IL107393A - Compositions and methods for forming a biodegradable implant insitu - Google Patents

Description

ηΐλπ ητηα piDnnm Άχ^ηηη *j->nv ητχ'^ mu'i/i Π'ΤΒΟΙΙ COMPOSITIONS AND METHODS FOR FORMING A BIODEGRADABLE IMPLANT IN-SITU The present invention relates to a composition for forming a polymer which is curable in-situ within a body to produce a biodegradable implant, and to methods for forming a biodegradable implant in-situ in a body.

The present specification is divided from Israel Specification 91850, filed September 29, 1989.

In order that the invention may be better understood and appreciated, description from Israel Specification 91850 is included herein, it being understood that this is for background purposes only, the subject matter of Israel Specification.91850 being specifically disclaimed and not forming a part of the present invention.

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. Patent No. 3,297,033 to Schmitt; the poly( L-lactide-co-glycolide ) sutures described in U.S. Patent No. 3,636,956 to Schneider; the poly(L-lactide-co-glycolide) surgical clips and staples described in U.S. Patent No. 4,523,591 to Kaplan, et al., and the drug-delivery systems described' in U.S. Patent No. 3,773,919 to Boswell, et al. ; U.S. Patent No., 3,887,699 to Yolles; U.S. Patent No. 4,155,992 to Schmitt; U.S. Patent No. 4,379,138 to Pitt, et . al., and U.S. Patents No. 4,130,639 and No. 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. Patent No. 2,933,477 to Hostettler and U.S. Patent 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. Patents No. 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. Patent No. 4,379,138 to Pitt, et al. Trihydroxy-functional copolymers of ε-caprolactone and ε-valerolactone have been crosslinked with diisocyanates, thereby affording bviodegradable polymers, as described in Pitt, et al., J. Polym. Sci. : Part A: Polym. Chem. , Vol. 25, pp. 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, - λ - 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 as 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 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 peridontal 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 its 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.

In Israel Specification 91850, there is described and claimed a composition for forming a biodegradable implant in-situ within a body, comprising an effective amount of a biodegradable thermoplastic biocompatible polymer dissolved within a biocompatible solvent, which is capable of dissipating upon placement within a body to form said implant.

Said Israel Specification also claims a method of forming an implant in-situ in a body, comprising the steps of: a) dissolving a biodegradable thermoplastic polymer in a biocompatible solvent to form a liquid; b) placing said liquid within said body, and c) allowing said solvent to dissipate to produce a solid implant.

While said Israel Specification relates to a thermoplastic system in which a solid, linear-chain, biodegradable polymer or copolymer is dissolved in a solvent which is non-toxic and water miscible to form a liquid solution, the present invention is directed to a thermosetting system comprising the synthesis of cross-linkable 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.

More specifically, according to the present invention there is now provided a composition for forming a polymer which is curable in-situ within a body to produce a biodegradable implant, comprising a liquid acrylic ester terminated prepolymer capable of being cured into said implant upon addition of an effective amount of a curing agent.

The invention also provides a method of forming an implant in-situ in a body, comprising the steps of: . a) placing a liquid, biocompatible polymer within said body and b) curing said polymer in-situ to form said implant, wherein said liquid polymer is an acrylic ester terminated prepolymer and a curing agent is added to said prepolymer prior to placement of said prepoly er and allowing said prepolymer to cure in-situ.

The invention further provides a method of forming a solid implant in-situ within a body, comprising the steps of: a) mixing together effective amounts of liquid acrylic ester terminated, biodegradable prepolymer and a curing agent to form a mixture in a liquid form, and b) delivering said mixture within said body while said mixture is in a liquid form, so as to allow said prepolymer to cure to form said solid implant.

The multifunctional polymers useful in the thermosetting system are first synthesized via copolymerizatio of either DL-lactide or 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 azobisiso-butyronitrile, 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 solidifes.

