Plasticizable implant material and method for producing the same

A plasticizable implant material for injection into the human body. The implant material is prepared through simultaneous selection of the comonomer ratio of .epsilon.-caprolactone and lactide and production parameters to provide a (co)polymer having an average molecular weight and caprolactone concentration that produces a phase transition temperature from solid to fluid within a temperature range of 37-55.degree. C., preferably within the range of 43-48.degree. C. When the temperature of the implant material is regulated to be within said range it can easily be transported to and placed at its target in the body by using a fluid transport and dosing device like a syringe or a heated press.

BACKGROUND OF THE INVENTION
 During the last years in the biomedical applications synthetic implant
 materials have been taken into use to an ever-increasing degree.
 Biomaterial is defined to be a synthetic structural material whose aim is
 to interact with the biological system, and to replace, to treat, promote
 healing and renewal of and to join tissue, organs or some function of the
 body. Applications of these materials are reviewed in a publication edited
 by Hocker et al. (Macromolecular Symposia, vol. 103, January 1996, Klee,
 D., Severich, B., Hocker, H., pp. 19-29). Among the most important present
 and future applications are fixation materials of different types for bone
 fracture treatment which can be used for manufacturing screws, nails or
 rods for the above mentioned application, just to mention an example.
 These materials can be either non-biodegradable ones, e.g., metals or
 metal alloys, or polymeric materials degradable at a controlled rate in
 the body.
 The most widely used biodegradable materials are high molecular weight
 lactide homopolymers, and lactide copolymers with, for example, glycolide.
 Useful parts or products are processed from these materials with
 processing methods for thermoplastics known in polymer technology, such as
 injection molding, hot pressing or extrusion.
 Dentistry is well familiar with polymeric materials, too. Typical polymeric
 dental filling materials are chemically (for example, photochemically)
 curable plastics based on methyl methacrylate, dimethyl acrylate and their
 derivatives.
 Fixation bone cements for orthopedic hip prostheses are also based on
 monomer combinations methacrylate type. In these applications the curing
 is based on redox initiated free radical polymerization, and on thus
 accomplished cross-linking and network formation.
 The methacrylate based implant materials are, however, neither
 biodegradable nor biocompatible to any particular extent, as Dr. Heikkila
 reports in his dissertation (Annales Universitatis Turkuensis Ser. D:
 Medica-Odontologica, tom. 240, 23.8. 1996, Turku/Finland, Heikkila, J.,
 Bioactive glass as a bone substitute in experimental and clinical bone
 defects, pp. 1-97, especially on p. 30).
 In the use of methacrylate based implant materials further problems are
 caused by the exposition of personnel to volatile compounds, and the heat
 released during the reaction which may lead to an excessive local
 temperature increase, and to tissue damages as a consequence.
 Another application for synthetic implant materials is controlled release
 of drugs or other bioactive substances when the idea is that the potent
 agent is released at a controlled rate from the polymeric matrix. As an
 example of this kind of application one can mention Norplant, a product
 brand and a trade mark of Leiras Co., which is based on a non-degradable
 polymeric material. A definitely formed device is implanted into the body
 by a surgical operation, and it is removed therefrom in a similar manner
 after a defined time when the active component has been released and
 diffused to the body.
 Specific needs for development in the present state of the art are
 connected to the following areas:
 Biocompatibility
 If the material is not biocompatible it may induce tissue inflammation,
 unwanted cell growth, or rejection. Biocompatibilities of the presently
 used bone cements based for the most part on poly(methyl methacrylate) are
 unsatisfactory. This causes a certain risk of loosening of the hip
 prosthesis even in the case that there exists connective tissue formation
 between the polymeric material and bone tissue. A better biocompatibility
 would be a significant benefit for these materials.
 Bioactivity
 A bioactive implant material makes possible an active interaction between
 the tissue and the implant. As an example can be taken a mechanism by
 which the tissue is enabled to reconstruct into the implanted material
 while the implanted material itself is gradually removed due to
 biodegradation. Heimke, G. and Griss, P. have in their publication (Tissue
 interactions to bone replacement materials, in Bioceramics of calcium
 phosphate, de Groot, K. (ed.), 1983, CRC Press, Boca Raton Fla., pp.
