using the device shown in FIG. Crude fibrous collagen sheets, either activated with PEG or prepared as usual for disruption testing, were cut to a size of 1 x 3 cm (referred to as "strips"). A plastic tab or adhesive tape was secured over one end to cover an area of 1 cm square of the strip. Hide grindate this plastic tab It was fixed to strip) with cyanoacrylate glue. A hole about 2 mm in diameter was drilled in the plastic tab or tape. Approximately 400 μl of PC buffer with pH 9.6 was placed over an area of 1 square cm of collagen sheet. The active PEG / collagen sheet was then placed on a buffered skin grind and held at light pressure for 30 seconds. In this way, the combined surface was 1 square cm. Hooks were then placed through the holes at the ends of each strip and the combined strips were hooked onto the appropriate supports. Metal washers (average weight per washer: 1.63 g) were added to the lower hooks to incrementally increase the shear force on the bonded surface. The weight required to pull and release the bonded surface was recorded. The properties of the bond surface and variations in the bond formulation were studied. This includes using a control strip containing collagen sponge on the surface but not active PEG and binding the cyanoacrylate to the skin grinding strip. [0122]

One set of samples included the binding (without active PEG) of one untreated skin grinding strip to the collagen sponge coated strip. Another set of samples was made with the same pair of strips, but this time in addition to the strips juxtaposed with the COH102 / COH206 mixture and gelled. The gelled mixture was as follows: 10 mg COH102 and 40 μl pH 6.0, 0.5 mM phosphate buffer 10 mg COH206 and 40 μl pH 8.9, 300 mM phosphate buffer. [0123]

Such a mixture gels in about 15 seconds, which is enough time for the gel mixture to be added to one strip, the second strip pressed onto it, and wait 30 seconds for gelation to occur. I left it. [0124]

For strip binding with cyanoacrylate, test strips were first squeezed with PC buffer (200 μl per square cm area). A second strip coated with cyanoacrylate was then pressed onto the first strip. As with the active PEG sheet, juxtaposition was maintained for a binding time of 30 seconds using only mild pressure. [0125]

(result) The acidic KN sheet and the basic KN sheet were subjected to the destructive test: 1 square cm piece was placed on the test skin grinding sheet already moistened with PC buffer. The acidic KN sheet showed very low burst intensities (2 mmHg and 4 mmHg). The basic KN sheet gave similar results to the previous results with the coated collagen sheet (324 mmHg, 286 mmHg and 311 mmHg). [0126]

Collagen sheet binding was compared under several conditions using the shear failure method (Table 7). Very poor binding strength was observed for the sheets bound by COH102 / COH206. The active PEG sheet showed strength similar to that observed with cyanoacrylate. [0127]

[Table 7]

<img file="JP2003508564A_D0007.tif" />* Final bond strength was not determined; in both cases, the collagen sheet was destroyed or the hook failed to secure, nor was the bond region. [0128]

(Example 6: Method for synthesizing high-strength bioadhesive) This example describes the synthesis of a copolymer of polyethylene glycol (PEG) and polycaprolactone (PCL) containing a pendant N-hydroxysuccinimidyl group. [0129]

(Scheme 1 (Fig. 7)) Bifunctional PEGs of various molecular weights are selected and bound to PCL by controlled ring-opening polymerization and added to the fixed and selected molecular weight PCLs (Hedrick et al, Macromolecules 1998, 31,8691-8705). .. The terminal OH group in PCL was converted to glutalyl succinimidyl (Abuchowski et al, Cancer Biochem 1984, 7, 175-180). Lysine ethyl ester was added to give a linear polymer with repeating PEG, PCL and lysine segments. The lysine segment has a pendant carboxyl ester. This ester is hydrolyzed to give a free carboxyl group (Nathan et al. Macromolecules 1992, 25, 4476-4484). This carboxyl group is then converted to glutalyl succinimidyl (Abuchowski). [0130]

This scheme produced a straight chain with a pendant reactive group (NHS), a degradable linkage (PCL), and a water-soluble group (PEG). Polymer stiffness can be increased by adding long PCL elements to short PEG elements. However, PEG must be long enough to ensure water solubility. This can be determined empirically. [0131]