For drug-delivery utilizing this system, biologically active agents are added to the liquid polymer systems in the uncatalyzed state.

In both the thermoplastic system of Israel Specification No. 91850 and the thermosetting system of the present invention, 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 which is capable of producing an effect on a body.

The invention will now be described in connection with certain preferred embodiments with reference to the following illustrative examples and figures so that it may be more fully understood.

With specific reference now to the examples and figures in detail, it is stressed that the particulars described and shown are by way of example and for purposes of illustrative discussion of the preferred embodiments of the present invention only, and are presented in the cause of providing what is believed to be the most useful and readily understood description of the principles and conceptual aspects of the invention. In this context, it is to be noted that only subject matter embraced in the scope of the claims appended hereto is intended to be included in the scope of the present invention, while subject matter of Israel Specification 91850, although described and exemplified to provide background and better understanding of the invention, is not intended for inclusion as part of the present invention.

In the drawings and tables: 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.

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 concentrations 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. Patent 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 CH2C12, filter out the Et3N"HCl, and then extract the CH2C12 layer with water. Stable emulsions were sometimes encountered during extraction. The acylations were later carried out in CH2C12 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 XH 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 XH 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 r 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 CH2C12. In some cases, fillers of 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 syringe and injected into a body. The mixture then 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 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 terms "drug" or "bioactive (biologically active) agent" as used herein include, 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, decongestants, miotics, anticholinergecs, sympathomimetics, sedatives, hypnotics, psychic energizers, tranquilizers, androgenic steroids, estrogens, progestational agents, humoral agents, prostaglandins, analgesics, antispasmodics, antimalarials, antihistamines, cardioactive agents, non-steroidal antiinflammatory 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 implant 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.

EXAMPLES The following examples are set forth as representat-ive of the present invention. These examples are not to- be construed as limiting the scope of the invention, as tfr 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.

EXAMPLE 2 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 homogenous solution of drug and polymer. Approximately 250 μΐ. 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.

EXAMPLE 3 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.

EXAMPLE 4 SaEt was added to the same polymer solution of DL-PLA in NMP as described in Example 3. A homogenous 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.

EXAMPLE 5 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 loadings 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 5 days.

EXAMPLE 6 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 or 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-.

EXAMPLE 7 SaEt was added to the same solution of DL-PLG in MP 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.

EXAMPLE 8 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 of 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.

EXAMPLE 9 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 ormulation 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.

EXAMPLE 10 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 as the catalyst. The polymer was dissolved in NMP to give a 50% by weight polymer solution. A quantity of the polymer solution (100 /i.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 μΐ,) 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.

EXAMPLE 11 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 g/mL after 3 days, 80.2 ^g/mL after 7 days, 67.8 μg/mL after 10 days, and 70.5 ^g/mL after 14 days.

EXAMPLE 12 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 eguiv/equiv OH) was added via a syringe. The addition funnel was charged with 15.4 g of acryloyl chloride (0.95 eguiv/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 CH2C12 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. Et3NlHCl was filtered out, additional CH2C12 (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 MgSO^/NajSO^, 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 l'REPOLYMERS SYNTHESIZED Mole ratio of monomers to initiator Hydroxyl No . , (ethylene glycol — 1.0) Cat lys ineq 011 (56.D/ (SnCl2), Theoretic l Sample no. DL- lactide (Ξ- caprolactone w % Mn , dal ons Observed Theore C96/»-l -l 2.4 5.0 0.03 993 100 11 C96A- 12 - 1. 6.1 32.8 0.05 5036 19.7 2 C96/+- 1.28-1 2.5 5.0 0.03 993 103 11 096 -136-1. 8.0 8.0 0.03 21.28 <Ί8 (est. ) 5 ΤΛΒΙ.Κ 2. SUMMARY OF ACRYLIC ES ER TERMINATED PREl'OLYMERS SYNTHESIZ Estimated Diol concentration Reaction conditions precursor acrylic groups, Sample no. sample no. meq/g Temp, "C Time, h Solvent C9f>/» -118-1 C96 -11 1.78 0 - RT 17 T1IF No pro 0964- 125-1 C96 - II -I 1.78 O-RT 17 THF Gelled.