 79-97) characterized the concept of bioactivity, and have been cited by
 Heikkila in his publication (FIG. 1 in the publication cited on page 1
 where a) bioincompatible materials, b) bioinert but by the interface
 biocompatible materials, and c) bioactive and biocompatible materials are
 presented schematically. In the case a) the implant is tolerated but no
 connection with bone is formed, in the case b) intimate contact without
 bone bonding occurs at the interface whereas in the case c) both intimate
 contact with chemical bone bonding and gradual transformation between bone
 and implant material will result).
 Bioactive materials have scarcely been reported in the literature.
 Especially in the case of bone cements bioactivity would be desirable and
 a significant benefit.
 Controlled Biodegradation
 Depending on the application and purpose of implant materials, they are
 expected to have either long lasting durability or controlled
 degradability in the body at a predetermined rate to harmless degradation
 products. The wanted degradation rate is depending substantially on the
 renewal rate of the tissue. In the case of bone tissue, it may be case of
 several months, or even of a time span in the range of half an year to one
 year.
 In the case of controlled drug delivery it is crucial what is the desired
 rate of release of the active ingredient from the biodegradable matrix.
 When the potent ingredient release is based on matrix degradation the rate
 of matrix degradation determines the release rate of the drug. When active
 agent is released from the matrix through diffusion, degradation of the
 matrix shall happen mainly only after the release of the active agent.
 Industrial Hygienic Aspects
 The materials in continuous clinical use have to be safe to the users in a
 sense of work safety and hygiene. This is a severe drawback with the
 present bone cements and dental filling materials which are based on
 methacrylates.
 Controlled Mechanical Properties
 The mechanical properties required from implant materials are depending on
 the application. With bone implants usually a compression strength of at
 least 50 Mpa is necessary, as well as bending strength and tensile
 strength values which are at the level of those of bone. On the other
 hand, even in the bone applications , in case of bone grafting by filling
 of fractures and cavities, one can fairly well apply implant materials of
 lower strength if only the use properties, mouldability, biocompatibility,
 and possible biodegradability are at an optimum level.
 In connection with soft tissue the requirements, on the other hand, are
 elasticity, flexibility and softness.
 Plasticizability and Hardening Thereafter
 The today used polymeric implant materials are either pieces of definite
 shape, i.e., processed before implanting to the final form using methods
 known in the plastics technology (as an example one can mention
 biodegradable bone nails based on polylactide, e.g., trade name Biofix),
 or bone cements based on methacrylate which typically have no
 biodergradability and lack bioactivity but as monomers, or as a blend of
 monomers, can be shaped in the target according to the needs, and can be
 hardened thereafter.
 In surgery there would be plenty of applications for plasticizable, and
 afterwards to solid curable biodegradable polymeric materials. Then the
 idea is, that the material is plastic in connection to the surgical
 operation , and can be shaped according to the target's shapes or can be
 forced to penetrate even into small cavities, fractures and pores.
 Thereafter it again reversibly becomes solid, mechanically tough material
 which, however, has the property of controlled degradation. Thus
 plasticizable material can then be of the type of wax, plastic or rubber.
 Better biocompatibility, bioactivity and wished mechanical properties as
 combined to the mouldability in the target and hardening occurring
 thereafter are properties which before the present invention have not been
 known in the state of the art.
 Applicability to the Matrix of Bioactive Components
 As examples of a bioactive active agent in an implant material can be
 mentioned different drugs, hormons or components activating tissue growth,
 such as hydroxyapatite in connection with bone tissue, and certain
 proteins.
 The matrix has to be such a material that the appropriate hormons can be
 easily blended with it to form a homogeneous mixture. Especially the upper
 temperature in this process is often limited because many drugs and
 hormons are heat sensitive.
 Easiness in Application (Usability and Transferability to the Target)
 The implant material is placed to the target in connection to a clinical
 situation, e.g., in connection to an operation. Then the applicability and
 the mouldability of the material has to be easy: it must be possible to,
 for example, inject it, or to place it with a special press, to the
 target, and its hardening has to have a certain induction period during
 which the material can be shaped. On the other hand, one has to take into
 account that the possible drug present and/or the contact with tissue do
 not allow use of methods where the temperature even for a short period
 exceeds the typical upper limit of 55.degree. C. The invention now present
 brings a novel solution which significantly improves applicability of
 biodegradable implants, for example, in regenerating surgery and in long
 lasting drug theraphy.