(Scheme 2 (Fig. 8)) A similar straight chain with pendant reactive groups but with PMMA (polymethylmethacrylate) repeating segments. Bifunctional PEGs of varying but selected molecular weight were activated with 2-bromo-2-methylpropionyl Br, and controlled segments of PMMA were grown at each end. The terminal contains a Br group, which can be converted to a -OH group by standard organic chemistry methods and further derivatized to glutalyl NHS and reacted with a lysine ethyl ester as in Scheme 1. , And finally the pendant carboxyethyl ester was hydrolyzed and liberated to carboxylate and derivatized to glutalyl NHS, and a similar copolymer with different physical properties was given again (eg PMMA has high tensile strength). Known to have strength). Sufficient PEG functionality may be provided due to its water solubility. It may be necessary to construct a PEG-PCL-PMMA-lys repeat unit, resulting in PCL providing the required biodegradable binding. The construction of such a segment is described in Hedrick. [0132]

Another approach for strengthening polymers is described in Liu, et al., Macromolecules, 29: 5508 (1996). PMMA blocks containing monomers such as .2-cinamoylethyl methacrylate (CEM) can be synthesized. The following structures are exemplary: water-soluble blocks (eg, hydroxyethyl methacrylate (HEMA)) and some amino and carboxy-containing monomers (eg, aminoethyl methacrylate (AEM) and mono-2- (acryloyl) ethyl succi). Nate (ARS)), and diblock polymethacrylates containing water-soluble blocks (eg, polyMMA-co-CEM) are synthesized. When the polymer is present in water or in polar media, the MMA CEM block avoids water and associates with separate regions. Lie, et As outlined in al., It is possible to change the ratio of water-soluble blocks to water-insoluble blocks in the polymer to obtain structures in different regions. The areas that meet and avoid water can be micelle-shaped (spherical), rod-like or plate-like. This is because the polymer chain contains a plurality of water-insoluble blocks. If precise conditions are found for rod-shaped poly MMA-CEM blocks, these blocks can be crosslinked with light of appropriate wavelength and intensity. The rod is then fixed and a water soluble polymer block (in this case poly HEMA-AEM-ARS) surrounds the rod. Nanometer-scale rods with reactive amino groups can then be obtained. The dimensions of such rods are expected to be 1-30 nm in diameter and hundreds to thousands of nm in length. Such rods should also function to strengthen the polymer gel network. This is because methylated collagen is believed to enhance the PEG network as described elsewhere in this application. [0133]

Hedrick describes how to make linear PEG-PCL-PMMA segments of controlled size but without reactive groups and with multiple pendant reactive groups. Nathan describes a method of binding PEG to lysine to give a repeating polymer having a reactive pendant group (but no degradable ester bond and no addition of PCL and PMMA for strength). Abuchowski provides a method for making glutalyl succinimidyl PEG. WO99 / 17417 and the references cited therein describe PEG polylactide-acrylates to be polymerized by photoradical and free radical processes. These materials do not mention adding PMMA for strength. Moreover, they make no mention of pendant NHS or other groups for covalent bonds (not driven by free radicals). [0134]

(Example 7: COH102 / 206 / methylated collagen and fibrous filler glass wool or Vicryl) (material) (a. Methylated collagen) Methylated collagen was prepared by modifying the procedure of Miyata et al, US Pat. No. 4,164,559. Dispersant of bovine pepsinized reconstituted collagen at 0.02 M sodium phosphate, 0.13 M NaCl pH 7.2 (3% w / v) (McPherson et al., Collagen) (Prepared by the method of Rel. Res. 5,119-135, 1985) was extruded onto the glass surface in a thin layer at room temperature and dried. Methanolic HCl was prepared by adding 104 g anhydrous sodium sulfate and 10.7 ml of concentrated HCl to 1300 ml of anhydrous methanol, and was tightly capped and left for 2 days. The dried collagen was cut into 1 x 5 cm strips and added to methanolic HCl (200 ml methanolic HCl: 1 g dry collagen) in an enclosed container and gently shaken at 20 ° C. for 3 days. The methanolic HCl was carefully decanted and removed, and the collagen was filtered through a calcined glass funnel to remove the remaining methanol. Complete methanol removal was completed overnight under reduced pressure. Methylated collagen was redissolved in distilled water and the pH was adjusted to about 4-6. The amount of water was calculated to achieve a final protein concentration of about 31 mg / ml. Samples of methylated collagen dissolved at lower protein concentrations were reconcentrated by short lyophilization to remove water. It was a completely transparent material containing dissolved methylated collagen, contained no fibers or opal glow, and had a viscous gel-like consistency. Preparations that still contained hazy or insoluble components (due to incomplete methylation of collagen) behaved poorly in adhesive formulations, swelled excessively, and had poor binding strength. The gel shown was produced. [0135]

