USRE37950E1 - Biogradable in-situ forming implants and methods of producing the same - Google Patents
Biogradable in-situ forming implants and methods of producing the same Download PDFInfo
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- USRE37950E1 USRE37950E1 US09/536,579 US53657900A USRE37950E US RE37950 E1 USRE37950 E1 US RE37950E1 US 53657900 A US53657900 A US 53657900A US RE37950 E USRE37950 E US RE37950E
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
- A61K9/00—Medicinal preparations characterised by special physical form
- A61K9/0012—Galenical forms characterised by the site of application
- A61K9/0019—Injectable compositions; Intramuscular, intravenous, arterial, subcutaneous administration; Compositions to be administered through the skin in an invasive manner
- A61K9/0024—Solid, semi-solid or solidifying implants, which are implanted or injected in body tissue
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L15/00—Chemical aspects of, or use of materials for, bandages, dressings or absorbent pads
- A61L15/16—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons
- A61L15/42—Use of materials characterised by their function or physical properties
- A61L15/44—Medicaments
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L15/00—Chemical aspects of, or use of materials for, bandages, dressings or absorbent pads
- A61L15/16—Bandages, dressings or absorbent pads for physiological fluids such as urine or blood, e.g. sanitary towels, tampons
- A61L15/42—Use of materials characterised by their function or physical properties
- A61L15/62—Compostable, hydrosoluble or hydrodegradable materials
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- A—HUMAN NECESSITIES
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- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L24/00—Surgical adhesives or cements; Adhesives for colostomy devices
- A61L24/001—Use of materials characterised by their function or physical properties
- A61L24/0042—Materials resorbable by the body
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L24/00—Surgical adhesives or cements; Adhesives for colostomy devices
- A61L24/04—Surgical adhesives or cements; Adhesives for colostomy devices containing macromolecular materials
- A61L24/046—Surgical adhesives or cements; Adhesives for colostomy devices containing macromolecular materials obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L26/00—Chemical aspects of, or use of materials for, wound dressings or bandages in liquid, gel or powder form
- A61L26/0061—Use of materials characterised by their function or physical properties
- A61L26/009—Materials resorbable by the body
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
- A61L27/14—Macromolecular materials
- A61L27/18—Macromolecular materials obtained otherwise than by reactions only involving carbon-to-carbon unsaturated bonds
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- A—HUMAN NECESSITIES
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- A61L27/00—Materials for grafts or prostheses or for coating grafts or prostheses
- A61L27/50—Materials characterised by their function or physical properties, e.g. injectable or lubricating compositions, shape-memory materials, surface modified materials
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
- A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
- A61F2/02—Prostheses implantable into the body
- A61F2/30—Joints
- A61F2002/30001—Additional features of subject-matter classified in A61F2/28, A61F2/30 and subgroups thereof
- A61F2002/30003—Material related properties of the prosthesis or of a coating on the prosthesis
- A61F2002/3006—Properties of materials and coating materials
- A61F2002/30062—(bio)absorbable, biodegradable, bioerodable, (bio)resorbable, resorptive
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- A—HUMAN NECESSITIES
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- A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
- A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
- A61F2/02—Prostheses implantable into the body
- A61F2/30—Joints
- A61F2002/30001—Additional features of subject-matter classified in A61F2/28, A61F2/30 and subgroups thereof
- A61F2002/30316—The prosthesis having different structural features at different locations within the same prosthesis; Connections between prosthetic parts; Special structural features of bone or joint prostheses not otherwise provided for
- A61F2002/30535—Special structural features of bone or joint prostheses not otherwise provided for
- A61F2002/30581—Special structural features of bone or joint prostheses not otherwise provided for having a pocket filled with fluid, e.g. liquid
- A61F2002/30583—Special structural features of bone or joint prostheses not otherwise provided for having a pocket filled with fluid, e.g. liquid filled with hardenable fluid, e.g. curable in-situ
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
- A61F2210/00—Particular material properties of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof
- A61F2210/0004—Particular material properties of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof bioabsorbable
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
- A61F2210/00—Particular material properties of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof
- A61F2210/0085—Particular material properties of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof hardenable in situ, e.g. epoxy resins
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
- A61L2300/20—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices containing or releasing organic materials
- A61L2300/252—Polypeptides, proteins, e.g. glycoproteins, lipoproteins, cytokines
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
- A61L2300/40—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices characterised by a specific therapeutic activity or mode of action
- A61L2300/404—Biocides, antimicrobial agents, antiseptic agents
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- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61L—METHODS OR APPARATUS FOR STERILISING MATERIALS OR OBJECTS IN GENERAL; DISINFECTION, STERILISATION OR DEODORISATION OF AIR; CHEMICAL ASPECTS OF BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES; MATERIALS FOR BANDAGES, DRESSINGS, ABSORBENT PADS OR SURGICAL ARTICLES
- A61L2300/00—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices
- A61L2300/40—Biologically active materials used in bandages, wound dressings, absorbent pads or medical devices characterised by a specific therapeutic activity or mode of action
- A61L2300/404—Biocides, antimicrobial agents, antiseptic agents
- A61L2300/406—Antibiotics
Definitions
- the present invention relates to a method and composition for producing biodegradable polymers, and more particularly to the use of such polymers for providing syringeable, in-situ forming, solid, biodegradable implants.
- Biodegradable polymers have been used for many years in medical applications. These include sutures, surgical clips, staples, implants, and drug delivery systems. The majority of these biodegradable polymers have been thermoplastic materials based upon glycolide, lactide, ⁇ -caprolactone, and copolymers thereof. Typical examples are the polyglycolide sutures described in U.S.-Pat. No. 3,297,033 to Schmitt, the poly(L-lactide-co-glycolide) sutures described in U.S. Pat. No. 3,636,956 to Schneider, the poly(L-lactide-co-glycolide) surgical clips and staples described in U.S. Pat. No.
- biodegradable polymers described in these patents are thermoplastic materials. Consequently, they can be heated and formed into various shapes such as fibers, clips, staples, pins, films, etc. Only when heated above their melting point do these polymers become liquid. During their normal use, they are solids.
- Thermoset biodegradable polymers have also been previously described for use in medical applications. These polymers have been formed by crosslinking reactions which lead to high-molecular-weight materials that do not melt or form flowable liquids at high temperatures. Typical examples of these materials are the crosslinked polyurethanes described in U.S. Pat. No. 2,933,477 to Hostettler and U.S. Pat. No. 3,186,971 to Hostettler et al. Copolymers based on ⁇ -caprolactone and L-lactide or DL-lactide crosslinked via peroxide initiators were described in U.S. Pat. Nos. 4,045,418 and 4,057,537, both to Sinclair.
- Crosslinked caprolactone copolymers have been prepared by incorporation of a bislactone into a monomer feed, as described in U.S. Pat. No. 4,379,138 to Pitt et al. Trihydroxy-functional copolymers of ⁇ -caprolactone and ⁇ -valerolactone have been crosslinked with diisocyanates, thereby affording biodegradable polymers, as described in Pitt et al., J. Polym. Sci.: Part A: Polym Chem. 25:955-966; 1987. These polymers are also solids when crosslinked or cured.
- biodegradable polymers Although these two classes of biodegradable polymers have many useful biomedical applications, there are several important limitations to their use in the body where body is defined as that of humans, animals, birds, fish, and reptiles. Because these polymers are solids, all instances involving their use have required initially forming the polymeric structures outside the body, followed by insertion of the solid structure into the body. For example, sutures, clips, and staples are all formed from thermoplastic biodegradable polymers prior to use. When inserted into the body, they retain their original shape rather than flow to fill voids or cavities where they may be most needed.
- drug-delivery systems using these biodegradable polymers have to be formed outside the body.
- the drug is incorporated into the polymer and the mixture shaped into a certain form such a cylinder, disc, or fiber for implantation.
- the drug-delivery system has to be inserted into the body through an incision. These incisions are often larger than desired by the medical profession and lead to a reluctance of the patients to accept such an implant or drug-delivery system.
- microspheres or microcapsules prepared from these polymers and containing drugs for release into the body are sometimes difficult to produce on a large scale, and their storage and injection characteristics present problems.
- one other major limitation of the microcapsule or small-particle system is their lack of reversibility without extensive surgical intervention. That is, if there are complications after they have been injected, it is considerably more difficult to remove them from the body than with solid implants.
- the present invention relates to the production and use of biodegradable polymers as prosthetic implants and controlled-release, drug-delivery systems which can be administered as liquids via, for example, a syringe and needle, but which coagulate or cure (“set”) shortly after dosing to form a solid.
- the implants are biodegradable because they are made from biodegradable polymers and copolymers comprising two types of polymer systems: thermoplastic and thermosetting.
- thermoplastic system in which a solid, linear-chain, biodegradable polymer or copolymer is dissolved in a solvent, which is nontoxic and water miscible, to form a liquid solution.
- a solvent which is nontoxic and water miscible
- the solvent dissipates or diffuses away from the polymer, leaving the polymer to coagulate or solidify into a solid structure.
- the placement of the solution can be anywhere within the body, including soft tissue such as muscle or fat, hard tissue such as bone, or a cavity such as the periodontal, oral, vaginal, rectal, nasal, or a pocket such as a periodontal pocket or the cul-de-sac of the eye.
- the biologically active agent is added to the polymer solution where it is either dissolved to form a homogeneous solution or dispersed to form a suspension or a dispersion of drug within the polymeric solution.
- the solvent diffuses away from the polymer-drug mixture and water diffuses into the mixture where it coagulates the polymer thereby trapping or encapsulating the drug within the polymeric matrix as the implant solidifies.
- the release of the drug then follows the general rules for diffusion or dissolution of a drug from within a polymeric matrix.
- thermosetting system comprising the synthesis of crosslinkable polymers which are biodegradable and which can be formed and cured in-situ.
- the thermosetting system comprises reactive, liquid, oligomeric polymers which contain no solvents and which cure in place to form solids, usually with the addition of a curing catalyst.
- the multifunctional polymers useful in the thermosetting system are first synthesized via copolymerization of either DL-lactide of L-lactide with ⁇ -caprolactone using a multifunctional polyol initiator and a catalyst to form polyol-terminated prepolymers.
- the polyol-terminated prepolymers are then converted to acrylic ester-terminated prepolymers, preferably by acylation of the alcohol terminus with acryloyl chloride via a Schotten-Baumann-like technique, i.e., reaction of acyl halides with alcohols.
- the acrylic ester-terminated prepolymers may also be synthesized in a number of other ways, including but not limited to, reaction of carboxylic acids (i.e., acrylic or methacrylic acid) with alcohols, reaction of carboxylic acid esters (i.e., methyl acrylate or methyl methacrylate) with alcohols by transesterification, and reaction of isocyanatoalkyl acrylates (i.e., isocyanatoethyl methacrylate) with alcohols.
- carboxylic acids i.e., acrylic or methacrylic acid
- carboxylic acid esters i.e., methyl acrylate or methyl methacrylate
- isocyanatoalkyl acrylates i.e., isocyanatoethyl methacrylate
- the liquid acrylic-terminated prepolymer is cured, preferably by the addition of benzoyl peroxide or azobisisobutyronitrile, to a more solid structure.
- the catalyst is added to the liquid acrylic-terminated prepolymer immediately prior to injection into the body. Once inside the body, the crosslinking reaction will proceed until sufficient molecular weight has been obtained to cause the polymer to solidify.
- the liquid prepolymer when injected, will flow into the cavity or space in which it is placed and assume that shape when it solidifies.
- biologically active agents are added to the liquid polymer systems in the uncatalyzed state.
- the polymer may be injected via syringe and needle into a body while it is in liquid form and then left in-situ to form a solid biodegradable implant structure.
- the need to form an incision is eliminated, and the implant will assume the shape of its cavity.
- a drug-delivery vehicle may be provided by adding a biologically active agent to the liquid prior to injection. Once the implant is formed, it will release the agent to the body and then biodegrade.
- biologically active agent means a drug or some other substance capable of producing an effect on a body.
- FIG. 1 illustrates the synthesis of acrylate-terminated prepolymers and subsequent crosslinking by free-radical initiators
- FIG. 2 illustrates structures for the random copolymer of ⁇ -caprolactone and L-lactide initiated with a diol
- Table 1 is a summary of the bifunctional PLC prepolymers synthesized
- Table 2 is a summary of the acrylic ester terminated prepolymers synthesized.
- Table 3 is a summary of curing studies.
- the present invention relates to biodegradable, in-situ forming implants and methods for producing the same.
- the present invention also relates to a liquid biodegradable polymeric delivery system that can be injected into a body where it forms a solid and releases a biologically active agent at a controlled rate.
- Two types of biodegradable polymeric systems are described: thermoplastic copolymers dissolved in a biocompatible solvent and thermosetting polymers that are liquids without the use of solvents.
- thermoplastic system in which a solid, linear-chain, biodegradable polymer is dissolved in a biocompatible solvent to form a liquid, which can then be administered via a syringe and needle.
- biodegradable polymers which can be used in this application are polylactides, polyglycolides, polycaprolactones, polyanhydrides, polyamides, polyurethanes, polyesteramides, polyorthoesters, polydioxanones, polyacetals, polyketals, polycarbonates, polyorthocarbonates, polyphosphazenes, polyhydroxybutyrates, polyhydroxyvalerates, polyalkylene oxalates, polyalkylene succinates, poly(malic acid), poly (amino acids), polyvinylpyrrolidone, polyethylene glycol, polyhydroxycellulose, chitin, chitosan, and copolymers, terpolymers, or combinations or mixtures of the above materials.
- the preferred polymers are those which have a lower degree of crystallization and are more hydrophobic. These polymers and copolymers are more soluble in the biocompatible solvents than the highly crystalline polymers such as polyglycolide and chitin which also have a high degree of hydrogen-bonding.
- Preferred materials with the desired solubility parameters are the polylactides, polycaprolactones, and copolymers of these with glycolide in which there are more amorphous regions to enhance solubility.
- the solvent for the biodegradable polymer be non-toxic, water miscible, and otherwise biocompatible. Solvents that are toxic should not be used to inject any material into a living body. The solvents must also be biocompatible so that they do not cause severe tissue irritation or necrosis at the site of implantation. Furthermore, the solvent should be water miscible so that it will diffuse quickly into the body fluids and allow water to permeate into the polymer solution and cause it to coagulate or solidify.
- solvents examples include N-methyl-2-pyrrolidone, 2-pyrrolidone, ethanol, propylene glycol, acetone, methyl acetate, ethyl acetate, methyl ethyl ketone, dimethylformamide, dimethyl sulfoxide, tetrahydrofuran, caprolactam, decylmethylsulfoxide, oleic acid, and 1-dodecylazacycloheptan-2-one.
- the preferred solvents are N-methyl-2-pyrrolidone, 2-pyrrolidone, dimethyl sulfoxide, and acetone because of their solrating solvating ability and their compatibility.
- the solubility of the biodegradable polymers in the various solvents will differ depending upon their crystallinity, their hydrophilicity, hydrogen-bonding, and molecular weight. Thus, not all of the biodegradable polymers will be soluble in the same solvent, but each polymer or copolymer should have its optimum solvent. Lower molecular-weight polymers will normally dissolve more readily in the solvents than high-molecular-weight polymers. As a result, the concentration of a polymer dissolved in the various solvents will differ depending upon type of polymer and its molecular weight. Conversely, the higher molecular-weight polymers will normally tend to coagulate or solidify faster than the very low-molecular-weight polymers. Moreover the higher molecular-weight polymers will tend to give higher solution viscosities than the low-molecular-weight materials. Thus for optimum injection efficiency, the molecular weight and the concentration of the polymer in the solvent have to be controlled.
- low-molecular-weight polylactic acid formed by the condensation of lactic acid will dissolve in N-methyl-2-pyrrolidone(NMP) to give a 73% by weight solution which still flows easily through a 23-gauge syringe needle
- NMP N-methyl-2-pyrrolidone
- DL-PLA higher molecular-weight poly(DL-lactide)
- the higher molecular-weight polymer solution coagulates immediately when placed into water.
- the low-molecular-weight polymer solution although more concentrated, tends to coagulate very slowly when placed into water.