Et3N-H C96 - 132-1 0964- 128-1 .8 T1IF 100 ppi workup .

C964- 137-1 C964- 128-1 I.8 0 Et70 Diffic C964-139-1 C96A-136-1 0.81 0 THF Gelled. residu refrig C96/i -IV. -1 0964 -136 0.81 0 CHz l2 No pro 096 - 1 6-1 C9M- A 0.33 0 No pro TABLE 3. SUMMARY OF CURING STUDIES Acryl ic Benzoyl Other Curing conditions Initial prcpol yiner peroxide acfcfi t ivcs, Shore A Sample no. sanplc no. wt X wt % Temp, °C Time, hardness Comncnt C964- 120-1 C964- 118-1 2.0 none 82 16 NDa Rubbery, breaks when bent 1 C964- 120-2 C9cV. -118- 1 1.0 none 82 16 83 Less brittle than C964-120- C964- 121-1 C964- 116-1 2.0 none 82 16 77 Rubbery, breaks when bent 1 C964-121-2 C9cV.-118- 1 1.0 none 82 16 80 Slightly stronger than C964 C964-121-3 C96 - 118-1 0.5 none 82 16 78 Slightly more elastic than 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 IMPTETA'346 82 2.5 94 Less rubbery than C964-120 C964- 122-2 C964-118-1 0.5 TMPTETA 46 82 2.5 91 Same as C964-122- 1 , more fl C964- 122-3 C96V'.- 118-1 1.0 TMPIETA 175 82 2.5 95 Hot rubbert at all, brittl C964- 122-4 C964-118-1 0.5 1ΜΡΙΕΤΛ 175 82 2.5 93 Simi tar to C964-122-3.

C964- 123-1 C964- 118-1 0.1 TMPTETA 46 82 2.5 89 Rubbery, stronger than C96 not flexible.

C964- 123-2 C964- 118-1 0.25 TMPTETA 46 82 2.5 83 About the seme as C964-123- C964- 123-3 C964- 118-1 0.1 TMPTETA 175 82 2.5 92 Hot rubbery; strong, brittl C964- 134-1 C964- 132-1 0.05 (AIBH)C none 60d 17 Liquid Mo cure.

C964-134-2 C964- 132-1 0.10 (AIBH) none 60d 17 Liquid Ho cure.

C964- 134-3 C964- 132-1 0.25 (AIBN) none 60d 17 L iquid Ho cure.

C964-134-4 C964- 132-1 0.50 (AIBH) none 60d 17 L iquid Ho cure.

C964-134-5 C964- 132-1 1.00 (AIBN) none 60d 17 I iquid SI ightly thickened.

C964- 135-1 C964-132- 1 0.05 none 80d 17 Liquid Ho cure.

C964- 135-2 C964-132-1 0.10 none 80d 17 Liquid Ho cure.

C964- 135-3 C964- 132-1 0.25 none 80d 1 Liquid Ho cure.

C964- 135-4 C964- 132-1 0.50 none 80d 17 I iquid No cure.

G964- 135-5 C964- 132-1 1.00 none 80d 17 Liquid Slightly thickened.

C964-135-6 C96 -128-1c 0.05 none 80d 1 L iquid No cure. (continued) TABLE 3 (continued) Acryl ic Benzoyl Other Cur ing condi t ions Initial prepolytner peroxide additives, Shore A Sample no. snn le no. wt V. wt % Temp, "C Time, h hardness Comnents C964-135-7 C964-128-1e 0.10 none 80d 17 L iquid Ho cure.

C964- 135-8 C964- 128-1° 0.25 none 80d 17 L iquid Ho cure.

C964- 135-9 C964-128-1e 0.50 none 80d 17 L iquid No cure.