(b. Adhesive without filler) For 0.5 ml of adhesive, 50 mg dry powdered COH102 (4-armed tetraglutarylsuccinimidyl polyethylene glycol, 10K, Figure 9a) and 50 mg dry powdered COH206 (4-armed tetrathiol polyethylene glycol, 10K; Figure 10b) was mixed with 400 mg methylated collagen at 31 mg protein / ml. Both PEG components were dissolved in an aqueous collagen solution to give a clear viscous liquid. The solution is spread over the tissue site with a spatula and it hardly flows under gravity. To cure the adhesive, 20-50 μl of buffer (either 134 mM sodium phosphate, 166 mM sodium carbonate, pH 8.9; or PC buffer, pH 9.6) was added to its surface. The buffer did not dilute the gel, but slowly leached. In 3-5 minutes, the surface of the gel was recognizablely cured. [0136]

For binding strength studies under hydrated conditions, the gel and substrate were cured on a bench for 20 minutes and then soaked in 50 mM sodium phosphate, 130 mM sodium chloride, pH 6.7, 37 ° C for over 2 hours. did. The bond strength test was performed in a tensioning device. [0137]

(c. Adhesive with filler) Vicryl is a (90:10) copolymer of glycolic acid and lactic acid, sold as a transplantable mesh by Ethicon Corporation (Polyglactin 910; Sommerville, New Jersey). [0138]

To this methylated collagen was added 19 mg of 1-2 cm long Vicryl threads unwound from a transplantable Vicryl mesh. In some cases, Vicryl fibers as short as 0.3 cm were also used. The yarn and viscous gel were mixed and then the PEG component was added as described above. The application and hardening to the tissue was as described above. The respective amounts added to the other fillers and 0.5 ml of adhesive were: glass wool 9 mg; fibrous collagen (Semed F collagen, Kensey-Nash Corporation) 8 mg; Dexon S (poly) Glycolidelactide suture "4-0"), 10 pieces 1 cm long; elastin fiber (bovine cervical ligament, 0.25 ~ 10 mm, Elastin Products Co., Inc, Owensville, MO), 40 mg; slenderless steel fiber (Bekaert Fiber Technologies, Marietta, GA), 14-28 mg (fiber is water or 1N Washed with HCl and removed with polyvinyl alcohol coating); Polylactide / glycolide (65; 35, 40-75,000 molecular weight, Aldrich Chemical Co., Zheng, J., and Homsby, PJ, Biotechnol.Progr.15,763-767 (1999) 2-4 diameter microparticles prepared by the method of), polylactide / glycolide microparticles prepared from 25 mg). [0139]

(d. Adhesive with methylated collagen replaced by another drug) Various long chain molecules were tested (eg, hyaluronic acid (Rooster Comb, Sigma Chemical Co., St. Louis, Missouri), chitosan (Sigma), and polylysine (Sigma)). For hyaluronic acid, the formula was: COH102, 50 mg, COH206, 50 mg, Vicryl, 14 mg, and 400 μl of hyaluronic acid, 2% (w / v) in water, pH adjusted to 4 for chitosan; for chitosan, In the same formula, 400 μl of 1% chitosan (w / v) was used in water at pH 4-5. For polylysine, COH 102, 40 mg, COH 206, 30 mg were dissolved together in 50 μl water; polylysine hydrobromide, 330 K, 40 mg was dissolved in 60 μl water; the two solutions were mixed together and 7 mg Vicryl fibrils were added. In addition, the polylactide / glycolide particles prepared as described above were tested as an alternative for methylated collagen; 16.5 mg of particles were suspended in 300 μl of water, and 50 mg. It was mixed with COH 102, 50 mg, COH 206, and 14 mg Vicryl. All gels were cured with a layer of pH 9.6 buffer as described above. [0140]

(e. Adhesive without filler and without methylated collagen) COH102 was dissolved in water at 20% (w / v); COH206 was dissolved in pH 8.9 buffer at 20%. The two solutions were quickly mixed and extruded into the site. Gelation was performed during drying. [0141]