- a solvent mixture can be used to increase the coagulation rate.
- one liquid component of the mixture is a good solvent for the polymer, and the other component is a poorer solvent or a non-solvent.
- the two liquids are mixed at a ratio such that the polymer is still soluble but precipitates with the slightest increase in the amount of non-solvent, such as water in a physiological environment.
- the solvent system must be miscible with both the polymer and water.
- An example of such a binary solvent system is the use of NMP and ethanol for low-molecular-weight DL-PLA. The addition of ethanol to the NMP/polymer solution increases its coagulation rate significantly.
- the polymer solution is placed in a syringe and injected through a needle into the body. Once in place, the solvent dissipates, the remaining polymer solidifies, and a solid structure is formed.
- the implant will adhere to its surrounding tissue or bone by mechanical forces and can assume the shape of its surrounding cavity.
- the biodegradable polymer solution can be injected subdermally like collagen to build up tissue or to fill in defects. It can also be injected into wounds including burn wounds to prevent the formation Of deep scars. Unlike collagen, the degradation time of the implant can be varied from a few weeks to years depending upon the polymer selected and its molecular weight.
- the injectable polymer solution can also be used to mend bone defects or to provide a continuous matrix when other solid biodegradable implants such as hydroxyapatite plugs are inserted into bone gaps.
- the injectable system can also be used to adhere tissue to tissue or other implants to tissue by virtue of its mechanical bonding or encapsulation of tissue and prosthetic devices.
- thermoplastic system Another envisioned use of the thermoplastic system is to provide a drug-delivery system.
- a bioactive agent is added to the polymer solution prior to injection, and then the polymer/solvent/agent mixture is injected into the body.
- the drug will also be soluble in the solvent, and a homogeneous solution of polymer and drug will be available for injection.
- the drug will not be soluble in the solvent, and a suspension or dispersion of the drug in the polymer solution will result. This suspension or dispersion can also be injected into the body. In either case, the solvent will dissipate and the polymer will solidify and entrap or encase the drug within the solid matrix.
- the release of drug from these solid implants will follow the same general rules for release of a drug from a monolithic polymeric device.
- the release of drug can be affected by the size and shape of the implant, the loading of drug within the implant, the permeability factors involving the drug and the particular polymer, and the degradation of the polymer.
- the above parameters can be adjusted by one skilled in the art of drug delivery to give the desired rate and duration of release.
- drug or bioactive (biologically active) agent includes without limitation physiologically or pharmacologically active substances that act locally or systemically in the body.
- Representative drugs and biologically active agents to be used with the syringeable, in-situ forming solid implant systems include, without limitation, peptide drugs, protein drugs, desensitizing agents, antigens, vaccines, anti-infectives, antibiotics, antimicrobials, antiallergenics, steroidal anti-inflammatory agents, antiallergenics, steroidal anti-inflammatory agents, decongestants, miotics, anticholinergios, sympathomimetics, sedatives, hypnotics, psychic energizers, tranquilizers, androgenic steroids, estrogens, progestational agents, humoral agents, prostaglandins, analgesics, antispasmodics, antimalarials, antihistamines, cardioactive agents, non-steroidal anti-inflammatory agents, antiparkinsonian agents, antihypertensive agents
- drugs or biologically active agents that can be released in an aqueous environment can be utilized in the described injectable delivery system.
- various forms of the drugs or biologically active agents may be used. These include without limitation forms such as uncharged molecules, molecular complexes, salts, ethers, esters, amides, etc., which are biologically activated when injected into the body.
- the amount of drug or biologically active agent incorporated into the injectable, in-situ, solid forming implant depends upon the desired release profile, the concentration of drug required for a biological effect, and the length of time that the drug has to be released for treatment. There is no critical upper limit on the amount of drug incorporated into the polymer solution except for that of an acceptable solution or dispersion viscosity for injection through a syringe needle. The lower limit of drug incorporated into the delivery system is dependent simply upon the activity of the drug and the length of time needed for treatment.
- the solid implant formed within the injectable polymer solution will slowly biodegrade within the body and allow natural tissue to grow and replace the impact as it disappears.
- the material when the material is injected into a soft-tissue defect, it will fill that defect and provide a scaffold for natural collagen tissue to grow.
- This collagen tissue will gradually replace the biodegradable polymer.
- the biodegradable polymer With hard tissue such as bone, the biodegradable polymer will support the growth of new bone cells which will also gradually replace the degrading polymer.
- the solid implant formed from the injectable system will release the drug contained within its matrix at a controlled rate until the drug is depleted.
- the polymer will degrade after the drug has been completely released.
- other drugs such as peptides or proteins, the drug will be completely released only after the polymer has degraded to a point where the non-diffusing drug has been exposed to the body fluids.
- the injectable, in-situ forming biodegradable implants can also be produced by crosslinking appropriately functionalized biodegradable polymers.
- the thermosetting system comprises reactive, liquid, oligomeric polymers which cure in place to form solids, usually with the addition of a curing catalyst.
- any of the biodegradable polymers previously described for the thermoplastic system can be used, the limiting criteria is that low-molecular-weight oligomers of these polymers or copolymers must be liquids and they must have functional groups on the ends of the prepolymer which can be reacted with acryloyl chloride to produce acrylic ester capped prepolymers.
- the preferred biodegradable system is that produced from poly(DL-lactide-co-caprolactone), or “DL-PLC”.
- Low-molecular-weight polymers or oligomers produced from these materials are flowable liquids at room temperature.
- Hydroxy-terminated PLC prepolymers may be synthesized via copolymerization of DL-lactide or L-lactide and ⁇ -caprolactone with a multifunctional polyol initiator and a catalyst.
- Catalysts useful for the preparation of these prepolymers are preferably basic or neutral ester-interchange (transesterification) catalysts.
- Metallic esters of carboxylic acids containing up to 18 carbon atoms such as formic, acetic, lauric, stearic, and benzoic are normally used as such catalysts.
- Stannous octoate and stannous chloride are the preferred catalysts, both for reasons of FDA compliance and performance.
- a bifunctional chain initiator such as ethylene glycol is employed.
- a trifunctional initiator such as trimethylolpropane produces a trifunctional polymer, etc.
- the amount of chain initiator used determines the resultant molecular weight of the polymer or copolymer.
- the concentration of bifunctional initiator is very low, each initiator molecule can initiate two polymer chains. In any case, the polymer chains are terminated by hydroxyl groups, as seen in FIG. 1 .
- the assumption has been made that only one polymer chain is initiated per bifunctional initiator molecule. This assumption allows the calculation of a theoretical molecular weight for the prepolymers.
- the diol prepolymers were converted to acrylic-ester-capped prepolymers via a reaction with acryloyl chloride under Schotten-Baumann-like conditions, as seen in FIG. 2 and summarized in Table 2.
- Other methods of converting the diol prepolymers to acrylic-ester-capped prepolymers may also be employed.
- the acrylic prepolymers and diol prepolymers were then cured, as summarized in Table 3.
- the general procedure for the curing of the prepolymers is now described: to 5.0 g of acrylic prepolymer contained in a small beaker was added a solution of benzoyl peroxide (BP) in approximately 1 mL of CH 2 Cl 2 . In some cases, fillers or additional acrylic monomers were added to the prepolymers prior to the introduction of the BP solution. The mixtures were stirred thoroughly and then poured into small petri dishes. The dishes were placed in a preheated vacuum oven for curing. Some of the samples were cured in air and not in vacuo, and these samples are so indicated in Table 3.
- BP benzoyl peroxide
- thermosetting system may be used wherever a biodegradable implant is desired.
- the liquid prepolymer/curing agent mixture may be placed into a sytringe and injected into a body.
- the mxiture than solidifies in-situ, thereby providing an implant without an incision.
- a drug-delivery system may be provided by adding a biologically active agent to the prepolymer prior to injection. Once in-situ, the system will cure to a solid; eventually, it will biodegrade, and the agent will be gradually released.
- Poly(DL-lactic acid) was prepared by the simple poly-condensation of lactic acid. No catalysts were used, and the reaction times were varied to produce polymers with different theoretical molecular weights. These polymers were designated as DL-PLA oligomers. A quantity of the solid oligomer was dissolved in NMP to give a 68:32 ratio of polymer to solvent. Sanguinarine chloride(SaCl), a benzophenanthridine alkaloid with antimicrobial activity especially toward periodontal pathogens, was added to the polymer solution to give a 2% by weight dispersion of the drug in the total mixture.
- SaCl Sanguinarine chloride
- the dispersion of drug and polymer solution was then injected into a dialysis tube (diameter of 1.5 mm) with a sterile disposable syringe without a needle.
- a dialysis tube dialysis tube (diameter of 1.5 mm) with a sterile disposable syringe without a needle.
- Each end of the 6-in. length of dialysis tubing was tied with a knot to prevent loss of the drug/polymer mass, and the tube with the injected material was placed in a pH 7 Sorenson's buffer receiving fluid maintained at 37° C.
- the drug/polymer mass coagulated into a solid mass, and the drug began to be released from the polymer as indicated by an orange-red color in the receiving fluid.
- the quantity of solution injected into the dialysis tube was about 250 ⁇ L or about 100 mg of solids.
- the dialysis tubing was selected to have a molecular-weight cutoff of about 3,500. With this molecular-weight cutoff, the SaCl released from the polymer could easily diffuse through the walls of the tubing, but any solid polymer would be retained.
- the dialysis tubing containing the drug/polymer matrix was removed frequently and placed in a bottle of fresh receiving fluid. The old receiving fluid containing the released drug was then acidified to a pH of 2.76 to convert all released drug to the iminium ion form of the drug, and the concentration of drug was determined by measuring the ultraviolet absorption (UV) at a wavelength of 237 nm. The cumulative mass of drug released and the cumulative fraction were then calculated and plotted as a function of time. Approximately 60% of the drug was released in the first day, 72% after 2 days, 85% after 5 days, 90% after 9 days, and 97% after 14 days.
- UV ultraviolet absorption
- SaEt dissolved in the polymer solution to give a homogeneous solution of drug and polymer.
- the release of SaEt was slower than that for SaCl as expected because of its lower water solubility. After the first day, approximately 45% was released, 52% after 2 days, 60% after 5 days, 70% after 9 days, and 80% after 14 days.
- Poly(DL-lactide) with an inherent viscosity of 0.08 dL/g and a theoretical molecular weight of 2,000 was prepared by the ring-opening polymerization of DL-lactide using lauryl alcohol as the initiator and stannous chloride as the catalyst. This polymer was then dissolved in NMP to give a 40% by weight polymer solution. SaCl was dispersed in the solution of this polymer in NMP to give a 1.5% by weight dispersion of the drug in the solution and the release rate determined as described in Example 1. The release rate of the drug from this higher molecular-weight polymer was slower than from the DL-PLA oligomer. After the first day, approximately 32% was released, 40% after 2 days, 45% after 5 days, and 50% after 15 days.
- SaEt was added to the same polymer solution of DL-PLA in NMP as described in Example 3. A homogeneous solution with the drug at 1.5% by weight was obtained. The release of drug from this solution determined using the same procedure described in Example 1 gave a much slower release of SaEt than from the DL-PLA oligomer. After the first day approximately 8% was released, 14% after 2 days, 20% after 5 days, 23% after 9 days, and 28% after 14 days.
- the 2% -loaded formulation gave 65% release after 1 day, 75% after 2 days, and 88% after 5 days; the 7%-loaded formulation gave 48% release after 1 day, 52% after 2 days, and 58% after 5 days; and the 14%-loaded formulation gave 38% release after 1 day, 43% after 2 days, and 49% after 5days.
- Poly(DL-lactide-co-glycolide) was prepared by the ring-opening polymerization of a mixture of DL-lactide and glycolide using lauryl alcohol as the initiator and stannous chloride as the catalyst. The proportions of the two monomers were adjusted so that the final copolymer(DL-PLG) had a 50:50 ratio of the two monomers as determined by nuclear magnetic resonance spectrophotometry. The initiator was also adjusted to give a copolymer with a theoretical molecular weight of 1500 daltons. The copolymer was dissolved in NMP to give a 70% by weight polymer solution. SaCl was added to this solution to give a 2% by weight dispersion of the drug in the polymer solution.
- the release of drug from this formulation was determined using the same procedure described in Example 1. A much lower release rate was obtained from the copolymer than from the DL-PLA oligomer of DL-PLA 2000 molecular weight materials. After 2 days approximately 7% of the drug was released, 10% after 5 days, 12% after 7 days, and 16% after 14 days.
- SaEt was added to the same solution of DL-PLG in NMP as described in Example 6 to give a 2% by weight solution of the drug.
- the release of drug from this formulation was determined by the same procedure as described previously.
- the release rate of SaEt from this formulation was identical to that for SaCl described in Example 6.
- Tetracycline as the free base was added to the same solution of DL-PLG in NMP as described in Example 6.
- the drug dissolved completely in the polymer solution to give a 2.4% by weight solution to the drug.
- the release of the drug from this formulation was determined by a similar procedure to that described in Example 1 except the receiving fluid was not acidified to a pH of 2.76 and the concentration of TCB was determined by UV absorption at the wavelength appropriate for the drug.
- the release of TCB from this formulation was more linear and at a much higher rate than that for SaCl or SaEt from the same copolymer. After 1 day approximately 44% of the drug was released, 54% after 2 days, 68% after 5 days, 73% after 6 days, 80% after 7 days, 87% after 9 days, 96% after 12 days, and 100% after 14 days.
- Tetracycline as the hydrochloride salt (TCH) was added to the same solution of DL-PLG in NMP as described in Example 6.
- the salt form of the drug also dissolved completely in the polymer solution.
- the release of drug from this formulation was determined as described in Example 8 and found to be similar to that for the free base except for a slightly lower rate. After 1 day approximately 32% of the drug was released, 40% after 2 days, 57% after 5 days, 64% after 6 days, 75% after 7 days, 82% after 9 days, 92% after 12 days, and 100% after 14 days.
- DL-PLA with an inherent viscosity of 0.26 dL/g and a theoretical molecular weight of approximately 10,000 daltons was prepared by the ring-opening polymerization of DL-lactide using lauryl alcohol as the initiator and stannous chloride at the catalyst.
- the polymer was dissolved in NMP to give a 50% by weight polymer solution.
- a quantity of the polymer solution (100 ⁇ L) was injected subdermally into rabbits, and the tissue reaction was compared to that of a USP negative plastic.
- the test sites were evaluated for signs of local irritation, in accordance with the Draize method, immediately after injection, at 1 and 6 hours post injection, and once daily thereafter until scheduled sacrifice at 7, 14 or 21 days.
- the reaction at the test sites was equivalent to that at the control USP negative plastic.
- the polymer solution (100 ⁇ L) was also administered subgingivally into sites created by dental extractions in Beagle dogs. Control sites were flushed with saline solution. The dogs were examined daily for signs of mortality, pharmacotoxic effects, body weights, and local gingival irritation. The animals were sacrificed at 15 and 21 days. No distinct differences were noted between the control and test sites.
- DL-PLA with an inherent viscosity of 0126 dL/g and a molecular weight of about 10,000 was dissolved in NMP to give a 50% by weight polymer solution.
- SaCl was added to the polymer solution to give a 2.4% by weight dispersion.
- This material was loaded into a 1-cc disposable syringe fitted with a 23-gauge blunted-end syringe needle, and the material was inserted into the periodontal pocket of a greyhound dog. The material flowed easily out of the narrow syringe tip. The polymer precipitated or coagulated into a film or solid mass when it contacted the saliva and fluid within the pocket.
- the dog was observed over a time of 2 weeks during which the mass of material remained within the pocket, adhering to tissue surrounding the pocket, and slowly changing color from a light orange to a pale white.
- the crevicular fluid from the pocket containing the implant was sampled during this 2-week period using Periostrips which are small strips of paper that are placed at the entrance to the periodontal pocket to wick up small quantities of the crevicular fluid within the pocket.
- the volume of fluid collected is determined using a Periotron which measures the changes in conductance of the paper strip. The Periotron is calibrated before use with a known volume of serum.