 SUMMARY OF THE INVENTION
 In comparison with the present state of the art, it has been possible
 according to the present invention surprisingly to achieve such a method
 which brings a significant improvement to it and opens up new
 possibilities in the use of body implantable polymeric materials. More
 specificly it has been possible to find such a composition of materials
 and a method, which enables the combination of several of the important
 properties in an exceptionally beneficial way from many applications'
 point of view. Below the invention's characteristic features are explained
 more in detail through examples which describe one successful way of its
 implementation.
 The developed material is based on as such known structural units used in
 syntheses of polymeric biomaterials, and on structural units derived from
 capronic acid. Structural units derived from lactic acid can be, in
 addition to lactic acid itself, L- , D-, and DL-lactide units. Structural
 units derived from capronic acid can be, e.g., .epsilon.-caprolactone. In
 addition as structural units can be organic carbonates, like
 trimethylenecarbonate.
 Lactones are cyclic esters based on hydroxyacids. Some of the most common
 lactones are L-lactide, DL-lactide, D-lactide and E-caprolactone. With
 these to the same group of cyclic monomers can also be included cyclic
 carbonates, like trimethylenecarbonate.
 The polymerization of lactones and cyclic carbonates can be carried out, as
 is well known. through catalytic ring opening polymerization. The catalyst
 used is typically some organometallic compound like tin(II)octoate (in
 other words stannous-2-ethylhexenoate), or trimethylaluminium.
 The molecular weight control in this type of polymerization is based on the
 optimal selection of polymerization temperature and time, and it is
 possible also using so called initiator compounds, of which typical are
 multifunctional alcohols, e.g., glycerol. During the polymerization the
 polymer chains will start to grow from the --OH groups so that the
 molecular weight will be the lower the more initiator is present. By
 choosing the structure of the multifunctional alcohol the shape of the
 forming molecule can be affected. So, for example, glycerol forms a
 comb-like and pentaerythritol a star-shaped molecular structure,
 respectively.
 The ring-opening polymerization of lactones is described, for example, in
 the following publication: Ylikangas, I., The polymerization of
 .epsilon.-caprolactone with stannous catalysts, Polyemer Technology
 Publication Series No. 15, Helsinki University of Technology, 1993, 1-23.
 Thus it is as such known that by using the above mentioned structural units
 one can form a biodegradable polymeric material which has either
 plastic-like or rubbery properties but in the present invention the
 structural units are used in a special way to achieve a biodegradable
 polymeric material which melts within a certain narrow temperature range
 and which, on the other hand, is either a mechanically tough plastic, a
 wax or a rubbery material below its melting temperature. In respect of
 implant application it is important that the material melts or plasticizes
 at a temperature which is somewhere 3-12.degree. C. above the body
 temperature, and becomes again a solid material at a temperature still
 above the body temperature. The melting and plasticization temperature of
 the material must, of course, not to be too high in respect to the body
 temperature, taking into account the above described temperature
 resistance of tissue and of drug-like active agents. When melting or
 plasticizing the material according to the invention becomes a viscose
 mass or fluid which can be shaped, injected or otherwise used to fill
 hollows, fractures and cavities as well as to replace different kinds of
 tissue defects, as in bone grafting.
 In the present invention it has become possible to adjust the temperature
 at which the material becomes plastic to within the temperature range of
 42-55.degree. C., or it can be adjusted to any temperature around this
 temperature range as well. The melting temperature of the material can be
 checked by means of measurements of entalphy changes, using differential
 scanning calorimetry (DSC) as the method.
 The control of the melting temperature of a polymeric material according to
 the invention is based on one hand on the specific monomer ratio selection
 in the starting materials, and, on the other hand, on the specific control
 of the molecular weight in the copolymerization.
 Both factors together namely affect the melting temperature of the
 copolymer, hence only certain combinations of them bring about the wished
 result. FIG. 1 is a graphical presentation the combinations of monomer
 compositions and average molecular weights for the L-lactide/caprolactone
 copolymers which produce a suitable melting temperature for implant use,
 indicated by points falling within the shaded area between curves in the
 monomer molar ratio versus molecular weight coordination.