(Mechanical test) The bond strength of the adhesive formulation was applied to the three types of tissues or tissue substitutes shown in FIG. The collagen membrane (Fig. 10a; sausage packaging; The SausageMaker, Inc., Buffalo, New York) was washed with isopropyl alcohol and water to remove impurities in lipids and salts and dried. Bonding of membranes with 1-3 mm overlap and 1 cm width was performed by spreading the adhesive over the sheet. The adhesive was cured on a bench for 20 minutes, then soaked at 37 ° C for 30 minutes to 2 hours, then pulled off in an Instron model 4202 test instrument (Canton, Massachusetts) using a 100N load cell. Coupling of the porcine carotid arteries (10b, Pelfreeze, Rogers, Arkansas) was also performed in the terminal-to-terminal structure. The severed carotid segment was flanked (46 mm diameter) and the adhesive spread; no supporting suture was applied. Incubation and testing were the same as described for collagen membranes. [0142]

Hair-removed bovine skin pieces were purchased from Spear Products, Inc., Quakertown, Pennsylvania for binding of bovine skin strips (10c). The pieces were 2-3 mm thick and almost uniform. A strip 0.4 cm wide was cut from a hide piece using a single-ended laser blade. The cutting strips were flanked at the ends and bonded by spreading 0.25 ml of "CT003" adhesive or a few drops of cyanoacrylate. Incubation and testing were the same as those described for collagen membranes. Figure 8 below shows that COH102 / COH206 / methylated collagen was superior in binding strength to unfilled formulas a and b when filled with glass wool (formula c) when tested on collagen membranes. I showed that. In fact, the bond strength was comparable to that obtained with commercially available cyanoacrylate adhesives (Table 9). Medical grade cyanoacrylate (Dermabond) formed stronger bonds even with collagen membranes (5.2 ± 1.9N force for 7 measurements) [0143]

[Table 8]

<img file="JP2003508564A_D0008.tif" />* The collagen membrane was torn, but the sealant bond was still intact. [0144]

[Table 9]

<img file="JP2003508564A_D0009.tif" /> Table 10 below presents data on the addition of different fillers, Vicryl yarns, to COH102 / 206 / methylated collagen. When a substrate such as cowhide or carotid artery was used, the substrate did not tear and the bond strength value was representative of the strength of the adhesive junction itself. Typically, these bonds were not adhered. That is, the tensile strength of the adhesive gel itself remained intact and was not a limiting factor. The binding strength observed in saline at 37 ° C was also comparable to that seen with cyanoacrylate for binding the same set of tissue substrates (Table 9). [0145]

[Table 10]

<img file="JP2003508564A_D0010.tif" />* Cowhide strip, 0.5 cm width, porcine carotid artery, 0.3-0.5 cm diameter, collagen membrane: sausage packaging, 0.2 mm thick, 1 cm. [0146]

(Curing different fillers) Table 11 presents the results for various filler materials. The test was performed on cowhide strips and soaked in saline for 2 hours at 37 ° C. The fibrous material appeared to be more effective than the spherical particles. Bonding of the filler to the gel was very important for improving strength. Collagen polyethylene glycol filaments were waxy and did not adhere to the gel. Therefore, despite their high aspect ratio, they were not effective fillers. [0147]

[Table 11]

<img file="JP2003508564A_D0011.tif" /> (Effect of substituting methylated collagen with other polymer molecules) Table 12 shows that no test material provided binding strength comparable to formulations containing methylated collagen. [0148]

[Table 12]

<img file="JP2003508564A_D0012.tif" /> (Effect of cross-linking) Table 13 below shows that when the gel was formed from other types of cross-linking reactions, its adhesive strength and bond strength were affected when tested against cowhide after incubation at 37 ° C. .. Material 1 was formed from COH206 and hydrogen peroxide. It oxidizes adjacent sulfhydryl groups into disulfide bonds. The gel forms rapidly and the gel can be supplemented with methylated collagen and Vicryl, but after a few hours in saline buffer the gel becomes very weak. This Vicryl fiber is easily pulled out. Material 2 utilizes the reaction of a sulfhydryl group from COH206 with a double bond of a 4-arm vinyl sulfone derivative of PEG (10K, Shearwater Polymers; Figure 11). The expected reaction, Michael-type addition, formed a thioether bond. Such gels had adequate tensile strength, but had poor adhesion to cowhide after incubation in saline. Materials 3 and 4 are COH204 (4 arms, tetrafunctional amino PEG, 10K, Shearwater) Polymers) were included; the amino functional group probably reacted with the succinimidyl ester of COH102 to form an amide bond (Fig. 12). These gels were comparable in performance to the gels formed from COH102 and COH206. (For proper reaction in the presence of methylated collagen, COH204 had to be titrated to pH 2-4 during reagent mixing; its pH increased with the addition of curing buffer. , Amino group reaction is possible). The presence of succinimidyl ester appeared to be important for achieving the best adhesion to the tissue substrate and for the good tensile strength of the gel. Other groups that react with amines, such as aldehydes (aldehydes conjugated to multi-arm PEG), are also expected to be effective adhesive forming reagents. [0149]