- the paper strip containing the collected fluid is then extracted with a solution of 0.5% by volume of hydrochloric acid in methanol and injected into a liquid chromatograph where the quantity of drug is determined by reference to a known concentration of the same compound.
- the quantity of SaCl extracted from the paper strip is divided by the quantity of crevicular fluid collected to calculate the concentration of drug in the fluid.
- the SaCl concentration in the crevicular fluid was 63.2 ⁇ g/mL after 3 days, 80.2 ⁇ g/mL after 7 days, 67.8 ⁇ g/mL after 10 days, and 70.5 ⁇ g/mL after 14 days.
- the addition funnel was charged with 15.4 g of acryloyl chloride (0.95 equiv/equiv OH) in 15 mL of THF, and the solution was added dropwise to the stirred reaction mixture over 1 hour. The mixture was stirred overnight and allowed to reach room temperature. The precipitated triethylamine hydrochloride was removed by filtration, and the filtrate was evaporated in vacuo, affording a pale yellow oil, which was the acrylate-terminated prepolymer.
- the acylations employing CH 2 Cl 2 as solvent were conducted in a similar manner. However, the reaction times at 0° C. were shortened to 1 hour, whereupon the reaction mixtures were allowed to reach room temperature over 1 hour.
- C964-139-1 C964-136-1 0.81 0 2 THF Gelled. Overnight exposure to residual Et 3 N.HCl in refrigerator.
- C964-144-1 C964-136-1 0.81 0 1 CH 2 Cl 2 No problems, stable.
- C964-146-1 C964-124-1 0.33 0 1 CH 2 Cl 2 No problems, stable.
- C964-121-3 C964-118-1 0.5 none 82 16 78 Slightly more elastic than C964-121-2.
- C964-121-4 C964-118-1 0.1 none 82 16 69 Same as C964-121-3.
- C964-122-1 C964-118-1 1.0 TMPTETA b 46 82 2.5 94 Less rubbery than C964-120 and C964-121, brittle.
- C964-122-2 C964-118-1 0.5 TMPTETA 46 82 2.5 91 Same as C964-122-1, more flexible.
- C964-122-3 C964-118-1 1.0 TMPTETA 175 82 2.5 95 Not rubbert at all, brittle, weak.
- C964-122-4 C964-118-1 0.5 TMPTETA 175 82 2.5 93 Similar to C964-122-3.
- C964-123-1 C964-118-1 0.1 TMPTETA 46 82 2.5 89 Rubbery, stronger than C964-120 and C964-121, not flexible.
- C964-123-2 C964-118-1 0.25 TMPTETA 46 82 2.5 83 About the same as C964-123-1, may be more brittle.
- C964-123-3 C964-132-1 0.1 TMPTETA 175 82 2.5 92 Not rubbery, strong, brittle.
- C964-134-1 C964-132-1 0.05 (AIBN) c none 60 d 17 Liquid No cure.
- C964-135-6 C964-128-1 e 0.05 none 80 d 17 Liquid No cure.
- C964-135-7 C964-128-1 e 0.10 none 80 d 17 Liquid No cure.
- C964-148-3 C964-144-1 0.25 none 80 2 66 C964-148-4 and C964-148-6 were about the same in toughness, and both were better than C964-148-3 and C964-148-5.
- C964-148-4 C964-144-1 0.50 none 80 2 68 C964-148-4 and C964-148-6 were about the same in toughness, and both were better than C964-148-3 and C964-148-5.
- C964-148-5 C964-144-1 1.00 none 80 2 67 C964-148-4 and C964-148-4 were about the same in toughness, and both were better than C964-148-3 and C964-148-5.
- C964-150-4 C964-146-1 0.50 none 80 2 56 Elastic, flexible, moderately strong.
- C964-150-5 C964-146-1 1.00 none 80 2 50 Elastic, flexible, moderately strong.
- b TMPTETA trimethylolpropane triethoxy triacrylate.
- c AIBN azobisisobutyronitrile.
- d Cured in air at atmospheric pressure. e Diol prepolymer used.
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Abstract
A biodegradable polymer is provided for use in providing syringeable, in-situ forming, solid biodegradable implants for animals. The polymer is placed into the animal in liquid form and cures to form the implant in-situ. A thermoplastic system to form said implant comprises the steps of dissolving a non-reactive polymer in biocompatible solvent to form a liquid, placing the liquid within the animal, and allowing the solvent to dissipate to produce the implant. An alternative, thermosetting system comprises mixing together effective amounts of a liquid acrylic ester terminated, biodegradable prepolymer and a curing agent, placing the liquid mixture within an animal and allowing the prepolymer to cure to form the implant. Both systems provide a syringeable, solid biodegradable delivery system by the addition of an effective level of biologically active agent to the liquid before injection into the body.
Description
This application is a continuation of U.S. patent application Ser. No. 08/210,891, filed Mar. 18, 1994, now U.S. Pat. No. 5,739,176 (issued Apr. 14, 1998), which is a continuation of U.S. patent application Ser. No. 07/788,032, filed Dec. 23, 1991, now U.S. Pat. No. 5,340,849 (issued Aug. 23, 1994), which is a division of U.S. patent application Ser. No. 07/513,782, filed Apr. 24, 1990, now U.S. Pat. No. 5,278,201 (issued Jan. 11, 1994), which is a division of U.S. patent application Ser. No. 07/252,645, filed Oct. 3, 1988, now U.S. Pat. No. B1-4,938,763 (issued Jul. 3, 1990; certificate issued Jul. 4, 1995).
The present invention relates to a method and composition for producing biodegradable polymers, and more particularly to the use of such polymers for providing syringeable, in-situ forming, solid, biodegradable implants.
Biodegradable polymers have been used for many years in medical applications. These include sutures, surgical clips, staples, implants, and drug delivery systems. The majority of these biodegradable polymers have been thermoplastic materials based upon glycolide, lactide, ε-caprolactone, and copolymers thereof. Typical examples are the polyglycolide sutures described in U.S.-Pat. No. 3,297,033 to Schmitt, the poly(L-lactide-co-glycolide) sutures described in U.S. Pat. No. 3,636,956 to Schneider, the poly(L-lactide-co-glycolide) surgical clips and staples described in U.S. Pat. No. 4,523,591 to Kaplan et al., and the drug-delivery systems described in U.S. Pat. No. 3,773,919 to Boswell et al., U.S. Pat. No. 3,887,699 to Yolles, U.S. Pat. No. 4,155,992 to Schmitt, U.S. Pat. No. 4,379,138 to Pitt et al., and U.S. Pat. Nos. 4,130,639 and 4,186,189 to Shalaby et al.
All of the biodegradable polymers described in these patents are thermoplastic materials. Consequently, they can be heated and formed into various shapes such as fibers, clips, staples, pins, films, etc. Only when heated above their melting point do these polymers become liquid. During their normal use, they are solids.
Thermoset biodegradable polymers have also been previously described for use in medical applications. These polymers have been formed by crosslinking reactions which lead to high-molecular-weight materials that do not melt or form flowable liquids at high temperatures. Typical examples of these materials are the crosslinked polyurethanes described in U.S. Pat. No. 2,933,477 to Hostettler and U.S. Pat. No. 3,186,971 to Hostettler et al. Copolymers based on ε-caprolactone and L-lactide or DL-lactide crosslinked via peroxide initiators were described in U.S. Pat. Nos. 4,045,418 and 4,057,537, both to Sinclair. Crosslinked caprolactone copolymers have been prepared by incorporation of a bislactone into a monomer feed, as described in U.S. Pat. No. 4,379,138 to Pitt et al. Trihydroxy-functional copolymers of ε-caprolactone and ε-valerolactone have been crosslinked with diisocyanates, thereby affording biodegradable polymers, as described in Pitt et al., J. Polym. Sci.: Part A: Polym Chem. 25:955-966; 1987. These polymers are also solids when crosslinked or cured.
Although these two classes of biodegradable polymers have many useful biomedical applications, there are several important limitations to their use in the body where body is defined as that of humans, animals, birds, fish, and reptiles. Because these polymers are solids, all instances involving their use have required initially forming the polymeric structures outside the body, followed by insertion of the solid structure into the body. For example, sutures, clips, and staples are all formed from thermoplastic biodegradable polymers prior to use. When inserted into the body, they retain their original shape rather than flow to fill voids or cavities where they may be most needed.
Similarly, drug-delivery systems using these biodegradable polymers have to be formed outside the body. In such instances, the drug is incorporated into the polymer and the mixture shaped into a certain form such a cylinder, disc, or fiber for implantation. With such solid implants, the drug-delivery system has to be inserted into the body through an incision. These incisions are often larger than desired by the medical profession and lead to a reluctance of the patients to accept such an implant or drug-delivery system.
The only way to avoid the incision with these polymers is to inject them as small particles, microspheres, or microcapsules. These may or may not contain a drug which can be released into the body. Although these small particles can be injected into the body with a syringe, they do not always satisfy the demand for a biodegradable implant. Because they are particles, they do not form a continuous film or solid implant with the structural integrity needed for certain prostheses. When inserted into certain body cavities such as the mouth, a periodontal pocket, the eye, or the vagina where there is considerable fluid flow, these small particles, microspheres, or microcapsules are poorly retained because of their small size and discontinuous nature. In addition, microspheres or microcapsules prepared from these polymers and containing drugs for release into the body are sometimes difficult to produce on a large scale, and their storage and injection characteristics present problems. Furthermore, one other major limitation of the microcapsule or small-particle system is their lack of reversibility without extensive surgical intervention. That is, if there are complications after they have been injected, it is considerably more difficult to remove them from the body than with solid implants.
Therefore, there exists a need for a method and composition which provides a biodegradable, polymeric structure useful in overcoming the above-described limitations.
There exists a further need for a method and composition for providing syringeable, in-situ forming, solid, biodegradable implants which can be used as prosthetic devices and/or controlled delivery systems.
Moreover, there exists a need for such a method and composition which can provide implants having a range of properties from soft to rigid, so as to be usable with both soft and hard tissue.
The present invention relates to the production and use of biodegradable polymers as prosthetic implants and controlled-release, drug-delivery systems which can be administered as liquids via, for example, a syringe and needle, but which coagulate or cure (“set”) shortly after dosing to form a solid. The implants are biodegradable because they are made from biodegradable polymers and copolymers comprising two types of polymer systems: thermoplastic and thermosetting.
A thermoplastic system is provided in which a solid, linear-chain, biodegradable polymer or copolymer is dissolved in a solvent, which is nontoxic and water miscible, to form a liquid solution. Once the polymer solution is placed into the body where there is sufficient water, the solvent dissipates or diffuses away from the polymer, leaving the polymer to coagulate or solidify into a solid structure. The placement of the solution can be anywhere within the body, including soft tissue such as muscle or fat, hard tissue such as bone, or a cavity such as the periodontal, oral, vaginal, rectal, nasal, or a pocket such as a periodontal pocket or the cul-de-sac of the eye. For drug-delivery systems, the biologically active agent is added to the polymer solution where it is either dissolved to form a homogeneous solution or dispersed to form a suspension or a dispersion of drug within the polymeric solution. When the polymer solution is exposed to body fluids or water, the solvent diffuses away from the polymer-drug mixture and water diffuses into the mixture where it coagulates the polymer thereby trapping or encapsulating the drug within the polymeric matrix as the implant solidifies. The release of the drug then follows the general rules for diffusion or dissolution of a drug from within a polymeric matrix.
Another embodiment of the invention is also provided, namely, a thermosetting system comprising the synthesis of crosslinkable polymers which are biodegradable and which can be formed and cured in-situ. The thermosetting system comprises reactive, liquid, oligomeric polymers which contain no solvents and which cure in place to form solids, usually with the addition of a curing catalyst.
The multifunctional polymers useful in the thermosetting system are first synthesized via copolymerization of either DL-lactide of L-lactide with ε-caprolactone using a multifunctional polyol initiator and a catalyst to form polyol-terminated prepolymers. The polyol-terminated prepolymers are then converted to acrylic ester-terminated prepolymers, preferably by acylation of the alcohol terminus with acryloyl chloride via a Schotten-Baumann-like technique, i.e., reaction of acyl halides with alcohols. The acrylic ester-terminated prepolymers may also be synthesized in a number of other ways, including but not limited to, reaction of carboxylic acids (i.e., acrylic or methacrylic acid) with alcohols, reaction of carboxylic acid esters (i.e., methyl acrylate or methyl methacrylate) with alcohols by transesterification, and reaction of isocyanatoalkyl acrylates (i.e., isocyanatoethyl methacrylate) with alcohols.
The liquid acrylic-terminated prepolymer is cured, preferably by the addition of benzoyl peroxide or azobisisobutyronitrile, to a more solid structure. Thus, for an implant utilizing these crosslinkable polymers, the catalyst is added to the liquid acrylic-terminated prepolymer immediately prior to injection into the body. Once inside the body, the crosslinking reaction will proceed until sufficient molecular weight has been obtained to cause the polymer to solidify. The liquid prepolymer, when injected, will flow into the cavity or space in which it is placed and assume that shape when it solidifies. For drug delivery utilizing this system, biologically active agents are added to the liquid polymer systems in the uncatalyzed state.
In both the thermoplastic and the thermosetting systems, the advantages of liquid application are achieved. For example, the polymer may be injected via syringe and needle into a body while it is in liquid form and then left in-situ to form a solid biodegradable implant structure. The need to form an incision is eliminated, and the implant will assume the shape of its cavity. Furthermore, a drug-delivery vehicle may be provided by adding a biologically active agent to the liquid prior to injection. Once the implant is formed, it will release the agent to the body and then biodegrade. The term “biologically active agent” means a drug or some other substance capable of producing an effect on a body.
It is an object of the present invention, therefore, to provide a method and composition for producing biodegradable polymers.
It is also an object of the present invention to provide such a polymer which may be useful in producing syringeable, in-situ forming, solid biodegradable implants.
It is a further object of the present invention to provide such an implant which can be used in a controlled-release delivery system for biological agents.
It is a further object of the present invention to provide implants having a range of properties from soft and elastomeric to hard and rigid, so as to be usable with both soft and hard tissue.
FIG. 1 illustrates the synthesis of acrylate-terminated prepolymers and subsequent crosslinking by free-radical initiators;
FIG. 2 illustrates structures for the random copolymer of ε-caprolactone and L-lactide initiated with a diol;
Table 1 is a summary of the bifunctional PLC prepolymers synthesized;
Table 2 is a summary of the acrylic ester terminated prepolymers synthesized; and
Table 3 is a summary of curing studies.
The present invention relates to biodegradable, in-situ forming implants and methods for producing the same. The present invention also relates to a liquid biodegradable polymeric delivery system that can be injected into a body where it forms a solid and releases a biologically active agent at a controlled rate. Two types of biodegradable polymeric systems are described: thermoplastic copolymers dissolved in a biocompatible solvent and thermosetting polymers that are liquids without the use of solvents.
A. Thermoplastic System
A thermoplastic system is provided in which a solid, linear-chain, biodegradable polymer is dissolved in a biocompatible solvent to form a liquid, which can then be administered via a syringe and needle. Examples of biodegradable polymers which can be used in this application are polylactides, polyglycolides, polycaprolactones, polyanhydrides, polyamides, polyurethanes, polyesteramides, polyorthoesters, polydioxanones, polyacetals, polyketals, polycarbonates, polyorthocarbonates, polyphosphazenes, polyhydroxybutyrates, polyhydroxyvalerates, polyalkylene oxalates, polyalkylene succinates, poly(malic acid), poly (amino acids), polyvinylpyrrolidone, polyethylene glycol, polyhydroxycellulose, chitin, chitosan, and copolymers, terpolymers, or combinations or mixtures of the above materials. The preferred polymers are those which have a lower degree of crystallization and are more hydrophobic. These polymers and copolymers are more soluble in the biocompatible solvents than the highly crystalline polymers such as polyglycolide and chitin which also have a high degree of hydrogen-bonding. Preferred materials with the desired solubility parameters are the polylactides, polycaprolactones, and copolymers of these with glycolide in which there are more amorphous regions to enhance solubility.