C964- 135-10 C964-128-1e 1.00 none B0d 17 L iquid Ho cure.

C964-135-11 C9cV« - 12 - 1 e 0.05 none 80d 17 WD Ho cure.

C964-135-12 C964-124-1e 0.10 none 80d 17 NO No cure.

C964- 135- 13 C964-124-1e 0.25 none 80d 17 HO Ho cure.

C964-135- 14 C964-124-1e 0.50 none 80d 17 HO Ho cure.

C96¾- 135- 15 C964-124-1e 1.00 none 80d 17 M0 No cure.

C964- 41-1 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 Ho cure.

C964-141-6 C964- 128-1 0.25 none 80 1 Liquid Ho cure.

C964- 141-7 0964- 128-1 0.50 none 80 1 L iquid No cure.

C964-141-8 C964- 128-1 1.00 none 80 1 L iquid Ho cure.

C964- 143- 1 C964- 137-1 0.25 Cnb-o-Si l ΡΤΠ, 80 1 74 No cure. ς . n u C964- 143-2 C964- 37-1 0.25 Cab-o-Sil PIG, 80 1 73 No cure.

C964- 143-3 C964- 137-1 0.25 L-PLA (1V=0.8), 80 1 75 Ho cure. c.. nu C964- 143-4 C964- 137-1 0.25 L-PLA (IV=0.8), 80 1 78 Ho cure.

J C964- 148-1 C964- 144-1 0.05 none 80 17 L iquid No cure.

C964- 148-2 C964-144-1 0.10 none 80 17 L iquid Ho cure.

C964- 148-3 C964- 144-1 0.25 none 80 2 66 C964-1 8- 4 anj C964-148-6 w toughness, and both were bet C964- 148-5. (continued) TABLE 3 (continued) Acrylic Benzoyl Other Curing conditions Initial prepolymer peroxide additives, Shore A Sample no. sanple no. wt 7. ut V. Temp, °C Time, h hardness Coninents C964- 148-4 C964- 144-1 0.50 noiie 80 2 68 C964-148-4 and C964-148-6 we toughness, and both were bet C964-148-5.

C964-148-5 C964- 144-1 1.00 none 80 2 67 C964-148-4 and C964- 148-6 we toughness, and both were bet C964-148-5.

C964- 148-6 C964-144-1 2.00 none 80 2 69 C964- 148-4 and C964- 148-6 we toughness, and both were bet C-964- 48-5.

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 C9 4-144-1 0.15 Cab-o-Si I N70- 1S 80 2 HO Seniles too porous, did not 5.0 for hardness measurement.

C964- 149-5 C964- 144-1 0.20 Cab-o- Si I N70- IS 80 2 no Samples too porous, did not 5.0 for hardness measurement.

C964- 149-6 C964-144-1 0.25 Cab-o-Si I N70- IS 80 2 NO Samples too porous, did not 5.0 for hardness measurement.

C964- 150-1 C964-K6-1 0.05 none 80 17 WD Only partially cured.

C964- 150-2 C964- 146-1 0.10 none 80 2 72 Elastic, flexible, moderatel C964-150-3 C964- 146-1 0.25 none 80 2 57 Elastic, flexible, moderatel C964-150-4 C964- 146-1 0.50 none 80 2 56 Elastic, flexible, moderatel C964- 150-5 C964- 146-1 1.00 none 80 2 50 Elastic, flexible, moderatel C964- 150-6 C964- 146-1 2.00 none 80 2 51 Elastic, flexible, moderatel "Result not determined.

^IMPTEIA = trimethylolpropane triethoxy triacrylate.

CAIBM = azobisisobutyronitrile. ¾ured in air at atmospheric pressure. eDiol prepolymer used. -24D - It will be evident to those skilled in the art that the invention is not limited to the details of the foregoing illustrative examples and that the present invention may be embodied in other specific forms without departing from the essential attributes thereof, and it is therefore desired that the present embodiments and examples be considered in all respects as illustrative and not restrictive, reference being made to the appended claims, rather than to the foregoing description, and all changes which come within the meaning and range of equivalency of the claims are therefore intended to be embraced therein.