[Table 13]

<img file="JP2003508564A_D0013.tif" /> (Persistence of binding under hydration conditions) Table 14 shows that the adhesives formed from COH102, COH206 and also COH204 form a bond that uses cowhide that lasts for long periods of immersion in saline buffer (37 ° C). Such stringent hydrolysis conditions mimic the in vivo environment. Bond decay was observed after more than 100 hours of hydrolysis. The attenuation of the bond strength was considered to be due to the hydrolysis of the carboxyl-ester and thio-ester (Fig. 13) network bonds. COH102 is a glutaryl succinimidyl ester; even after the reaction of the succinimidyl ester with the terminal carboxyl, the carboxyl ester that binds the glutaryl moiety to the major PEG chain remains; this bond and thioester. The bond could be hydrolyzed. [0150]

[Table 14]

<img file="JP2003508564A_D0014.tif" /> (Formulas associated with succinimidyl esters and low molecular weight compounds with amino or thiol reactive groups) Table 15 shows the bond strength of the low molecular weight PEG derivative on the cowhide strip when adhering, as well as the bond strength when supplemented with collagen and Vicryl. GLYC-20HS is a trifunctional succinimidyl succinate (2600 molar weight) of 3-armed PEG formed from a glycerol core (NOF Corporation, Japan). COH201 is tetra-amino (4-armed PEG, 2000 molar weight) (Shearwater Polymers). This polymer was Vicryl packed, which appears to have a small effect on bond strength. The following ratios were used: methylated collagen, 500 μl (22 mg / ml in water, 2707-30B); GLYC-20HS, 48 mg; 60 μl 60% aqueous solution titrated to pH 1-2 with COH201, 6M HCl; Vicryl yarn, 26 mg .. [0151]

[Table 15]

<img file="JP2003508564A_D0015.tif" /> (Burst test on collagen disc and slit defect in carotid artery) The ability of stickiness intended for use in surgical applications is often measured by the ability to seal fluid leaks. Two types of leak or fluid pressure tests were used: (a. Burst test on collagen disc) Using the device depicted in Figure 15, the collagen mat was mounted on a brass platform and secured with a second brass ring through the first brass ring. A hole was drilled in the lower brass platform and connected to a line filled with water. Water was flushed with a syringe pump at 5 ml / min. The shunt line was connected to the pressure gauge. The test collagen mat was also perforated (2 mm diameter hole). The adhesive preparation was applied to the mat to cover the perforations. The adhesive could cure for 3 minutes (longer if it needed to result in hardening to a hard rubber), then hydraulically applied. The pressure required to rupture the seal was recorded. Instead of cyanoacrylate, a small (4 x 4 mm) piece of collagen mat was glued to the lower perforated mat. [0152]

(b. Slit defect on the carotid artery) FIG. 16 illustrates a pressured carotid artery model. Pig arteries (Pel Freeze Biologicals, Rogers, Arkansas) were connected to a water line. Water was flushed with a peristaltic pump. The end of the line had a flow limit placed on the end of the line so that pressures up to 10 psi and above could be applied to the line by increasing the pump speed. A first intact artery was placed in the system and subjected to hydraulic pressure to ensure that it could withstand the desired pressure without leaks. Arterial compartments without side branches are preferred; sometimes leaking branches are clamped to stop leakage. A slit approximately 2 mm long was cut laterally to the circumference in four compartments in the artery (Fig. 16). The suture is then mimicked using the amputated artery, and the suture is a support suture (stay). suture) was applied. The cut sites were then glued everywhere in an attempt to seal them. Immediately prior to applying the adhesive, buffer (134 mM sodium phosphate and 166 mM sodium carbonate, pH 8.9) was applied to the arterial tissue. The adhesive soul was washed with a further small amount of this buffer to cure the gel. After a curing time of 8 minutes, the bonded joints were subjected to hydraulic pressure. Pressure was increased by 1 psi increase and held at each pressure for 1 minute before further increase. Leaks were recorded as positive if they drip faster than one drop every 10 seconds. [0153]