It is also preferred that the solvent for the biodegradable polymer be non-toxic, water miscible, and otherwise biocompatible. Solvents that are toxic should not be used to inject any material into a living body. The solvents must also be biocompatible so that they do not cause severe tissue irritation or necrosis at the site of implantation. Furthermore, the solvent should be water miscible so that it will diffuse quickly into the body fluids and allow water to permeate into the polymer solution and cause it to coagulate or solidify. Examples of such solvents include N-methyl-2-pyrrolidone, 2-pyrrolidone, ethanol, propylene glycol, acetone, methyl acetate, ethyl acetate, methyl ethyl ketone, dimethylformamide, dimethyl sulfoxide, tetrahydrofuran, caprolactam, decylmethylsulfoxide, oleic acid, and 1-dodecylazacycloheptan-2-one. The preferred solvents are N-methyl-2-pyrrolidone, 2-pyrrolidone, dimethyl sulfoxide, and acetone because of their solrating solvating ability and their compatibility.
The solubility of the biodegradable polymers in the various solvents will differ depending upon their crystallinity, their hydrophilicity, hydrogen-bonding, and molecular weight. Thus, not all of the biodegradable polymers will be soluble in the same solvent, but each polymer or copolymer should have its optimum solvent. Lower molecular-weight polymers will normally dissolve more readily in the solvents than high-molecular-weight polymers. As a result, the concentration of a polymer dissolved in the various solvents will differ depending upon type of polymer and its molecular weight. Conversely, the higher molecular-weight polymers will normally tend to coagulate or solidify faster than the very low-molecular-weight polymers. Moreover the higher molecular-weight polymers will tend to give higher solution viscosities than the low-molecular-weight materials. Thus for optimum injection efficiency, the molecular weight and the concentration of the polymer in the solvent have to be controlled.
For example, low-molecular-weight polylactic acid formed by the condensation of lactic acid will dissolve in N-methyl-2-pyrrolidone(NMP) to give a 73% by weight solution which still flows easily through a 23-gauge syringe needle, whereas a higher molecular-weight poly(DL-lactide) (DL-PLA) formed by the additional polymerization of DL-lactide gives the same solution viscosity when dissolved in NMP at only 50% by weight. The higher molecular-weight polymer solution coagulates immediately when placed into water. The low-molecular-weight polymer solution, although more concentrated, tends to coagulate very slowly when placed into water.
For polymers that tend to coagulate slowly, a solvent mixture can be used to increase the coagulation rate. Thus one liquid component of the mixture is a good solvent for the polymer, and the other component is a poorer solvent or a non-solvent. The two liquids are mixed at a ratio such that the polymer is still soluble but precipitates with the slightest increase in the amount of non-solvent, such as water in a physiological environment. By necessity, the solvent system must be miscible with both the polymer and water. An example of such a binary solvent system is the use of NMP and ethanol for low-molecular-weight DL-PLA. The addition of ethanol to the NMP/polymer solution increases its coagulation rate significantly.
It has also been found that solutions containing very high concentrations of high-molecular-weight polymers sometimes coagulate or solidify slower than more dilute solutions. It is suspected that the high concentration of polymer impedes the diffusion of solvent from within the polymer matrix and consequently prevents the permeation of water into the matrix where it can precipitate the polymer chains. Thus, there is an optimum concentration at which the solvent can diffuse out of the polymer solution and water penetrates within to coagulate the polymer.
In one envisioned use of the thermoplastic system, the polymer solution is placed in a syringe and injected through a needle into the body. Once in place, the solvent dissipates, the remaining polymer solidifies, and a solid structure is formed. The implant will adhere to its surrounding tissue or bone by mechanical forces and can assume the shape of its surrounding cavity. Thus, the biodegradable polymer solution can be injected subdermally like collagen to build up tissue or to fill in defects. It can also be injected into wounds including burn wounds to prevent the formation Of deep scars. Unlike collagen, the degradation time of the implant can be varied from a few weeks to years depending upon the polymer selected and its molecular weight. The injectable polymer solution can also be used to mend bone defects or to provide a continuous matrix when other solid biodegradable implants such as hydroxyapatite plugs are inserted into bone gaps. The injectable system can also be used to adhere tissue to tissue or other implants to tissue by virtue of its mechanical bonding or encapsulation of tissue and prosthetic devices.
Another envisioned use of the thermoplastic system is to provide a drug-delivery system. In this use, a bioactive agent is added to the polymer solution prior to injection, and then the polymer/solvent/agent mixture is injected into the body. In some cases, the drug will also be soluble in the solvent, and a homogeneous solution of polymer and drug will be available for injection. In other cases, the drug will not be soluble in the solvent, and a suspension or dispersion of the drug in the polymer solution will result. This suspension or dispersion can also be injected into the body. In either case, the solvent will dissipate and the polymer will solidify and entrap or encase the drug within the solid matrix. The release of drug from these solid implants will follow the same general rules for release of a drug from a monolithic polymeric device. The release of drug can be affected by the size and shape of the implant, the loading of drug within the implant, the permeability factors involving the drug and the particular polymer, and the degradation of the polymer. Depending upon the bioactive agent selected for delivery, the above parameters can be adjusted by one skilled in the art of drug delivery to give the desired rate and duration of release.
The term drug or bioactive (biologically active) agent as used herein includes without limitation physiologically or pharmacologically active substances that act locally or systemically in the body. Representative drugs and biologically active agents to be used with the syringeable, in-situ forming solid implant systems include, without limitation, peptide drugs, protein drugs, desensitizing agents, antigens, vaccines, anti-infectives, antibiotics, antimicrobials, antiallergenics, steroidal anti-inflammatory agents, antiallergenics, steroidal anti-inflammatory agents, decongestants, miotics, anticholinergios, sympathomimetics, sedatives, hypnotics, psychic energizers, tranquilizers, androgenic steroids, estrogens, progestational agents, humoral agents, prostaglandins, analgesics, antispasmodics, antimalarials, antihistamines, cardioactive agents, non-steroidal anti-inflammatory agents, antiparkinsonian agents, antihypertensive agents, β-adrenergic blocking agents, nutritional agents, and the benzophenanthridine alkaloids. To those skilled in the art, other drugs or biologically active agents that can be released in an aqueous environment can be utilized in the described injectable delivery system. Also, various forms of the drugs or biologically active agents may be used. These include without limitation forms such as uncharged molecules, molecular complexes, salts, ethers, esters, amides, etc., which are biologically activated when injected into the body.
The amount of drug or biologically active agent incorporated into the injectable, in-situ, solid forming implant depends upon the desired release profile, the concentration of drug required for a biological effect, and the length of time that the drug has to be released for treatment. There is no critical upper limit on the amount of drug incorporated into the polymer solution except for that of an acceptable solution or dispersion viscosity for injection through a syringe needle. The lower limit of drug incorporated into the delivery system is dependent simply upon the activity of the drug and the length of time needed for treatment.
In all cases, the solid implant formed within the injectable polymer solution will slowly biodegrade within the body and allow natural tissue to grow and replace the impact as it disappears. Thus, when the material is injected into a soft-tissue defect, it will fill that defect and provide a scaffold for natural collagen tissue to grow. This collagen tissue will gradually replace the biodegradable polymer. With hard tissue such as bone, the biodegradable polymer will support the growth of new bone cells which will also gradually replace the degrading polymer. For drug-delivery systems, the solid implant formed from the injectable system will release the drug contained within its matrix at a controlled rate until the drug is depleted. With certain drugs, the polymer will degrade after the drug has been completely released. With other drugs such as peptides or proteins, the drug will be completely released only after the polymer has degraded to a point where the non-diffusing drug has been exposed to the body fluids.
B. Thermosetting System
The injectable, in-situ forming biodegradable implants can also be produced by crosslinking appropriately functionalized biodegradable polymers. The thermosetting system comprises reactive, liquid, oligomeric polymers which cure in place to form solids, usually with the addition of a curing catalyst. Although any of the biodegradable polymers previously described for the thermoplastic system can be used, the limiting criteria is that low-molecular-weight oligomers of these polymers or copolymers must be liquids and they must have functional groups on the ends of the prepolymer which can be reacted with acryloyl chloride to produce acrylic ester capped prepolymers.
The preferred biodegradable system is that produced from poly(DL-lactide-co-caprolactone), or “DL-PLC”. Low-molecular-weight polymers or oligomers produced from these materials are flowable liquids at room temperature. Hydroxy-terminated PLC prepolymers may be synthesized via copolymerization of DL-lactide or L-lactide and ε-caprolactone with a multifunctional polyol initiator and a catalyst. Catalysts useful for the preparation of these prepolymers are preferably basic or neutral ester-interchange (transesterification) catalysts. Metallic esters of carboxylic acids containing up to 18 carbon atoms such as formic, acetic, lauric, stearic, and benzoic are normally used as such catalysts. Stannous octoate and stannous chloride are the preferred catalysts, both for reasons of FDA compliance and performance.
If a bifunctional polyester is desired, a bifunctional chain initiator such as ethylene glycol is employed. A trifunctional initiator such as trimethylolpropane produces a trifunctional polymer, etc. The amount of chain initiator used determines the resultant molecular weight of the polymer or copolymer. At high concentration of chain initiator, the assumption is made that one bifunctional initiator molecule initiates only one polymer chain. On the other hand, when the concentration of bifunctional initiator is very low, each initiator molecule can initiate two polymer chains. In any case, the polymer chains are terminated by hydroxyl groups, as seen in FIG. 1. In this example, the assumption has been made that only one polymer chain is initiated per bifunctional initiator molecule. This assumption allows the calculation of a theoretical molecular weight for the prepolymers.
A list of the bifunctional PLC prepolymers that were synthesized is given in Table 1. Appropriate amounts of DL-lactide, ε-caprolactone, and ethylene glycol were combined in a flask under nitrogen and then heated in an oil bath at 155° C. to melt and mix the monomers. The copolymerizations were then catalyzed by the addition of 0.03 to 0.05 wt % SnCl2. The reaction was allowed to proceed overnight. The hydroxyl numbers of the prepolymers were determined by standard titration procedure. The Gardner-Holdt viscosities of the liquid prepolymers Were also determined using the procedures outlined in ASTM D 1545. The highest molecular-weight prepolymer (MW=5000) was a solid at room temperature; therefore, its Gardner-Holdt viscosity could not be determined.
The diol prepolymers were converted to acrylic-ester-capped prepolymers via a reaction with acryloyl chloride under Schotten-Baumann-like conditions, as seen in FIG. 2 and summarized in Table 2. Other methods of converting the diol prepolymers to acrylic-ester-capped prepolymers may also be employed.
Both THF and dichloromethane were evaluated as solvents in the acylation reactions. Several problems were encountered when THF was used as the solvent. The triethylamine hydrochloride formed as a by-product in the reaction was so finely divided that it could not be efficiently removed from the reaction mixture by filtration. Triethylamine hydrochloride (Et3N-HCl) has been reported to cause polymerization of acrylic species (U.S. Pat. No. 4,405,798). In several instances, where attempts to remove all of the Et3N.HCl failed, the acrylic-ester-capped prepolymers gelled prematurely. Thus, to effectively remove all of the Et3N.HCl, it was necessary to extract the prepolymers with water. For reactions carried out in THF, it is preferred that one first evaporate the THF in vacuo, redissolve the oil in CH2Cl2, filter out the Et3N-HCl, and then extract the CH2Cl2 layer with water. Stable emulsions were sometimes encountered during extraction. The acylations were later carried out in CH2Cl2 instead of THF. The filtration of Et3N.HCl from the reaction mixture was found to be much easier using this solvent, and the organic fraction could be extracted directly with water after filtration.
Both diol and acrylic prepolymers were examined by IR and 1H NMR spectroscopy. The salient feature of the IR spectra of diol prepolymers is a prominent O—H stretch centered at approximately 3510 cm−1. Upon acylation, the intensity of the O—H stretch decreases markedly, and new absorbances at approximately 1640 cm−1 appear. These new absorbances are attributed to the C—C stretch associated with acrylic groups. Likewise, the presences of acrylic ester groups is apparent in the 1H NMR spectra, the characteristic resonances for the vinyl protons falling in the range of 5.9 to 6.6 ppm.
The acrylic prepolymers and diol prepolymers were then cured, as summarized in Table 3. The general procedure for the curing of the prepolymers is now described: to 5.0 g of acrylic prepolymer contained in a small beaker was added a solution of benzoyl peroxide (BP) in approximately 1 mL of CH2Cl2. In some cases, fillers or additional acrylic monomers were added to the prepolymers prior to the introduction of the BP solution. The mixtures were stirred thoroughly and then poured into small petri dishes. The dishes were placed in a preheated vacuum oven for curing. Some of the samples were cured in air and not in vacuo, and these samples are so indicated in Table 3.
This thermosetting system may be used wherever a biodegradable implant is desired. For example, because the prepolymer remains a liquid for a short time after addition of the curing agent, the liquid prepolymer/curing agent mixture may be placed into a sytringe and injected into a body. The mxiture than solidifies in-situ, thereby providing an implant without an incision. Furthermore, a drug-delivery system may be provided by adding a biologically active agent to the prepolymer prior to injection. Once in-situ, the system will cure to a solid; eventually, it will biodegrade, and the agent will be gradually released.
The following examples are set forth as representative of the present invention. These examples are not to be construed as limiting the scope of the invention as these and other equivalent embodiments will be apparent in view of the present disclosure, figures, and accompanying claims.
Poly(DL-lactic acid) was prepared by the simple poly-condensation of lactic acid. No catalysts were used, and the reaction times were varied to produce polymers with different theoretical molecular weights. These polymers were designated as DL-PLA oligomers. A quantity of the solid oligomer was dissolved in NMP to give a 68:32 ratio of polymer to solvent. Sanguinarine chloride(SaCl), a benzophenanthridine alkaloid with antimicrobial activity especially toward periodontal pathogens, was added to the polymer solution to give a 2% by weight dispersion of the drug in the total mixture. The dispersion of drug and polymer solution was then injected into a dialysis tube (diameter of 1.5 mm) with a sterile disposable syringe without a needle. Each end of the 6-in. length of dialysis tubing was tied with a knot to prevent loss of the drug/polymer mass, and the tube with the injected material was placed in a pH 7 Sorenson's buffer receiving fluid maintained at 37° C. Upon immersion in the receiving fluid, the drug/polymer mass coagulated into a solid mass, and the drug began to be released from the polymer as indicated by an orange-red color in the receiving fluid. The quantity of solution injected into the dialysis tube was about 250 μL or about 100 mg of solids.
The dialysis tubing was selected to have a molecular-weight cutoff of about 3,500. With this molecular-weight cutoff, the SaCl released from the polymer could easily diffuse through the walls of the tubing, but any solid polymer would be retained. The dialysis tubing containing the drug/polymer matrix was removed frequently and placed in a bottle of fresh receiving fluid. The old receiving fluid containing the released drug was then acidified to a pH of 2.76 to convert all released drug to the iminium ion form of the drug, and the concentration of drug was determined by measuring the ultraviolet absorption (UV) at a wavelength of 237 nm. The cumulative mass of drug released and the cumulative fraction were then calculated and plotted as a function of time. Approximately 60% of the drug was released in the first day, 72% after 2 days, 85% after 5 days, 90% after 9 days, and 97% after 14 days.
Ethoxydihydrosanguinarine(SaEt), the ethanol ester of sanguinarine, was added to the same DL-PLA oligomer/NMP solution described in Example 1. SaEt dissolved in the polymer solution to give a homogeneous solution of drug and polymer. Approximately 250 μL of the solution was added to receiving fluid and the release of drug measured as described in Example 1. The release of SaEt was slower than that for SaCl as expected because of its lower water solubility. After the first day, approximately 45% was released, 52% after 2 days, 60% after 5 days, 70% after 9 days, and 80% after 14 days.
Poly(DL-lactide) with an inherent viscosity of 0.08 dL/g and a theoretical molecular weight of 2,000 was prepared by the ring-opening polymerization of DL-lactide using lauryl alcohol as the initiator and stannous chloride as the catalyst. This polymer was then dissolved in NMP to give a 40% by weight polymer solution. SaCl was dispersed in the solution of this polymer in NMP to give a 1.5% by weight dispersion of the drug in the solution and the release rate determined as described in Example 1. The release rate of the drug from this higher molecular-weight polymer was slower than from the DL-PLA oligomer. After the first day, approximately 32% was released, 40% after 2 days, 45% after 5 days, and 50% after 15 days.