 In several applications it is wanted that the implanted material is
 degradable in a controlled manner, or vice versa mechanically stable at
 least for a certain period of time. The first stage in bidegradation of
 the materials according to the present invention is hydrolysis which
 splits polymer chains to shorter segments until the molecular size is at a
 level where the own enzymatic functions of the body can convert the
 degradation products to compounds which are natural in the body.
 The hydrophilic character of the polymer is an essential parameter
 affecting the degradation rate. Thus in copolymers of the present
 invention it is possible to control hydrolytic degradation rate through
 control of monomer composition, and so also hydrophilic character, which
 according to what has been presented above directly affects the
 degradation rate of the material in the body. One has to note, however,
 that the exact rate of the degradation in the body depends on the end use,
 and has to be always investigated case by case. FIG. 2 indicates the
 dependence between the hydrolysis rate and monomer composition in the
 L-lactide/.epsilon.-caprolactone copolymers prepared in the examples.
 An essential feature of the invention is that if the composition of the
 material comprises only or for the most part .epsilon.-caprolactone, the
 rest being L-lactide, DL-lactide, D-lactide or trimethyl carbonate, the
 polymer is almost stable in the body, or degrades very slowly, typically
 during several years. Through selection of the monomer composition keeping
 this in mind, and through adjusting the molecular weight of the forming
 polymer by controlling the polymerization parameters, one can get benefit
 from the well-known biocompatibility and bioactivity of poly(hydroxy
 acid)s. On the other hand, a waxy version of the copolymers according to
 the invention can be made to degrade even really fast, just by controlling
 both average molecular weight and monomer composition as described above.
 In that case the degradation rate in the body is typically from some days
 to some weeks.
 A material according to the invention is in a molten state, i.e., while
 being at a temperature exceeding 37.degree. C., preferably within a
 temperature range from 42 to 55.degree. C., and most preferably within a
 temperature range from 43 to 48.degree. C., a plastic mass or a viscous
 fluid. This makes it possible to shape the material by hand or by using
 specific tools to a desired shape to meet the needs of the target, or
 alternatively makes it possible that the material can, by different
 transportation methods, be used to fill different and variably shaped
 spaces, hollows, cavities and fractures numerously existing in surgery,
 dentistry and medicine.
 Especially advantegous in the invention is an unexpected observation that
 the plasticizable material according to the invention can above its
 melting temperature be in such a state that it can be applied to the
 target using a specific press, the schematic drawing of which is presented
 in FIG. 3. Said press essentially consists of a heated cylinder 3, a die 4
 (which can be of different kinds according to the needs of the end uses),
 a piston 5, trigger 2, and a press mechanism 8. The temperature of the
 cylinder 3 can be regulated at a desired level by a thermostat so that the
 implantable material is either in a molten state or in a plasticized state
 inside the cylinder in front of the piston 5, and thus it can be
 transferred using the press device through the die to the target area.
 Because the temperature of the molten or plasticized polymeric material is
 higher than the body temperature but in any case such that there is no
 danger of tissue damage, the implant material can be transferred directly
 to the target area where it solidifies when cooling down.
 Advantages of the method according to the invention are the simplicity for
 the user, e.g., for the surgeon, hygienicity because the implant material
 is all the time in a closed chamber, and the point that there are no
 volatile components utilized in the method (an advantage in respect of
 industrial hygiene and working safety), and further that there is no heat
 generation in connection with the solidification (as usually is the case
 in the solidification of bone cements) and the risk of tissue damages can
 thus be avoided. As an additional benefit one can mention that by using
 differently shaped longer dies 2 it becomes possible to inject the mass to
 narrow, deep and variously shaped channels to their very end or to their
 deepest points, to fractures, to cavities and the like as target areas.
 Another preferred embodiment of the implantation method according to the
 invention is the use of such a device which does not contain any specific
 piston, but the piston 5 is replaced with a rod made up of polymeric
 material according to the invention which penetrates through the unheated
 part of the cylinder to the heated part where it melts or plasticizes. The
 feeding mechanism 8 pushes the rod into the cylinder, and correspondingly
 the molten or plasticized polymeric material is extruded out through the
 die 4.