Table 16 shows the burst intensities of COH102 / 206 / methylated collagen / Vicryl for holes of various diameters (against collagen membranes with a cure time of 8 minutes; cured with pH 8.9 buffer; using a spatula. Spread 0.5 ml sample over the entire hole). A hole with a diameter of 5 mm is the largest defect that can be considered in surgical applications. Because supportive sutures were used to close this largest defect, and the largest spacing between such sutures was estimated to be 5 mm. Even in such large holes, the adhesive was able to maintain pressure near or greater than the maximum expected (ie, 4 psi) in hypertensive patients. The third data entry emphasizes the need to have good gel hardening at the interface between the gel and the collagen disc. Addition of curing buffer to this surface prior to application improves this short-term binding. [0154]

(Table 16: Burst intensity of COH102 / 206 / methylated collagen / Vicryl) [0155]

[Table 16]

<img file="JP2003508564A_D0016.tif" /> Table 17 shows data for the closure of large slit defects in the carotid artery (4 x 2 mm slits cut into an artery with a diameter of 4-6 mm). The COH102 / 206 / Methylated Collagen / Vicryl formulation is comparable in performance to cyanoacrylates. Note that poorer results are observed for smaller arteries that extend at lower pressures. [0156]

(Table 17: Burst intensity test for porcine carotid artery) [Table 17]

<img file="JP2003508564A_D0017.tif" /> (Example 8) (Derivatives of penta-erythritol and trimethylolpropane functionalized with thiols and acrylates, + fillers (eg Vicryl)) This type of adhesive was prepared from a liquid reagent that was not diluted with water (no aqueous buffer added). The coupling reaction involving the addition of a thiol group to the activated double bond of the acrylate is shown in FIG. Table 19 below shows the reactive pairs TPETA (tri-methylolpropane ethoxylate tri-acrylate (molecular weight 912), Aldrich Chemical Co .; Figure 18) and TMPE-SH (tri-methylol-propaneethoxylate tri-thiol). (Molecular weight 1140) (Fig. 19) (Trimethylol-propane ethoxylate (molecular weight 1100), synthesized from Aldrich Chemical Co.)) is shown. To drive this reaction, a base was needed to convert the thiol to a thiolate anion. T403 (Tri-Amino 3-Arm Propylene Oxide (Molecular Weight 970) (Texaco Chemical) Co., Fig. 2c) was used. The gel formed was strong enough and only Vicryl was added as a filler. No methylated collagen was added (methylated collagen (dissolved in water) was unlikely to remain in solution with liquid TPETA and TMPE not diluted with water). This prescription was as follows: (Table 18) (TPETA / TMPE-SH / T403 / Vicryl) [0157]

[Table 18]

<img file="JP2003508564A_D0018.tif" /> TMPE-SH and T403 were mixed together, then Vicryl yarn was added, and finally TPETA was added and mixed with the other ingredients. The mixture was spread on a moist test substrate with a spatula as described above. The mixture became sticky in about 3 minutes and was a hard gel in 5 minutes. After this, the test mixture was immersed in saline after 20 minutes as described above. The gel time was freely controllable by the amount of T403 added. [0158]

Table 19 shows that binding was achieved for all three tissue substrates, but the binding strength was lower than that of the COH102 / 206 formulation. Bonding failures have always been adhesives-these gels showed high tensile strength but did not adhere well to the tissue. However, strong adhesion to tissue is not required for certain clinical applications. This feature was particularly pronounced for collagen membranes. This chemical coupling reaction apparently does not provide a suitable mechanism for covalent attachment to tissues, as with COH102 (where succinimidyl esters can, in principle, react with amino groups on the tissue). .. Free sulfhydryl groups on tissue proteins can react with acrylate groups in TPETA, but such groups are relatively rare on proteins. [0159]

[Table 19]

<img file="JP2003508564A_D0019.tif" /> Table 20 shows TPETA / TMPESH formulations, which show persistent binding after immersion in saline, but these are during prolonged immersion in saline at 37 ° C. , Shows only loose binding to small pieces of cowhide. This is believed to be due to the relative stability of the thio-ether bond compared to the bond present in the COH102 / 206 gel. The molar ratio of the materials used was 0.59: 0.52: 0.031 and TPETA: TMPE-SH: T403: Vicryl was 15 mg and mg / 0.39 ml. [0160]

[Table 20]

<img file="JP2003508564A_D0020.tif" /> (Example 9) (Tensile strength of hydrated adhesive material) The adhesive formulations described herein usually lack adhesiveness-its adhesive strength does not appear to be a limiting factor. The data below compare the tensile strength of this adhesive itself to other adhesives and hydrogels. [0161]