SaEt was added to the same polymer solution of DL-PLA in NMP as described in Example 3. A homogeneous solution with the drug at 1.5% by weight was obtained. The release of drug from this solution determined using the same procedure described in Example 1 gave a much slower release of SaEt than from the DL-PLA oligomer. After the first day approximately 8% was released, 14% after 2 days, 20% after 5 days, 23% after 9 days, and 28% after 14 days.
The effect of drug loading on the release of drug from the polymer solutions were demonstrated by adding SaCl to a 40% by weight of DL-PLA oligomer in NMP. The drug was dispersed in the polymer solution to give 2, 7 and 14% by weight dispersions. The release of drug from these formulations using the same procedure as described in Example 1 showed that the higher drug loading gave a lower fractional rate of release as normally obtained for matrix delivery systems with diffusional release. The 2% -loaded formulation gave 65% release after 1 day, 75% after 2 days, and 88% after 5 days; the 7%-loaded formulation gave 48% release after 1 day, 52% after 2 days, and 58% after 5 days; and the 14%-loaded formulation gave 38% release after 1 day, 43% after 2 days, and 49% after 5days.
Poly(DL-lactide-co-glycolide) was prepared by the ring-opening polymerization of a mixture of DL-lactide and glycolide using lauryl alcohol as the initiator and stannous chloride as the catalyst. The proportions of the two monomers were adjusted so that the final copolymer(DL-PLG) had a 50:50 ratio of the two monomers as determined by nuclear magnetic resonance spectrophotometry. The initiator was also adjusted to give a copolymer with a theoretical molecular weight of 1500 daltons. The copolymer was dissolved in NMP to give a 70% by weight polymer solution. SaCl was added to this solution to give a 2% by weight dispersion of the drug in the polymer solution. The release of drug from this formulation was determined using the same procedure described in Example 1. A much lower release rate was obtained from the copolymer than from the DL-PLA oligomer of DL-PLA 2000 molecular weight materials. After 2 days approximately 7% of the drug was released, 10% after 5 days, 12% after 7 days, and 16% after 14 days.
SaEt was added to the same solution of DL-PLG in NMP as described in Example 6 to give a 2% by weight solution of the drug. The release of drug from this formulation was determined by the same procedure as described previously. The release rate of SaEt from this formulation was identical to that for SaCl described in Example 6.
Tetracycline as the free base (TCB) was added to the same solution of DL-PLG in NMP as described in Example 6. The drug dissolved completely in the polymer solution to give a 2.4% by weight solution to the drug. The release of the drug from this formulation was determined by a similar procedure to that described in Example 1 except the receiving fluid was not acidified to a pH of 2.76 and the concentration of TCB was determined by UV absorption at the wavelength appropriate for the drug. The release of TCB from this formulation was more linear and at a much higher rate than that for SaCl or SaEt from the same copolymer. After 1 day approximately 44% of the drug was released, 54% after 2 days, 68% after 5 days, 73% after 6 days, 80% after 7 days, 87% after 9 days, 96% after 12 days, and 100% after 14 days.
Tetracycline as the hydrochloride salt (TCH) was added to the same solution of DL-PLG in NMP as described in Example 6. The salt form of the drug also dissolved completely in the polymer solution. The release of drug from this formulation was determined as described in Example 8 and found to be similar to that for the free base except for a slightly lower rate. After 1 day approximately 32% of the drug was released, 40% after 2 days, 57% after 5 days, 64% after 6 days, 75% after 7 days, 82% after 9 days, 92% after 12 days, and 100% after 14 days.
DL-PLA with an inherent viscosity of 0.26 dL/g and a theoretical molecular weight of approximately 10,000 daltons was prepared by the ring-opening polymerization of DL-lactide using lauryl alcohol as the initiator and stannous chloride at the catalyst. The polymer was dissolved in NMP to give a 50% by weight polymer solution. A quantity of the polymer solution (100 μL) was injected subdermally into rabbits, and the tissue reaction was compared to that of a USP negative plastic. The test sites were evaluated for signs of local irritation, in accordance with the Draize method, immediately after injection, at 1 and 6 hours post injection, and once daily thereafter until scheduled sacrifice at 7, 14 or 21 days. The reaction at the test sites was equivalent to that at the control USP negative plastic. The polymer solution (100 μL) was also administered subgingivally into sites created by dental extractions in Beagle dogs. Control sites were flushed with saline solution. The dogs were examined daily for signs of mortality, pharmacotoxic effects, body weights, and local gingival irritation. The animals were sacrificed at 15 and 21 days. No distinct differences were noted between the control and test sites.
DL-PLA with an inherent viscosity of 0126 dL/g and a molecular weight of about 10,000 was dissolved in NMP to give a 50% by weight polymer solution. SaCl was added to the polymer solution to give a 2.4% by weight dispersion. This material was loaded into a 1-cc disposable syringe fitted with a 23-gauge blunted-end syringe needle, and the material was inserted into the periodontal pocket of a greyhound dog. The material flowed easily out of the narrow syringe tip. The polymer precipitated or coagulated into a film or solid mass when it contacted the saliva and fluid within the pocket. The dog was observed over a time of 2 weeks during which the mass of material remained within the pocket, adhering to tissue surrounding the pocket, and slowly changing color from a light orange to a pale white. The crevicular fluid from the pocket containing the implant was sampled during this 2-week period using Periostrips which are small strips of paper that are placed at the entrance to the periodontal pocket to wick up small quantities of the crevicular fluid within the pocket. The volume of fluid collected is determined using a Periotron which measures the changes in conductance of the paper strip. The Periotron is calibrated before use with a known volume of serum. The paper strip containing the collected fluid is then extracted with a solution of 0.5% by volume of hydrochloric acid in methanol and injected into a liquid chromatograph where the quantity of drug is determined by reference to a known concentration of the same compound. The quantity of SaCl extracted from the paper strip is divided by the quantity of crevicular fluid collected to calculate the concentration of drug in the fluid. With this technique, the concentration of SaCl within the crevicular fluid from the periodontal pocket with the polymeric delivery system was determined to be almost constant during the 2 weeks of observation. The SaCl concentration in the crevicular fluid was 63.2 μg/mL after 3 days, 80.2 μg/mL after 7 days, 67.8 μg/mL after 10 days, and 70.5 μg/mL after 14 days.
An illustrative method for the synthesis of an acrylate terminated prepolymer is described. To an oven-dried, 500-mL, three-necked, round-bottom flask fitted with an addition funnel, gas inlet adapter, mechanical stirrer assembly, and rubber septum was added, under nitrogen, 100.0 g of difunctional hydroxy-terminated prepolymer and 200 mL of freshly distilled THF (from CaH2). The flask was cooled in an ice bath, and 24 mL of dry triethylamine (0.95 equiv/equiv OH) was added via a syringe. The addition funnel was charged with 15.4 g of acryloyl chloride (0.95 equiv/equiv OH) in 15 mL of THF, and the solution was added dropwise to the stirred reaction mixture over 1 hour. The mixture was stirred overnight and allowed to reach room temperature. The precipitated triethylamine hydrochloride was removed by filtration, and the filtrate was evaporated in vacuo, affording a pale yellow oil, which was the acrylate-terminated prepolymer. The acylations employing CH2Cl2 as solvent were conducted in a similar manner. However, the reaction times at 0° C. were shortened to 1 hour, whereupon the reaction mixtures were allowed to reach room temperature over 1 hour. Et3N1HCl was filtered out, additional CH2Cl2 (approximately 800 mL) was added to the filtrate, and the filtrate was extracted several times with 250 mL portions of water. The organic layer was dried over MgSO4/Na2SO4, filtered, and reduced to an oil in vacuo. The bottles of acrylic prepolymers were wrapped in foil and stored in a refrigerator to safeguard against premature crosslinking.
TABLE 1 |
SUMMARY OF DIOL PREPOLYMERS SYNTHESIZED |
Mole ratio of monomers | Gardiner Holdt | ||||
to initiator | Catalyst | Hydroxyl No., | viscosity, | ||
(ethylene glycol = 1.0) | (SnCl2), | Theoretical | meq OH (56.1)/g | approx. Stokes |
Sample no. | DL-lactide | ε-caprolactone | wt % | Mn, daltons | Observed | Theoretical | (T = 22.2° C.) |
C964-114-1 | 2.4 | 5.0 | 0.03 | 993 | 100 | 113 | 28.0 |
C964-124-1 | 6.1 | 32.8 | 0.05 | 5036 | 19.7 | 22.3 | Solid |
C964-128-1 | 2.5 | 5.0 | 0.03 | 993 | 103 | 113 | 28.2 |
C964-136-1 | 8.0 | 8.0 | 0.03 | 2128 | 48 (est.) | 52.7 | 1375 |
TABLE 2 |
SUMMARY OF ACRYLIC ESTER TERMINATED PREPOLYMERS SYNTHESIZED |
Estimated | ||||
Diol | concentration | |||
precursor | of acrylic groups, | Reaction conditions |
Sample no. | sample no. | meq/g | Temp, ° C. | Time, h | Solvent | Comments |
C964-118-1 | C964-114-1 | 1.78 | O-RT | 17 | THF | No problems, stable. |
C964-125-1 | C964-114-1 | 1.78 | O-RT | 17 | THF | Gelled. Overnight exposure to |
Et3N.HCl at RT. | ||||||
C964-132-1 | C964-128-1 | 1.84 | 0 | 2 | THF | 100 ppm MEHQ added before |
workup. | ||||||
C964-137-1 | C964-128-1 | 1.84 | 0 | 2 | Et2O | Difficult workup. Low yield. |
C964-139-1 | C964-136-1 | 0.81 | 0 | 2 | THF | Gelled. Overnight exposure to |
residual Et3N.HCl in | ||||||
refrigerator. | ||||||
C964-144-1 | C964-136-1 | 0.81 | 0 | 1 | CH2Cl2 | No problems, stable. |
C964-146-1 | C964-124-1 | 0.33 | 0 | 1 | CH2Cl2 | No problems, stable. |
TABLE 3 |
SUMMARY OF CURING STUDIES |
Acrylic | Benzoyl | Other | Initial | |||
prepolymer | peroxide | additives, | Curing conditions | Shore A |
Sample no. | sample no. | wt % | wt % | Temp, ° C. | Time, h | hardness | Comments |
C964-120-1 | C964-118-1 | 2.0 | none | 82 | 16 | NDa | Rubbery, breaks when bent 180°, weak. |
C964-120-2 | C964-118-1 | 1.0 | none | 82 | 16 | 83 | Less brittle than C964-120-1. |
C964-121-1 | C964-118-1 | 2.0 | none | 82 | 16 | 77 | Rubbery, breaks when bent 180°, weak. |
C964-121-2 | C964-118-1 | 1.0 | none | 82 | 16 | 80 | Slightly stronger than C964-121-1. |
C964-121-3 | C964-118-1 | 0.5 | none | 82 | 16 | 78 | Slightly more elastic than C964-121-2. |
C964-121-4 | C964-118-1 | 0.1 | none | 82 | 16 | 69 | Same as C964-121-3. |
C964-122-1 | C964-118-1 | 1.0 | TMPTETAb 46 | 82 | 2.5 | 94 | Less rubbery than C964-120 and C964-121, |
brittle. | |||||||
C964-122-2 | C964-118-1 | 0.5 | TMPTETA 46 | 82 | 2.5 | 91 | Same as C964-122-1, more flexible. |
C964-122-3 | C964-118-1 | 1.0 | TMPTETA 175 | 82 | 2.5 | 95 | Not rubbert at all, brittle, weak. |
C964-122-4 | C964-118-1 | 0.5 | TMPTETA 175 | 82 | 2.5 | 93 | Similar to C964-122-3. |
C964-123-1 | C964-118-1 | 0.1 | TMPTETA 46 | 82 | 2.5 | 89 | Rubbery, stronger than C964-120 and C964-121, |
not flexible. | |||||||
C964-123-2 | C964-118-1 | 0.25 | TMPTETA 46 | 82 | 2.5 | 83 | About the same as C964-123-1, may be more |
brittle. | |||||||
C964-123-3 | C964-132-1 | 0.1 | TMPTETA 175 | 82 | 2.5 | 92 | Not rubbery, strong, brittle. |
C964-134-1 | C964-132-1 | 0.05 (AIBN)c | none | 60d | 17 | Liquid | No cure. |
C964-134-2 | C964-132-1 | 0.10 (AIBN) | none | 60d | 17 | Liquid | No cure. |
C964-134-3 | C964-132-1 | 0.25 (AIBN) | none | 60d | 17 | Liquid | No cure. |
C964-134-4 | C964-132-1 | 0.50 (AIBN) | none | 60d | 17 | Liquid | No cure. |
C964-134-5 | C964-132-1 | 1.00 (AIBN) | none | 60d | 17 | Liquid | Slightly thickened. |
C964-135-1 | C964-132-1 | 0.05 | none | 80d | 17 | Liquid | No cure. |
C964-135-2 | C964-132-1 | 0.10 | none | 80d | 17 | Liquid | No cure. |
C964-135-3 | C964-132-1 | 0.25 | none | 80d | 17 | Liquid | No cure. |
C964-135-4 | C964-132-1 | 0.50 | none | 80d | 17 | Liquid | No cure. |
C964-135-5 | C964-132-1 | 1.00 | none | 80d | 17 | Liquid | Slightly thickened. |
C964-135-6 | C964-128-1e | 0.05 | none | 80d | 17 | Liquid | No cure. |
C964-135-7 | C964-128-1e | 0.10 | none | 80d | 17 | Liquid | No cure. |
C964-135-8 | C964-128-1e | 0.25 | none | 80d | 17 | Liquid | No cure. |
C964-135-9 | C964-128-1e | 0.50 | none | 80d | 17 | Liquid | No cure. |
C964-135-10 | C964-128-1e | 1.00 | none | 80d | 17 | Liquid | No cure. |
C964-135-11 | C964-124-1e | 0.05 | none | 80d | 17 | ND | No cure. |
C964-135-12 | C964-124-1e | 0.10 | none | 80d | 17 | ND | No cure. |
C964-135-13 | C964-124-1e | 0.25 | none | 80d | 17 | ND | No cure. |
C964-135-14 | C964-124-1e | 0.50 | none | 80d | 17 | ND | No cure. |
C964-135-15 | C964-124-1e | 1.00 | none | 80d | 17 | ND | No cure. |
C964-141-1 | C964-137-1 | 0.10 | none | 80 | 1 | 66 | Flexible elastomer. |
C964-141-2 | C964-137-1 | 0.25 | none | 80 | 1 | 71 | Flexible elastomer. |
C964-141-3 | C964-137-1 | 0.50 | none | 80 | 1 | 72 | Flexible elastomer. |
C964-141-4 | C964-137-1 | 1.00 | none | 80 | 1 | 72 | Flexible elastomer. |
C964-141-5 | C964-128-1 | 0.10 | none | 80 | 1 | Liquid | No cure. |
C964-141-6 | C964-128-1 | 0.25 | none | 80 | 1 | Liquid | No cure. |
C964-141-7 | C964-128-1 | 0.50 | none | 80 | 1 | Liquid | No cure. |
C964-141-8 | C964-128-1 | 1.00 | none | 80 | 1 | Liquid | No cure. |
C964-143-1 | C964-137-1 | 0.25 | Cab-o-Sil PTG, | 80 | 1 | 74 | No cure. |
5.0 | |||||||
C964-143-2 | C964-137-1 | 0.25 | Cab-o-Sil PTG, | 80 | 1 | 73 | No cure. |
2.5 | |||||||
C964-143-3 | C964-137-1 | 0.25 | L-PLA | 80 | 1 | 75 | No cure. |
(IV = 0.8), 5.0 | |||||||
C964-143-4 | C964-137-1 | 0.25 | L-PLA | 80 | 1 | 78 | No cure. |
(IV = 0.8), 2.5 | |||||||
C964-148-1 | C964-144-1 | 0.05 | none | 80 | 17 | Liquid | No cure. |
C964-148-2 | C964-144-1 | 0.10 | none | 80 | 17 | Liquid | No cure. |
C964-148-3 | C964-144-1 | 0.25 | none | 80 | 2 | 66 | C964-148-4 and C964-148-6 were about the |
same in toughness, and both were better than | |||||||
C964-148-3 and C964-148-5. | |||||||
C964-148-4 | C964-144-1 | 0.50 | none | 80 | 2 | 68 | C964-148-4 and C964-148-6 were about the |
same in toughness, and both were better than | |||||||
C964-148-3 and C964-148-5. | |||||||
C964-148-5 | C964-144-1 | 1.00 | none | 80 | 2 | 67 | C964-148-4 and C964-148-4 were about the |
same in toughness, and both were better than | |||||||
C964-148-3 and C964-148-5. | |||||||
C964-148-6 | C964-144-1 | 2.00 | none | 80 | 2 | 69 | C964-148-4 and C964-148-4 were about the |
same in toughness, and both were better than | |||||||
C964-148-3 and C-964-148-5. | |||||||
C964-149-1 | C964-144-1 | 0.15 | none | 80 | 2 | 64 | |
C964-149-2 | C964-144-1 | 0.20 | none | 80 | 2 | 64 | |
C964-149-3 | C964-144-1 | 0.25 | none | 80 | 2 | 66 | |
C964-149-4 | C964-144-1 | 0.15 | Cab-o-Sil N70-TS | 80 | 2 | ND | Samples too porous, did not have any flat area |
5.0 | for hardness measurement. | ||||||
C964-149-5 | C964-144-1 | 0.20 | Cab-o-Sil N70-TS | 80 | 2 | ND | Samples too porous, did not have any flat area |
5.0 | for hardness measurement. | ||||||
C964-149-6 | C964-144-1 | 0.25 | Cab-o-Sil N70-TS | 80 | 2 | ND | Samples too porous, did not have any flat area |
5.0 | for hardness measurement. | ||||||
C964-150-1 | C964-146-1 | 0.05 | none | 80 | 17 | ND | Only partially cured. |
C964-150-2 | C964-146-1 | 0.10 | none | 80 | 2 | 72 | Elastic, flexible, moderately strong. |
C964-150-3 | C964-146-1 | 0.25 | none | 80 | 2 | 57 | Elastic, flexible, moderately strong. |
C964-150-4 | C964-146-1 | 0.50 | none | 80 | 2 | 56 | Elastic, flexible, moderately strong. |
C964-150-5 | C964-146-1 | 1.00 | none | 80 | 2 | 50 | Elastic, flexible, moderately strong. |
C964-150-6 | C964-146-1 | 2.00 | none | 80 | 2 | 51 | Elastic, flexible, moderately strong. |
aResult not determined. | |||||||
bTMPTETA = trimethylolpropane triethoxy triacrylate. | |||||||
cAIBN = azobisisobutyronitrile. | |||||||
dCured in air at atmospheric pressure. | |||||||
eDiol prepolymer used. |
Claims (36)
1. A method of forming a solid, biodegradable implant in-situ in a body, comprising:
(a) dissolving a solid, water-insoluble, biodegradable, linear thermoplastic polymer in a biocompatible, water-soluble organic solvent to form a flowable composition; the organic solvent being capable of dissolving the thermoplastic polymer and being capable of dissipating or diffusing into a body fluid upon placement within a body; and
(b) placing the composition into an implant site within the body; and
(c) allowing the organic solvent to dissipate or diffuse into body fluid, and the thermoplastic polymer to coagulate or solidify to produce the biodegradable solid implant, wherein the proportion of polymer in solvent and the polymer molecular weight are effective to provide said dissipating or diffusing function and said coagulating or solidifying function.