 Still one advantageous method to implantate according to the invention is
 the one where the temperature of the cylinder 3 is raised above one
 hundred degrees centigrade and is held there for the time required for
 completing heat sterilization. In this case any separate sterilization is
 not needed, and the sterilized material remains in its closed space
 sterile and ready for use. Naturally this form of implantation can be
 applied provided that there are no heat sensitive active agents, for
 example drugs, present as blend components. After the sterilization
 treatment the material is allowed to cool down to the temperature range of
 37-55.degree. C. so that it remains both sterile and in a molten or
 plasticized state for the implanting.
 One specific way of implantation is the one where the polymer according to
 the invention is produced in a way that its melt viscosity is low enough
 to allow its injection to the target area with a suitable injection
 syringe. In this case the polymeric material when cooled down to its solid
 state at the body temperature is a waxy material.
 The implant material according to the present invention can be applied in a
 way that the material forms a matrix in the controlled release of drugs,
 or it forms a matrix in blends containing bioactive inorganic materials
 where the bioactive component can be selected from the group including,
 for example, hydroxyapatite, a filler based on coral, bone allocrafts or
 its particles, titanium particles and carbon fibers, just to mention some.
 In this form of application the blend component can, of course, also be
 soluble to the polymeric matrix, as is the case in connection with the
 most drugs. On the other hand, the solid components in the mixtures may be
 platelets or fibrous particles so that the composition is comparable to
 the polymer composites, and in which case the blend component essentially
 improves mechanical properties. A special advantage is achieved in the
 aspect that the use of a biodegradable polymeric matrix yields adhesion
 between particles, plasticizability, ease of processing, and improved
 biocompatibility to this kind of materials containing solid particles.
 In connection with drugs, hormones, or other corresponding active agents,
 the method according to the invention enables the controlled release and
 targeted location to the body, for example, in connection with an
 operation, or by injecting.
 Both composite materials containing solid fillers and/or reinforcements and
 blends containing soluble components can be located according to the
 invention, for example, as is presented in FIG. 3, easily and safely as a
 plastic mass or a viscous fluid directly to the end use target.
 Furthermore it is characteristic for the method according to the invention
 that also other methods suitable for the dosing and transferring plastic
 masses and viscous fluids are suitable for use, and they are in accordance
 with the invention.

DESCRIPTION OF PREFERRED EMBODIMENTS
 EXAMPLES 1-14
 The Used Chemicals
 The copolymers were prepared from .epsilon.-caprolactone monomer
 (.epsilon.-CL), &gt;99% purity, Fluka Chemika nr. 21510, lot 335334/1 794,
 and D,L-lactide (D,L-LA), Purac, lot DF 386H. As catalyst was used
 tin(II)octoate (Stannous 2-ethylhexanoate; SnOct), 95% purity , Sigma nr.
 s-3252, lot 112H0248. As the initiator was used glycerol, 99,5% purity,
 Fluka BioChemika nr. 49767, lot 42489/2 1094.
 The Purification and Storage of the Used Chemicals
 In the used .epsilon.-caprolactone there was molecular sieves (adding date
 15.02.1995), and the bottle was stored in a dark place at a temperature of
 23.degree. C. The caprolactone was not distilled again.
 D,L-lactide was purified with recrystallication from toluene (b.p.
 110.degree. C.) using a mass ratio of 1:2 toluene/lactide. The lactide
 dissolved to a hot toluene was poured from round bottom flask to a
 decanter. The lactide solubilized to the toluene was let to recrystallize
 overnight at 23.degree. C. After filtration the crystallized lactide was
 dried under reduced pressure for 4 days at +40.degree. C. and 4 mbar. The
 same stages were repeated once. In the polymerizations was thus used twice
 recrystallized D,L-lactide which was stored in an exsiccator in a
 refrigerator at +4.degree. C. The tin octoate and the glycerol were used
 as such.
 They were stored in a dark place at +23.degree. C.