(Measurement of tensile strength after immersion in physiological saline at 37 ° C for 2 hours) The adhesive is formed into a roughly columnar shape and cured on a polystyrene weighing boat at room temperature for about 20 minutes. The sample was then immersed in water at 37 ° C for 2 to 2 1/2 hours. The sample was removed from water, bottled, and taped to either end using cyanoacrylate adhesive. The sample was then attached to an Instron tensile tester, taped to one end and stretched until broken. With respect to the cross-sectional area of the sample at the point of failure, the final tensile strength is N / cm²Shown in. In some cases, the stress-strain curve as the strain ΔL / L₀Calculated using (where ΔL = LL₀And L is the sample length at time t, and L₀Is the original sample length). To prepare a control sample consisting of cyanoacrylate alone, cyanoacrylate was drawn on the liver of the cow. After about 20 minutes, the cyanoacrylate was cured into small pieces. It was removed from the liver (which is a very fragile tissue) and the remaining liver particles were removed. In this way, clean, relatively uniformly cured pieces of cyanoacrylate polymer were obtained. [0162]

The data shown in Table 21 show that adhesives containing tensile strength enhancers exhibit higher tensile strength than hydrogels without such enhancers. Acrylate-thiol-class adhesives containing Vicryl (eg, TPT-SH: trimethylolpropane tris (3-mercaptopropionate), FIG. 20) had higher tensile strength than cyanoacrylate alone. Prescription without tensile strength enhancer is about 8 ~ 20N / cm²Has a tensile strength of. COH102 / 206 (20%) is a simple hydrogel. It had much lower tensile strength than the corresponding formulations containing methylated collagen and Vicryl. The measurement was performed after immersion in physiological saline at 37 ° C for 2 hours. [0163]

[Table 21]

<img file="JP2003508564A_D0021.tif" /> (Example 10: Linear copolymer having pendant reactive group) The purpose of this experiment was to find a means for synthesizing linear polymers with the pendant reactive groups shown in FIG. A further objective was to allow not only pendant reactive groups, but also monomer subunits or polymer blocks that could impart tenacity to gels formed from polymers could be inserted into the linear polymer chains. Thus, this experiment is an extension of Example 6 above. [0164]

All of the polymers prepared were in the poly-acrylate / methacrylate family. Linear polymers were prepared according to the procedure outlined in "Composite poly (2-hydroxyethylcrylic) membranes as rate-controlling barriers for transdermal applications", Biomaterials 18,527-533, 1997 by Sun, YM et al. The monomers contained in this polymerization are HEMA (hydroxy-ethyl-methacrylate), MMA (methylmethacrylate), ARS (mono-2 (acrylicoxy) ethylsuccinate), and AEM (2-amino-ethyl-methacrylate). (Fig. 14). [0165]

Polymerization to form a poly HEMA with carboxyl and amino side chains using the following components: HEMA, 1260 μl; ARS, 240 μl; AEM, 348 mg; azobis-isobutyro-nitrile, 54 mg; and dry ethanol, 34 ml. , 75 ° C for 4 hours. The precipitation reaction was carried out using dry petroleum ether. [0166]

Polymerization was initiated using azobis-isobutyro-nitrile in dry ethanol with vigorous stirring. AEM (not sufficiently soluble in ethanol) is first dissolved in water (50% w / v), the solution is prepared to pH 2-4 using HCl, and solutions of other components in ethanol. Added. The entire reactant was flushed with nitrogen and conducted in a screw cap bottle in a water tank. Free radical inhibitors in HEMA and MMA are passed through a column of aluminum oxide 90 (Active neutral, activity 1,70-230 mesh, EM Science, Gibbstown, New Jersey) immediately prior to addition to the reaction mixture. Removed by. After 4 hours at 70-80 ° C, the solution was cooled, filtered to remove insoluble AEM and other particulates, and the polymer was precipitated by adding 10 ml petroleum ether. The polymer coagulated on the surface of the glass beaker and transferred the ethanol solution used. The polymer was dried under reduced pressure for several days. [0167]

Table 22 shows that the polymer is apparently acquired and rapidly forms a gel with COH102 for the first two samples in the table, suggesting the presence of an AEM moiety in the polymer chain. In the case of early polymer preparation, the relative amounts of amino and carboxyl functional groups on the polymer were inferred from the titration curve. For each 20 HEMA units, there were approximately one AEM subunit and two ARS subunits along the chain. See Figure 23. [0168]