2. The method of claim 1 , wherein the solid, linear thermoplastic polymer is selected from the group consisting of polylactides, polyglycolides, polycaprolactones, polydioxanones, polycarbonates, polyhydroxybutyrates, polyalkylene oxalates, polyanhydrides, polyamides, polyesteramides, polyacetates polyacetals, polyketals, polyorthocarbonates, polyphosphazenes, polyhydroxyvalerates, polyalkylene succinates, poly(amino acids), polyorthoesters, and copolymers, terpolymers, or combinations and mixtures thereof, each named thermoplastic polymer, copolymer, terpolymer, combination or mixture being biodegradable.
3. The method of claim 1 , wherein the solid, linear thermoplastic polymer is selected from the group consisting of polylactides, polyglycolides, polycaprolactones, and copolymers thereof.
4. The method of claim 1 , wherein the organic solvent is selected from the group consisting of N-methyl-2-pyrrolidone, 2-pyrrolidone, acetone, methyl acetate, ethyl acetate, methyl ethyl ketone, dimethylformamide, dimethyl sulfoxide, tetrahydrofuran, caprolactam, decylmethylsulfoxide, oleic acid, and 1-dodecylazacycloheptan-2-one, and combinations and mixtures thereof.
5. The method of claim 1 , further comprising combining a biologically-active agent with the solid, linear thermoplastic polymer and organic solvent.
6. A pharmaceutical composition for forming a solid biodegradable implant in situ within a body, comprising:
an effective amount of a solid, water-insoluble, biodegradable, linear thermoplastic polymer dissolved in an effective amount of a biocompatible, water-soluble organic solvent, wherein the solvent is capable of dissolving the thermoplastic polymer and is capable of dissipating or diffusing into a body fluid and the polymer is capable of coagulating or solidifying upon contact with body fluid; and
wherein the pharmaceutical composition has a flowable consistency and the molecular weight of the polymer, the amounts of the polymer and the solvent in the pharmaceutical composition are effective to form the solid implant in situ when the pharmaceutical composition contacts body fluid.
7. The pharmaceutical composition of claim 6 , wherein the solid, linear thermoplastic polymer is selected from the group consisting of:
polylactides, polyglycolides, polycaprolactones, polydioxanones, polycarbonates, polyhydroxybutyrates, polyalkylene oxalates, polyanhydrides, polyamides, polyesteramides, polyacetates polyacetals, polyketals, polyorthocarbonates, polyphosphazenes, polyhydroxyvalerates., polyalkylene succinates, poly(amino acids), polyorthoesters, and copolymers, terpolymers, or combinations and mixtures thereof each named thermoplastic polymer, copolymer, terpolymer, combination and mixture being biodegradable.
8. The pharmaceutical composition of claim 6 , wherein the solid, linear thermoplastic polymer is selected from the group consisting of:
polylactides, polyglycolides, polycaprolactones, and copolymers thereof.
9. The pharmaceutical composition of claim 6 , wherein the organic solvent is selected from the group consisting of:
N-methyl-2-pyrrolidone, 2-pyrrolidone, acetone, methyl acetate, ethyl acetate, methyl ethyl ketone, dimethylformamide, dimethyl sulfoxide, tetrahydrofuran, caprolactam, decylmethylsulfoxide, oleic acid and 1-dodecylazacycloheptan-2-one, and combinations and mixtures thereof.
10. The pharmaceutical composition of claim 6 , further comprising a biologically-active agent combined with the solid, linear thermoplastic polymer and organic solvent.
11. The pharmaceutical composition of claim 10 , wherein the biologically-active agent is an antibiotic, antimicrobial, peptide drug, protein drug, or a combination thereof.
12. The pharmaceutical composition of claim 10 , wherein the biologically-active agent is a tetracycline.
13. The pharmaceutical composition of claim 10 , wherein the composition is a homogenous solution, a suspension, or a dispersion.
14. A method of delivering a biologically-active agent to an animal a living body comprising:
administering a pharmaceutical composition according to claim 10 to the animal living body, wherein the pharmaceutical composition forms a solid implant and the biologically-active agent is released from the implant into the animal living body.
15. An in situ formed biodegradable implant comprising a mixture of a water-insoluble, biodegradable, solid, linear thermoplastic polymer, a biologically active agent and a water-soluble, biocompatible organic solvent, wherein the implant is surrounded by a body tissue, at least some of the polymer is solidified and the remainder of the polymer is in solution with the solvent.
16. An implant according to claim 15 wherein the organic solvent has a water solubility, a biocompatibility and a thermoplastic polymer dissolubility exemplified by a compound selected from the group consisting of N-methyl pyrrolidone, 2 -pyrrolidone, ethanol, propylene glycol, acetone, methyl acetate, ethyl acetate, methyl ethyl ketone, dimethylformamide, dimethyl sulfoxide, tetrahydrofuran, caprolactam, decylmethyl sulfoxide, oleic acid and 1 -dodecylazacycloheptan- 2 -one.
17. An implant according to claim 16 wherein the compound is N-methyl pyrrolidone or dimethyl sulfoxide.
18. An implant according to claim 16 wherein the compound is ethyl acetate or oleic acid of 1-dodecylazacycloheptan- 2 -one.
19. An implant according to claim 15 wherein essentially all of the solid, linear thermoplastic polymer is solidified.
20. An implant according to claim 19 , wherein at least some organic solvent is present in the implant.
21. An implant according to claim 19 wherein essentially no organic solvent is present in the implant.
22. An implant according to claim 15 wherein the biodegradable, solid, linear thermoplastic polymer is selected from the group consisting of polylactides, polyglycolides, polycaprolactones, polydioxanones, polycarbonates, polyhydroxybutyrates, polyalkyleneoxalates, polyanhydrides, polyamides, polyesteramides, polyacetals, polyketals, polyorthocarbonates, polyphosphazenes, polyhydroxyvalerates, polyalkylenesuccinates, poly(amino acids), polyorthoesters, and copolymers, terpolymers, combinations and mixtures thereof.
23. An implant according to claim 22 wherein the linear, solid, thermoplastic polymer is a homopolymer or copolymer of lactide, glycolide, caprolactone or a combination thereof.
24. An implant according to claim 15 that is elastomeric.
25. An implant according to claim 15 that has a rigidity like that of hard tissue.
26. An implant according to claim 15 that is a coagulate.
27. An implant according to claim 15 that is a solid.
28. An implant according to claim 15 wherein the biologically active agent is an antibiotic, antimicrobial, peptide drug, protein drug, or a combination thereof.
29. And implant according to claim 15 wherein the biologically active agent is a peptide drug.
30. An implant according to claim 26 that has fully coagulated.
31. An implant according to claim 27 that has fully solidified.
32. An implant according to claim 15 wherein essentially all of the solid, linear, thermoplastic polymer is a solid and is elastomeric.
33. An implant according to claim 15 wherein essentially all of the solid, linear, thermoplastic polymer is a solid and is usable with hard tissue.
34. An in situ formed biodegradable implant comprising a mixture of a water insoluble, biodegradable, solid, linear thermoplastic polymer, and a water-soluble, biocompatible organic solvent, wherein the implant is surrounded by a body tissue, at least some of the polymer is solidified, and the remainder of the polymer is in solution with the solvent.
35. An in situ formed biodegradable implant comprising a mixture of a water insoluble, biodegradable, solid, linear thermoplastic polymer, and a biologically active agent, wherein the implant is surrounded by a body tissue and the polymer is a solid.
36. An in situ formed biodegradable implant comprising a mixture of a water insoluble, biodegradable, solid, linear thermoplastic polymer, and a biologically active agent, wherein the implant is surrounded by a body tissue and the polymer is a solid.
Priority Applications (1)
Application Number | Priority Date | Filing Date | Title |
---|---|---|---|
US09/536,579 USRE37950E1 (en) | 1990-04-24 | 2000-03-28 | Biogradable in-situ forming implants and methods of producing the same |
Applications Claiming Priority (6)
Application Number | Priority Date | Filing Date | Title |
---|---|---|---|
US07/513,782 US5278201A (en) | 1988-10-03 | 1990-04-24 | Biodegradable in-situ forming implants and methods of producing the same |
US07/788,032 US5340849A (en) | 1988-10-03 | 1991-12-23 | Biodegradable in-situ forming implants and methods for producing the same |
US08/210,891 US5739176A (en) | 1988-10-03 | 1994-03-18 | Biodegradable in-situ forming implants and methods of producing the same |
US08/532,540 US5733950A (en) | 1988-10-03 | 1995-09-25 | Biodegradable in-situ forming implants and methods of producing the same |
US25264598A | 1998-10-03 | 1998-10-03 | |
US09/536,579 USRE37950E1 (en) | 1990-04-24 | 2000-03-28 | Biogradable in-situ forming implants and methods of producing the same |
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Application Number | Title | Priority Date | Filing Date |
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US08/532,540 Reissue US5733950A (en) | 1988-10-03 | 1995-09-25 | Biodegradable in-situ forming implants and methods of producing the same |
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USRE37950E1 true USRE37950E1 (en) | 2002-12-31 |
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US09/536,579 Expired - Lifetime USRE37950E1 (en) | 1990-04-24 | 2000-03-28 | Biogradable in-situ forming implants and methods of producing the same |
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US20020161170A1 (en) * | 1999-07-30 | 2002-10-31 | Takehisa Matsuda | Coumarin endcapped absorbable polymers |
US6565601B2 (en) * | 2000-11-15 | 2003-05-20 | Micro Therapeutics, Inc. | Methods for vascular reconstruction of diseased arteries |
US20040234616A1 (en) * | 2003-03-04 | 2004-11-25 | The Technology Development Company Ltd. | Long acting injectable insulin composition and methods of making and using thereof |
US20060121085A1 (en) * | 2003-03-11 | 2006-06-08 | Warren Stephen L | Formulations for cell-schedule dependent anticancer agents |
WO2007084460A2 (en) | 2006-01-18 | 2007-07-26 | Qps, Llc | Pharmaceutical compositions with enhanced stability |
US20080020016A1 (en) * | 2006-07-11 | 2008-01-24 | Quest Pharmaceutical Services, Llc | Pharmaceutical compositions for sustained release delivery of peptides |
US20080069865A1 (en) * | 2004-01-13 | 2008-03-20 | Southard Jeffrey L | Controlled Release Cgrp Delivery Composition for Cardiovascular and Renal Indications |
US20090023643A1 (en) * | 2004-01-13 | 2009-01-22 | Vasogenix Pharmaceuticals, Inc. | Methods For Treating Acute Myocardial Infarction By Administering Calcitonin Gene Related Peptide And Compositions Containing The Same |
US20090023644A1 (en) * | 2004-01-13 | 2009-01-22 | Southard Jeffery L | Methods of using cgrp for cardiovascular and renal indications |
US20090198330A1 (en) * | 2008-02-01 | 2009-08-06 | Kesten Randy J | Compressible cosmetic prosthesis with improved shape retention |
US20090198328A1 (en) * | 2008-02-01 | 2009-08-06 | Kesten Randy J | Breast implant with adjustable compression response |
US20090305957A1 (en) * | 2006-10-11 | 2009-12-10 | Qlt Usa, Inc. | Preparation of biodegradable polyesters with low-burst properties by supercritical fluid extraction |
US20090325879A1 (en) * | 2008-06-03 | 2009-12-31 | Qlt Usa, Inc. | Controlled release copolymer formulation with improved release kinetics |
US20100160892A1 (en) * | 2008-12-23 | 2010-06-24 | Tice Thomas R | Implantable suction cup composites and implants comprising same |
US20100160891A1 (en) * | 2008-12-23 | 2010-06-24 | Tipton Arthur J | Elastic implantable composites and implants comprising same |
US20100292195A1 (en) * | 2007-02-15 | 2010-11-18 | Qlt Usa, Inc. | Low burst polymers and methods to produce polymer |
US20110027389A1 (en) * | 2008-01-14 | 2011-02-03 | Dunn Richard L | Low Viscosity Liquid Polymeric Delivery System |
US20110171305A1 (en) * | 2008-06-03 | 2011-07-14 | Qlt Usa Inc. | Dehydrated hydrogel inclusion complex of a bioactive agent with flowable drug delivery system |
US7998202B2 (en) | 2004-09-16 | 2011-08-16 | Evera Medical, Inc. | Tissue implant having a biased layer and compliance that simulates tissue |
US8048101B2 (en) | 2004-02-25 | 2011-11-01 | Femasys Inc. | Methods and devices for conduit occlusion |
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US8052669B2 (en) | 2004-02-25 | 2011-11-08 | Femasys Inc. | Methods and devices for delivery of compositions to conduits |
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US9238127B2 (en) | 2004-02-25 | 2016-01-19 | Femasys Inc. | Methods and devices for delivering to conduit |
US9259701B2 (en) | 2010-09-30 | 2016-02-16 | Q Chip Ltd. | Method for making solid beads |
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US9333070B2 (en) | 2008-02-01 | 2016-05-10 | Evera Medical, Inc. | Breast implant with internal flow dampening |
US9480643B2 (en) | 2008-12-23 | 2016-11-01 | Surmodics Pharmaceuticals, Inc. | Implantable composites and implants comprising same |
US9554826B2 (en) | 2008-10-03 | 2017-01-31 | Femasys, Inc. | Contrast agent injection system for sonographic imaging |
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US10563153B2 (en) | 2010-05-20 | 2020-02-18 | Ecolab Usa Inc. | Rheology modified low foaming liquid antimicrobial compositions and methods of use thereof |