Claims (27)

- 25 - WHAT IS CLAIMED IS:

1. A composition for forming a polymer which is curable in-situ within a body to produce a biodegradable implant, comprising a liquid acrylic-ester-terminated prepolymer capable of being cured into said implant upon addition of an effective amount of a curing agent.

2. The composition of claim 1, wherein said liquid acrylic ester terminated prepolymer is a product of a conversion of a polyol-terminated prepolymer.

3. The composition of claim 2, wherein said polyol-terminated prepolymer is a product of co-polymerization of DL-lactide and ε-caprolactone with a polyol initiator.

4. The composition off claim 2, wherein said polyol-terminated prepolymer is a product of co-polymerization of L-lactide and ε-caprolactone with a polyol initiator.

5. The composition of claim 1, wherein said curing agent is azobisisbutyronitrile.

6. The composition of claim 1, wherein said curing agent is benzoyl peroxide. - 26 - 107,393/2

7. The composition of claim 1, and further comprising an effective amount of a biologically active agent.

8. A method of forming an implant in-situ in a body of a non-human animal, comprising the steps of: a) placing a liquid, biocompatible polymer within said body; and b) curing said polyment in situ to form said implant, wherein said liquid polymer is an acrylic-ester-terminated prepolymer and a curing agent is added to said prepolymer prior to placement of said prepolymer and allowing said prepolymer to cure in-situ.

9. The method of claim 8, and further comprising the step of synthesizing said prepolymer via copolymerization of DL-lactide with ε-caprolactone.

10. The method of claim 8, and further comprising the step of synthesizing said prepolymer via copolymerization of L-lactide with ε-caprolactone.

11. A method of forming a solid implant in-situ within a body of a non-human animal, comprising the steps of: a) mixing together effective amounts of liquid acrylic-ester-terminated, biodegradable prepolymer and a curing agent to form a mixture in a liquid form; and b) delivering said mixture within said body while said mixture is in a liquid form, so as to allow said prepolymer to cure to form said solid implant. - 27 -

12. The method of claim 11, and further comprising the step of forming said liquid acrylic-ester-terminated prepolymer by converting a polyol-terminated prepolymer.

13. The method of claim 12, and further comprising the step of forming said polyol-terminated prepolymer by copolymerization of DL-lactide and ε-caprolactone with a polyol initiator.

14. The method of claim 13, and further comprising the step of adding a catalyst to said copolymerization step.

15. The method of claim 14, wherein said catalyst is stannous octoate.

16. The method of claim 14, wherein said catalyst is stannous chloride. ί

17. The method of claim 12, and further comprising the step of forming said polyol-terminated prepolymer by copolymerization of L-lactide and ε-caprolactone with a polyol initiator.

18. The method of claim 17, and further comprising the step of adding a catalyst to said copolymerization step. - 28 -

19. The method of claim 18, wherein said catalyst is stannous octoate.

20. The method of claim 18, wherein said catalyst is stannous chloride.

21. The method of claim 11, wherein said curing agent is azobisisobutyronitrile .

22. The method of claim 11, wherein said curing agent is benzoyl peroxide.

23. The method of claim 11, and further comprising the step of adding a biologically active agent to said prepolymer and curing agent mixture to provide, upon curing, a biodegradable implant which releases said biologically active agent by diffusion or erosion as said implant biodegrades.

24. The method of claim 11, wherein said delivering step comprises injecting said mixture into said body by means of a syringe and needle.

25. A biodegradable implant for a body, produced according to the method of claim 8. - 29 - 107,393/2

26. A biodegradable implant for a body, produced according to the method of claim 11.

27. A biodegradable implant for a body, produced from a composition according to claim 1. for the Applicant: GOLLER