No harder polymer gel was formed than the hydrogel formed from COH102 and COH206 alone. Copolymers with high MMA are chalky and easily crumble. MMA was originally selected because pMMA itself is a hard, water-insoluble polymer. Therefore, these classes of polymer chains with pendant reactive groups are clearly not suitable for use in the present invention. Further modifications of these acrylic acid / methacrylic acid monomer ratios, or the introduction of other monomers of the allylic acid family, can result in polymers with the desired properties, and such optimization studies are described herein. It can be easily carried out using the teachings of. [0169]

(Table 22) (Properties of synthetic polyacrylic acid / methacrylic acid polymer) [0170]

[Table 22]

<img file="JP2003508564A_D0022.tif" />1. Approximately 50% (w / v) solution or dispersion of polymer in pH 9.6 buffer (described above) was mixed with 50% (w / v) solution of COH102 in water. Gel formation from seconds to hours was scored as a "+". 2. Molecular weight compared to pMMA standard, HPLC on series PL gel, molecular sieving column in dimethylformamide solvent at 1.0 ml / min flow rate. [0171]

Numerous modifications can be made to the above system without departing from its basic teachings. Although the present invention has been described in significant detail with reference to one or more specific embodiments, those skilled in the art will make modifications to the embodiments specifically disclosed in this application. It is understood that these modifications and improvements are still within the scope and purpose of the invention, as set forth in the claims above. All publications, patents, and patent applications disclosed herein are specifically and individually so that such publications, patents, or patent applications are incorporated herein by reference. Incorporated herein by reference, as shown in.

[Simple explanation of drawings]

[Figure 1]

FIG. 1 shows the structure of linear and branched polymers with biodegradable bonds (O) and functional groups attached to them (R). [Fig. 2a]

FIG. 2a shows the structure of pentaerythritol tetrakis (trimercaptopropionate, PESH-P). [Fig. 2b]

Figure 2b shows a pentaerythritol tetraacrylate (PETA) saw. [Fig. 2c]

Figure 2c shows the structure of T403 bound to COH102 or acrylate. [Fig. 3]

FIG. 3 shows the bond formation between PESH-P and PETA. [Fig. 4a]

FIG. 4a shows the structure of polytetramethylene oxide diamine that binds to COH102 or acrylate. [Fig. 4b]

FIG. 4b shows the structure of polypropylene oxide diacrylate attached to a thiol group or an amino group. [Fig. 4c]

FIG. 4c shows the structure of polypropylene oxide diamino that binds to COH102 or acrylate. [Fig. 5]

FIG. 5 shows the results of Example 3. [Fig. 6]

FIG. 6 shows an apparatus for the shear failure test of Example 5. [Fig. 7]

FIG. 7 shows the synthetic scheme of Example 6. [Fig. 8]

FIG. 8 shows the synthetic scheme of Example 6. [Fig. 9]

FIG. 9 shows the structure of COH102 (9a) and COH206 (b). [Fig. 10]

Figure 10 shows the types of devices that are useful for measuring tensile strength. [Fig. 11]

FIG. 11 shows the structure of a 4-arm vinyl sulfone derivative of PEG. [Fig. 12]

FIG. 12 shows the formation of amide and carboxyester bonded PEG conjugates from succinimidyl-glutaryl-PEG and aminoPEG. [Fig. 13]

FIG. 13 shows the formation of thioester-bound PEG conjugates from succinimidyl PEG and thiol PEG. [Fig. 14]

FIG. 14 shows the structures of HEMA (hydroxyethyl methacrylate), MMA (methyl methacrylate), mono-2- (acrylic oxy) ethyl succinate and 2-aminoethyl methacrylate. [Fig. 15]

FIG. 15 shows a device that is useful for testing the burst strength of collagen membranes. [Fig. 16]

FIG. 16 shows a device (compressed carotid model) useful for testing the burst intensity of repaired arterial slit defects. [Fig. 17]

FIG. 17 shows the formation of a thioether bond from the reaction of an acrylate with a thiol. [Fig. 18]

FIG. 18 shows the structure of TPETA (trimethylolpropane ethoxylate triacrylate). [Fig. 19]

FIG. 19 shows the structure of TMPE-SH (trimethylolpropane ethoxylate trithiol). [Fig. 20]

FIG. 20 shows the structure of TPT-SH (trimethylolpropane tris (3-mercaptopropionate)). [Fig. 21]

FIG. 21 shows the structure of glycerol-dimethacrylate. [Fig. 22]

FIG. 22 shows the structure of polyhydroxyethyl methacrylate-co-aminoethyl methacrylate-co-mono-2- (acrylic oxy) ethyl succinate.