US10758623B2 (en) | 2013-12-09 | 2020-09-01 | Durect Corporation | Pharmaceutically active agent complexes, polymer complexes, and compositions and methods involving the same |
US11717555B2 (en) | 2017-12-18 | 2023-08-08 | Foresee Pharmaceuticals Co., Ltd. | Pharmaceutical compositions having a selected release duration |
US11865205B2 (en) | 2015-11-16 | 2024-01-09 | Medincell S.A. | Method for morselizing and/or targeting pharmaceutically active principles to synovial tissue |
Citations (60)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
US2155658A (en) | 1936-01-08 | 1939-04-25 | Chemische Forschungs Gmbh | Surgical and medical preparations |
US3068188A (en) | 1957-03-27 | 1962-12-11 | Du Pont | Composition comprising a synthetic linear polymer, organic solvent, and an inorganic salt |
US3218283A (en) | 1962-07-26 | 1965-11-16 | Monsanto Co | Novel solutions of poly-(acrylic anhydride) and poly-(methacrylic anhydride) polymers |
US3219527A (en) | 1963-06-17 | 1965-11-23 | Loyola University | Peridontal pack or dressing composition |
US3328246A (en) | 1957-08-20 | 1967-06-27 | Bicrex Lab Ltd | Dental compositions and methods of making same |
US3463158A (en) | 1963-10-31 | 1969-08-26 | American Cyanamid Co | Polyglycolic acid prosthetic devices |
US3520949A (en) | 1966-07-26 | 1970-07-21 | Nat Patent Dev Corp | Hydrophilic polymers,articles and methods of making same |
FR2126270A1 (en) | 1971-02-23 | 1972-10-06 | Du Pont | |
US3696811A (en) | 1971-06-17 | 1972-10-10 | Squibb & Sons Inc | Periodontal bandage and backing therefor |
US3760034A (en) | 1971-01-26 | 1973-09-18 | Union Carbide Corp | Graft copolymers of lactone polyesters |
US3767784A (en) | 1970-12-01 | 1973-10-23 | S Gluck | Composition for the protection and treatment of injured body tissue and method of utilizing the same |
US3887699A (en) | 1969-03-24 | 1975-06-03 | Seymour Yolles | Biodegradable polymeric article for dispensing drugs |
US3931678A (en) | 1974-09-24 | 1976-01-13 | Loctite (Ireland) Limited | Dental filling method and composition formed thereby |
US3935308A (en) | 1974-08-08 | 1976-01-27 | The United States Of America As Represented By The Secretary Of The Navy | Wound covering and method of application |
US3939111A (en) | 1974-07-01 | 1976-02-17 | The B. F. Goodrich Company | Stable polyurethane solutions |
US4148871A (en) | 1977-10-11 | 1979-04-10 | Pitt Colin G | Sustained subdermal delivery ofdrugs using poly(ε-caprolactone) and its copolymers |
US4161948A (en) | 1977-01-18 | 1979-07-24 | Battelle Memorial Institute | Synthetic membrane for wound-dressings |
US4408023A (en) | 1980-11-12 | 1983-10-04 | Tyndale Plains-Hunter, Ltd. | Polyurethane diacrylate compositions useful for contact lenses and the like |
US4439420A (en) | 1982-11-16 | 1984-03-27 | Ethicon, Inc. | Absorbable hemostatic composition |
US4443430A (en) | 1982-11-16 | 1984-04-17 | Ethicon, Inc. | Synthetic absorbable hemostatic agent |
US4447562A (en) | 1981-07-15 | 1984-05-08 | Ivani Edward J | Amino-polysaccharides and copolymers thereof for contact lenses and ophthalmic compositions |
US4450150A (en) | 1973-05-17 | 1984-05-22 | Arthur D. Little, Inc. | Biodegradable, implantable drug delivery depots, and method for preparing and using the same |
US4491479A (en) | 1982-11-12 | 1985-01-01 | Lauchenauer Alfred E | Shaped semi-solid articles |
WO1985000969A1 (en) | 1983-08-31 | 1985-03-14 | S.K.Y. Polymers Ltd. | Injectable physiologically acceptable polymeric compositions |
EP0140766A2 (en) | 1983-10-07 | 1985-05-08 | The Forsyth Dental Infirmary For Children | Intra pocket drug delivery devices for treatment of periondontal diseases |
WO1985002092A1 (en) | 1983-11-14 | 1985-05-23 | Bio-Mimetics Inc. | Bioadhesive compositions and methods of treatment therewith |
EP0159293A1 (en) | 1984-04-02 | 1985-10-23 | Battelle Memorial Institute | Film-forming polymeric material for use in dressings, and its production |
EP0169016A2 (en) | 1984-07-16 | 1986-01-22 | Celtrix Pharmaceuticals, Inc. | Polypeptide cartilage-inducing factors found in bone |
US4568536A (en) | 1985-02-08 | 1986-02-04 | Ethicon, Inc. | Controlled release of pharmacologically active agents from an absorbable biologically compatible putty-like composition |
US4570629A (en) | 1982-03-17 | 1986-02-18 | University Of Illinois Foundation | Hydrophilic biopolymeric copolyelectrolytes, and biodegradable wound dressing comprising same |
US4582640A (en) | 1982-03-08 | 1986-04-15 | Collagen Corporation | Injectable cross-linked collagen implant material |
US4595713A (en) | 1985-01-22 | 1986-06-17 | Hexcel Corporation | Medical putty for tissue augmentation |
US4631188A (en) | 1983-08-31 | 1986-12-23 | S.K.Y. Polymers, Ltd. (Kingston Technologies) | Injectable physiologically-acceptable polymeric composition |
US4650665A (en) | 1985-02-08 | 1987-03-17 | Ethicon, Inc. | Controlled release of pharmacologically active agents from an absorbable biologically compatible putty-like composition |
US4663077A (en) | 1984-06-11 | 1987-05-05 | Morton Thiokol Inc. | Microbiocidal compositions comprising an aryl alkanol and a microbiocidal compound dissolved therein |
US4677139A (en) | 1982-04-07 | 1987-06-30 | Feinmann Bernhard P P | Material and method for dentistry |
EP0241178A1 (en) | 1986-03-25 | 1987-10-14 | Rohto Pharmaceutical Co., Ltd. | Pharmaceutical composition for treating periodontal diseases |
US4702917A (en) | 1985-11-18 | 1987-10-27 | Research Triangle Institute | Porous bioabsorbable polyesters |
EP0244118A1 (en) | 1986-04-29 | 1987-11-04 | Pharmetrix Corporation | Controlled release drug delivery system for the peridontal pocket |
US4715369A (en) | 1980-12-29 | 1987-12-29 | Teijin Limited | Method of treating an injured part on the oral mucosa and the covering material for use thereof |
US4745160A (en) | 1984-06-26 | 1988-05-17 | Imperial Chemical Industries Plc | Biodegradable amphipathic copolymers |
US4767627A (en) | 1985-05-29 | 1988-08-30 | Merck & Co., Inc. | Drug delivery device which can be retained in the stomach for a controlled period of time |
US4767861A (en) | 1986-01-28 | 1988-08-30 | Vipont Laboratories | Recovery of benzo-c-phenanthridine alkaloids |
US4772470A (en) | 1985-04-27 | 1988-09-20 | Nitto Electric Industrial Co., Ltd. | Oral bandage and oral preparations |
US4774227A (en) | 1986-02-14 | 1988-09-27 | Collagen Corporation | Collagen compositions for bone repair containing autogeneic marrow |
US4793336A (en) | 1981-03-25 | 1988-12-27 | Wang Paul Y | Wound coverings and processes for their preparation |
US4800219A (en) | 1986-12-22 | 1989-01-24 | E. I. Du Pont De Nemours And Company | Polylactide compositions |
US4804691A (en) | 1987-08-28 | 1989-02-14 | Richards Medical Company | Method for making a biodegradable adhesive for soft living tissue |
US4920203A (en) | 1987-12-17 | 1990-04-24 | Allied-Signal Inc. | Medical devices fabricated from homopolymers and copolymers having recurring carbonate units |
US4938763A (en) | 1988-10-03 | 1990-07-03 | Dunn Richard L | Biodegradable in-situ forming implants and methods of producing the same |
US4981696A (en) | 1986-12-22 | 1991-01-01 | E. I. Du Pont De Nemours And Company | Polylactide compositions |
US4983689A (en) | 1987-05-07 | 1991-01-08 | Yu Simon H | Process for making macromolecular monomers of polylactones with terminal acryloyl unsaturation and block copolymers thereof |
US5013553A (en) | 1987-06-30 | 1991-05-07 | Vipont Pharmaceutical, Inc. | Drug delivery devices |
US5077049A (en) | 1989-07-24 | 1991-12-31 | Vipont Pharmaceutical, Inc. | Biodegradable system for regenerating the periodontium |
EP0537559A1 (en) | 1991-10-15 | 1993-04-21 | Atrix Laboratories, Inc. | Polymeric compositions useful as controlled release implants |
EP0539751A1 (en) | 1991-10-28 | 1993-05-05 | Atrix Laboratories, Inc. | Biodegradable polymer composition |
EP0560014A1 (en) | 1992-03-12 | 1993-09-15 | Atrix Laboratories, Inc. | Biodegradable film dressing and method for its formation |
US5286763A (en) | 1983-03-22 | 1994-02-15 | Massachusetts Institute Of Technology | Bioerodible polymers for drug delivery in bone |
US5324520A (en) | 1988-12-19 | 1994-06-28 | Vipont Pharmaceutical, Inc. | Intragingival delivery systems for treatment of periodontal disease |
EP0649662A1 (en) | 1993-09-28 | 1995-04-26 | Atrix Laboratories, Inc. | Biodegradable implant precursor |
-
2000
- 2000-03-28 US US09/536,579 patent/USRE37950E1/en not_active Expired - Lifetime
Patent Citations (69)
Publication number | Priority date | Publication date | Assignee | Title |
---|---|---|---|---|
US2155658A (en) | 1936-01-08 | 1939-04-25 | Chemische Forschungs Gmbh | Surgical and medical preparations |
US3068188A (en) | 1957-03-27 | 1962-12-11 | Du Pont | Composition comprising a synthetic linear polymer, organic solvent, and an inorganic salt |
US3328246A (en) | 1957-08-20 | 1967-06-27 | Bicrex Lab Ltd | Dental compositions and methods of making same |
US3218283A (en) | 1962-07-26 | 1965-11-16 | Monsanto Co | Novel solutions of poly-(acrylic anhydride) and poly-(methacrylic anhydride) polymers |
US3219527A (en) | 1963-06-17 | 1965-11-23 | Loyola University | Peridontal pack or dressing composition |
US3463158A (en) | 1963-10-31 | 1969-08-26 | American Cyanamid Co | Polyglycolic acid prosthetic devices |
US3520949A (en) | 1966-07-26 | 1970-07-21 | Nat Patent Dev Corp | Hydrophilic polymers,articles and methods of making same |
US3887699A (en) | 1969-03-24 | 1975-06-03 | Seymour Yolles | Biodegradable polymeric article for dispensing drugs |
US3767784A (en) | 1970-12-01 | 1973-10-23 | S Gluck | Composition for the protection and treatment of injured body tissue and method of utilizing the same |
US3760034A (en) | 1971-01-26 | 1973-09-18 | Union Carbide Corp | Graft copolymers of lactone polyesters |
US3755558A (en) | 1971-02-23 | 1973-08-28 | Du Pont | Polylactide drug mixtures for topical application atelet aggregation |
FR2126270A1 (en) | 1971-02-23 | 1972-10-06 | Du Pont | |
US3696811A (en) | 1971-06-17 | 1972-10-10 | Squibb & Sons Inc | Periodontal bandage and backing therefor |
US4450150A (en) | 1973-05-17 | 1984-05-22 | Arthur D. Little, Inc. | Biodegradable, implantable drug delivery depots, and method for preparing and using the same |
US3939111A (en) | 1974-07-01 | 1976-02-17 | The B. F. Goodrich Company | Stable polyurethane solutions |
US3935308A (en) | 1974-08-08 | 1976-01-27 | The United States Of America As Represented By The Secretary Of The Navy | Wound covering and method of application |
US3931678A (en) | 1974-09-24 | 1976-01-13 | Loctite (Ireland) Limited | Dental filling method and composition formed thereby |
US4161948A (en) | 1977-01-18 | 1979-07-24 | Battelle Memorial Institute | Synthetic membrane for wound-dressings |
US4148871A (en) | 1977-10-11 | 1979-04-10 | Pitt Colin G | Sustained subdermal delivery ofdrugs using poly(ε-caprolactone) and its copolymers |
US4408023A (en) | 1980-11-12 | 1983-10-04 | Tyndale Plains-Hunter, Ltd. | Polyurethane diacrylate compositions useful for contact lenses and the like |
US4715369A (en) | 1980-12-29 | 1987-12-29 | Teijin Limited | Method of treating an injured part on the oral mucosa and the covering material for use thereof |
US4793336A (en) | 1981-03-25 | 1988-12-27 | Wang Paul Y | Wound coverings and processes for their preparation |
US4447562A (en) | 1981-07-15 | 1984-05-08 | Ivani Edward J | Amino-polysaccharides and copolymers thereof for contact lenses and ophthalmic compositions |
US4582640A (en) | 1982-03-08 | 1986-04-15 | Collagen Corporation | Injectable cross-linked collagen implant material |
US4570629A (en) | 1982-03-17 | 1986-02-18 | University Of Illinois Foundation | Hydrophilic biopolymeric copolyelectrolytes, and biodegradable wound dressing comprising same |
US4677139A (en) | 1982-04-07 | 1987-06-30 | Feinmann Bernhard P P | Material and method for dentistry |
US4491479A (en) | 1982-11-12 | 1985-01-01 | Lauchenauer Alfred E | Shaped semi-solid articles |
US4439420A (en) | 1982-11-16 | 1984-03-27 | Ethicon, Inc. | Absorbable hemostatic composition |
US4443430A (en) | 1982-11-16 | 1984-04-17 | Ethicon, Inc. | Synthetic absorbable hemostatic agent |
US5286763A (en) | 1983-03-22 | 1994-02-15 | Massachusetts Institute Of Technology | Bioerodible polymers for drug delivery in bone |
WO1985000969A1 (en) | 1983-08-31 | 1985-03-14 | S.K.Y. Polymers Ltd. | Injectable physiologically acceptable polymeric compositions |
US4631188A (en) | 1983-08-31 | 1986-12-23 | S.K.Y. Polymers, Ltd. (Kingston Technologies) | Injectable physiologically-acceptable polymeric composition |
EP0140766A2 (en) | 1983-10-07 | 1985-05-08 | The Forsyth Dental Infirmary For Children | Intra pocket drug delivery devices for treatment of periondontal diseases |
WO1985002092A1 (en) | 1983-11-14 | 1985-05-23 | Bio-Mimetics Inc. | Bioadhesive compositions and methods of treatment therewith |
EP0159293A1 (en) | 1984-04-02 | 1985-10-23 | Battelle Memorial Institute | Film-forming polymeric material for use in dressings, and its production |
US4663077A (en) | 1984-06-11 | 1987-05-05 | Morton Thiokol Inc. | Microbiocidal compositions comprising an aryl alkanol and a microbiocidal compound dissolved therein |
US4745160A (en) | 1984-06-26 | 1988-05-17 | Imperial Chemical Industries Plc | Biodegradable amphipathic copolymers |
CA1261549A (en) | 1984-07-16 | 1989-09-26 | Saeid Seyedin | Polypeptide cartilage-inducing factors found in bone |
EP0169016A2 (en) | 1984-07-16 | 1986-01-22 | Celtrix Pharmaceuticals, Inc. | Polypeptide cartilage-inducing factors found in bone |
US4595713A (en) | 1985-01-22 | 1986-06-17 | Hexcel Corporation | Medical putty for tissue augmentation |
US4650665A (en) | 1985-02-08 | 1987-03-17 | Ethicon, Inc. | Controlled release of pharmacologically active agents from an absorbable biologically compatible putty-like composition |
US4568536A (en) | 1985-02-08 | 1986-02-04 | Ethicon, Inc. | Controlled release of pharmacologically active agents from an absorbable biologically compatible putty-like composition |
US4772470A (en) | 1985-04-27 | 1988-09-20 | Nitto Electric Industrial Co., Ltd. | Oral bandage and oral preparations |
US4767627A (en) | 1985-05-29 | 1988-08-30 | Merck & Co., Inc. | Drug delivery device which can be retained in the stomach for a controlled period of time |
US4702917A (en) | 1985-11-18 | 1987-10-27 | Research Triangle Institute | Porous bioabsorbable polyesters |
US4767861A (en) | 1986-01-28 | 1988-08-30 | Vipont Laboratories | Recovery of benzo-c-phenanthridine alkaloids |
US4774227A (en) | 1986-02-14 | 1988-09-27 | Collagen Corporation | Collagen compositions for bone repair containing autogeneic marrow |
EP0241178A1 (en) | 1986-03-25 | 1987-10-14 | Rohto Pharmaceutical Co., Ltd. | Pharmaceutical composition for treating periodontal diseases |
US4780320A (en) | 1986-04-29 | 1988-10-25 | Pharmetrix Corp. | Controlled release drug delivery system for the periodontal pocket |
EP0244118A1 (en) | 1986-04-29 | 1987-11-04 | Pharmetrix Corporation | Controlled release drug delivery system for the peridontal pocket |
US4800219A (en) | 1986-12-22 | 1989-01-24 | E. I. Du Pont De Nemours And Company | Polylactide compositions |
US4981696A (en) | 1986-12-22 | 1991-01-01 | E. I. Du Pont De Nemours And Company | Polylactide compositions |
US4983689A (en) | 1987-05-07 | 1991-01-08 | Yu Simon H | Process for making macromolecular monomers of polylactones with terminal acryloyl unsaturation and block copolymers thereof |
US5013553A (en) | 1987-06-30 | 1991-05-07 | Vipont Pharmaceutical, Inc. | Drug delivery devices |
US4804691A (en) | 1987-08-28 | 1989-02-14 | Richards Medical Company | Method for making a biodegradable adhesive for soft living tissue |
US4920203A (en) | 1987-12-17 | 1990-04-24 | Allied-Signal Inc. | Medical devices fabricated from homopolymers and copolymers having recurring carbonate units |
US4938763A (en) | 1988-10-03 | 1990-07-03 | Dunn Richard L | Biodegradable in-situ forming implants and methods of producing the same |
US5278202A (en) | 1988-10-03 | 1994-01-11 | Atrix Laboratories, Inc. | Biodegradable in-situ forming implants and methods of producing the same |
US5278201A (en) | 1988-10-03 | 1994-01-11 | Atrix Laboratories, Inc. | Biodegradable in-situ forming implants and methods of producing the same |
US5340849A (en) | 1988-10-03 | 1994-08-23 | Atrix Laboratories, Inc. | Biodegradable in-situ forming implants and methods for producing the same |
US4938763B1 (en) | 1988-10-03 | 1995-07-04 | Atrix Lab Inc | Biodegradable in-situ forming implants and method of producing the same |
US5324520A (en) | 1988-12-19 | 1994-06-28 | Vipont Pharmaceutical, Inc. | Intragingival delivery systems for treatment of periodontal disease |
US5077049A (en) | 1989-07-24 | 1991-12-31 | Vipont Pharmaceutical, Inc. | Biodegradable system for regenerating the periodontium |
US5324519A (en) | 1989-07-24 | 1994-06-28 | Atrix Laboratories, Inc. | Biodegradable polymer composition |
US5368859A (en) | 1989-07-24 | 1994-11-29 | Atrix Laboratories, Inc. | Biodegradable system for regenerating the periodontium |
EP0537559A1 (en) | 1991-10-15 | 1993-04-21 | Atrix Laboratories, Inc. | Polymeric compositions useful as controlled release implants |
EP0539751A1 (en) | 1991-10-28 | 1993-05-05 | Atrix Laboratories, Inc. | Biodegradable polymer composition |
EP0560014A1 (en) | 1992-03-12 | 1993-09-15 | Atrix Laboratories, Inc. | Biodegradable film dressing and method for its formation |
EP0649662A1 (en) | 1993-09-28 | 1995-04-26 | Atrix Laboratories, Inc. | Biodegradable implant precursor |
Non-Patent Citations (6)
Title |
---|
"Encyclopedia of Polymer Science and Engineering", John Wiley & Sons, Inc., vol. 2, 236-237, (1985). |
Billmeyer, Textbook of Polymer Science, 3rd Ed., 390-391, (1984). |
Biocompatibility of Clinical Implant Materisl, (1976). |
Gilding, Biodegradable Polymers, 9, 210-232. |
Hawley's Condensed Chemical Dictionary (11th Ed.), Van Nostand Reinhold Co., NY, NY, 224, 555 and 567, (1987). |
Holland, "Polymers-for Biodegradable Medical Devices, 1, The Potential of Polyesters and Controlled Macromolecular Release Systems", J. of Controlled Release 4, 155-180, (1986). |
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---|---|---|---|---|
US7012126B2 (en) * | 1999-07-30 | 2006-03-14 | Ethicon, Inc. | Coumarin endcapped absorbable polymers |
US20020169275A1 (en) * | 1999-07-30 | 2002-11-14 | Takehisa Matsuda | Coumarin endcapped absorbable polymers |
US20020173616A1 (en) * | 1999-07-30 | 2002-11-21 | Takehisa Matsuda | Coumarin endcapped absorbable polymers |
US20020161170A1 (en) * | 1999-07-30 | 2002-10-31 | Takehisa Matsuda | Coumarin endcapped absorbable polymers |
US7144976B2 (en) | 1999-07-30 | 2006-12-05 | Ethicon, Inc. | Coumarin endcapped absorbable polymers |
US7105629B2 (en) | 1999-07-30 | 2006-09-12 | Takehisa Matsuda | Coumarin endcapped absorbable polymers |
US20030163192A1 (en) * | 2000-11-15 | 2003-08-28 | George Wallace | Methods for vascular reconstruction of diseased arteries |
US7011677B2 (en) * | 2000-11-15 | 2006-03-14 | Micro Therapeutics, Inc. | Methods for vascular reconstruction of diseased arteries |
US6565601B2 (en) * | 2000-11-15 | 2003-05-20 | Micro Therapeutics, Inc. | Methods for vascular reconstruction of diseased arteries |
US20040234615A1 (en) * | 2003-03-04 | 2004-11-25 | The Technology Development Company Ltd. | Oral insulin composition and methods of making and using thereof |
US20040234616A1 (en) * | 2003-03-04 | 2004-11-25 | The Technology Development Company Ltd. | Long acting injectable insulin composition and methods of making and using thereof |
US20100167985A1 (en) * | 2003-03-04 | 2010-07-01 | The Technology Development Company, Inc. | Oral insulin composition and methods of making and using thereof |
US20100166855A1 (en) * | 2003-03-04 | 2010-07-01 | Vladimir Sabetsky | Oral insulin composition and methods of making and using thereof |
US7544656B2 (en) | 2003-03-04 | 2009-06-09 | The Technology Development Company, Ltd. | Long acting injectable insulin composition and methods of making and using thereof |
US20060121085A1 (en) * | 2003-03-11 | 2006-06-08 | Warren Stephen L | Formulations for cell-schedule dependent anticancer agents |
US7976847B2 (en) | 2004-01-13 | 2011-07-12 | Vasogenix Pharmaceuticals, Inc. | Controlled release CGRP delivery composition for cardiovascular and renal indications |
US20080069865A1 (en) * | 2004-01-13 | 2008-03-20 | Southard Jeffrey L | Controlled Release Cgrp Delivery Composition for Cardiovascular and Renal Indications |
US20090023643A1 (en) * | 2004-01-13 | 2009-01-22 | Vasogenix Pharmaceuticals, Inc. | Methods For Treating Acute Myocardial Infarction By Administering Calcitonin Gene Related Peptide And Compositions Containing The Same |
US20090023644A1 (en) * | 2004-01-13 | 2009-01-22 | Southard Jeffery L | Methods of using cgrp for cardiovascular and renal indications |
US10292732B2 (en) | 2004-02-25 | 2019-05-21 | Femasys, Inc. | Methods and devices for conduit occlusion |
US8316853B2 (en) | 2004-02-25 | 2012-11-27 | Femasys Inc. | Method and devices for conduit occlusion |
US9402762B2 (en) | 2004-02-25 | 2016-08-02 | Femasys Inc. | Methods and devices for conduit occlusion |
US11779372B2 (en) | 2004-02-25 | 2023-10-10 | Femasys Inc. | Methods and devices for conduit occlusion |
US9308023B2 (en) | 2004-02-25 | 2016-04-12 | Femasys Inc. | Methods and devices for conduit occlusion |
US9238127B2 (en) | 2004-02-25 | 2016-01-19 | Femasys Inc. | Methods and devices for delivering to conduit |
US9220880B2 (en) | 2004-02-25 | 2015-12-29 | Femasys Inc. | Methods and devices for delivery of compositions to conduits |
US9839444B2 (en) | 2004-02-25 | 2017-12-12 | Femasys Inc. | Methods and devices for conduit occlusion |
US9034053B2 (en) | 2004-02-25 | 2015-05-19 | Femasys Inc. | Methods and devices for conduit occlusion |
US10111687B2 (en) | 2004-02-25 | 2018-10-30 | Femasys, Inc. | Methods and devices for conduit occlusion |
US8726906B2 (en) | 2004-02-25 | 2014-05-20 | Femasys Inc. | Methods and devices for conduit occlusion |
US8695606B2 (en) | 2004-02-25 | 2014-04-15 | Femasys Inc. | Methods and devices for conduit occlusion |
US8336552B2 (en) | 2004-02-25 | 2012-12-25 | Femasys Inc. | Methods and devices for conduit occlusion |
US8324193B2 (en) | 2004-02-25 | 2012-12-04 | Femasys Inc. | Methods and devices for delivery of compositions to conduits |
US8048101B2 (en) | 2004-02-25 | 2011-11-01 | Femasys Inc. | Methods and devices for conduit occlusion |
US8048086B2 (en) | 2004-02-25 | 2011-11-01 | Femasys Inc. | Methods and devices for conduit occlusion |
US8052669B2 (en) | 2004-02-25 | 2011-11-08 | Femasys Inc. | Methods and devices for delivery of compositions to conduits |
US8316854B2 (en) | 2004-02-25 | 2012-11-27 | Femasys Inc. | Methods and devices for conduit occlusion |
US7998202B2 (en) | 2004-09-16 | 2011-08-16 | Evera Medical, Inc. | Tissue implant having a biased layer and compliance that simulates tissue |
US7998201B2 (en) | 2004-09-16 | 2011-08-16 | Evera Medical, Inc. | Methods of forming a tissue implant having a tissue contacting layer held under compression |
US8313763B2 (en) | 2004-10-04 | 2012-11-20 | Tolmar Therapeutics, Inc. | Sustained delivery formulations of rapamycin compounds |
WO2007084460A2 (en) | 2006-01-18 | 2007-07-26 | Qps, Llc | Pharmaceutical compositions with enhanced stability |
US10646572B2 (en) | 2006-01-18 | 2020-05-12 | Foresee Pharmaceuticals Co., Ltd. | Pharmaceutical compositions with enhanced stability |
US20070196416A1 (en) * | 2006-01-18 | 2007-08-23 | Quest Pharmaceutical Services | Pharmaceutical compositions with enhanced stability |
US20080020016A1 (en) * | 2006-07-11 | 2008-01-24 | Quest Pharmaceutical Services, Llc | Pharmaceutical compositions for sustained release delivery of peptides |
US8840915B2 (en) | 2006-07-11 | 2014-09-23 | Foresec Pharmaceuticals, LLC | Pharmaceutical compositions for sustained release delivery of peptides |
US20090305957A1 (en) * | 2006-10-11 | 2009-12-10 | Qlt Usa, Inc. | Preparation of biodegradable polyesters with low-burst properties by supercritical fluid extraction |
US8076448B2 (en) | 2006-10-11 | 2011-12-13 | Tolmar Therapeutics, Inc. | Preparation of biodegradable polyesters with low-burst properties by supercritical fluid extraction |
US8324343B2 (en) | 2006-10-11 | 2012-12-04 | Tolmar Therapeutics, Inc. | Preparation of biodegradable polyesters with low-burst properties by supercritical fluid extraction |
US10206970B2 (en) | 2007-02-15 | 2019-02-19 | Tolmar Therapeutics, Inc. | Low-burst polymers and methods to produce polymer |
US11147880B2 (en) | 2007-02-15 | 2021-10-19 | Tolmar Therapeutics, Inc. | Low-burst polymers and methods to produce polymer |
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US9655970B2 (en) | 2007-02-15 | 2017-05-23 | Tolmar Therapeutics, Inc. | Low-burst polymers and methods to produce polymers |
US9974824B2 (en) | 2007-02-15 | 2018-05-22 | Tolmar Therapeutics, Inc. | Low-burst polymers and methods to produce polymer |
US20100292195A1 (en) * | 2007-02-15 | 2010-11-18 | Qlt Usa, Inc. | Low burst polymers and methods to produce polymer |
US11717572B2 (en) | 2007-02-15 | 2023-08-08 | Tolmar, Inc. | Low-burst polymers and methods to produce polymer |
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US9090737B2 (en) | 2007-11-13 | 2015-07-28 | Surmodics, Inc. | Viscous terpolymers as drug delivery platform |
US20110027389A1 (en) * | 2008-01-14 | 2011-02-03 | Dunn Richard L | Low Viscosity Liquid Polymeric Delivery System |
US8187640B2 (en) | 2008-01-14 | 2012-05-29 | Dunn Research & Consulting, Llc | Low viscosity liquid polymeric delivery system |
US20090198330A1 (en) * | 2008-02-01 | 2009-08-06 | Kesten Randy J | Compressible cosmetic prosthesis with improved shape retention |
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US9333070B2 (en) | 2008-02-01 | 2016-05-10 | Evera Medical, Inc. | Breast implant with internal flow dampening |
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US10149912B2 (en) | 2008-06-03 | 2018-12-11 | Indivior Uk Limited | Dehydrated hydrogel inclusion complex |
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US20090325879A1 (en) * | 2008-06-03 | 2009-12-31 | Qlt Usa, Inc. | Controlled release copolymer formulation with improved release kinetics |
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US11980395B2 (en) | 2008-10-03 | 2024-05-14 | Femasys Inc. | Methods and devices for sonographic imaging |
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US8974808B2 (en) | 2008-12-23 | 2015-03-10 | Surmodics, Inc. | Elastic implantable composites and implants comprising same |
US9415197B2 (en) | 2008-12-23 | 2016-08-16 | Surmodics, Inc. | Implantable suction cup composites and implants comprising same |
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US20100160892A1 (en) * | 2008-12-23 | 2010-06-24 | Tice Thomas R | Implantable suction cup composites and implants comprising same |
US11268049B2 (en) | 2010-05-20 | 2022-03-08 | Ecolab Usa Inc. | Rheology modified low foaming liquid antimicrobial compositions and methods of use thereof |
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