 Preparations for Polymerization
 At preceding night the used lactide has been placed into a vacuum chamber
 at +40.degree. C. and 4 mbar. The two-piece polymerization reactor (volume
 about 0.7 liter) was assembled, and the condition of the Teflon gasket
 belonging to the reactor was checked. The proper closure of the upper part
 and the lower part of the reactor was ensured by a iron wire closing
 device. The male parts of the glass joints belonging to the reactor were
 wiped slightly with a vacuum grease.
 Polymerization
 The oil thermostat used for the reactor heating was regulated to
 140.degree. C. The oil temperature varies during a polymerization within
 5.degree. C. above and below the set temperature. Lactide was weighed
 first about 10 g into a small decanter (accuracy 0.0001 g). On the lactide
 the tin octoate and the glycerol was weighed using a Pasteur pipet. After
 this the decanter was poured into the reactor, and the rest of the lactide
 was weighed with another balance (accuracy 0.01 g). .epsilon.-Caprolactone
 was then either poured or pipeted on lactide.
 The magnetic agitator has been added to the reactor before the reactor
 halves were closed. The reactor was placed into a thermostat, and the
 agitation was adjusted to the speed of 250 1/min. The reactor was flushed
 with Argon (AGA, grade S, 99.99%) for about 15 min. Argon was fed to the
 reactor through a glycerol trap. Finally the outside of the reactor was
 wrapped with a aluminium foil. When the forming copolymer started to
 become more viscous the agitation speed was adjusted again to the speed of
 125 l/min.
 The Prepared Copolymers and their Analysis
 Table 1 summarizes the copolymerizations and their results using
 .epsilon.-caprolactone and D,L-lactide (.epsilon.-CL/D,L-LA). In all the
 polymerizations the temperature was 140.degree. C. and the polymerization
 time was 24 h (except in Example nr. 3 where it was 29 h). Molecular
 weights determined by gel permeation chromatography (GPC) are presented in
 Table 1 in terms of number average molecular weight Mn, weight average
 molecular weight Mw, and the polydispersity PD calculated as the ratio of
 the previous ones Mw/Mn. In the same Table 1 there are also presented the
 transition temperatures of the polymerization products. i.e., melting
 temperature Tm and glass transition temperature Tg, determined using
 differential scanning calorimetry (DSC).
 GPC Measurements
 The GPC-samples for molecular weight measurements were prepared by
 dissolution of 15 mg of sample into 10 ml of chloroform. As columns were
 used columns of Polymer Laboratories Ltd with pore diameters of 100-10000
 .ANG.. The used detector was RI-, i.e., refractive index detector,
 manufactured by Waters, and a 55 min run time with a flow rate of 1 ml/min
 were used. To determine the molecular weights of the samples were used
 polystyrene (PS) standards manufactured by Polymer Laboratories, and the
 calibration curve based on the same. Because there is no experimental
 Mark-Houwink constants a and K available, the molecular weights in the
 Table 1 are not absolute molecular weights for the samples but relative
 values in comparison with PS standards.
 TABLE 1
 SnOct- Glycerol
 .epsilon.-CL conc. conc.
 D,L-LA- mol/ mol/ GPC results DSC
 results
 ratio mol .SIGMA. mol .SIGMA. M.sub.n M.sub.w
 T.sub.m T.sub.g
 Example (M-%) monomers monomers (g/mol) (g/mol) PD (.degree.
 C.) (.degree. C.)
 1 100/0 0.0001 0.005 -- -- -- 56 --
 2 80/20 0.0001 0.005 35000 50000 1.4 47 --
 3 80/20 0.0001 0.005 40000 60000 1.5 42 --
 4 80/20 0.0001 0.005 40000 60000 1.5 45 --
 5 80/20 0.0001 0.0005 165000 272000 1.65 46
 -53
 6 80/20 0.0001 0.25 -- -- -- -- --
 7 100/0 0.0001 0.25 4300 5200 1.2 35 --
 8 100/0 0.0001 0.05 445 729 1.6 -- --
 9 100/0 0.0001 0.05 -- -- -- -- --
 10 100/0 0.0001 0.25 2000 2600 1.3 -- --
 11 100/0 0.0001 0.0125 -- -- -- 53 --
 12 100/0 0.0001 0.023 10000 12000 1.2 -- --
 13 10010 0.0001 0.034 -- -- -- 43 --
 14 80/20 0.0001 0.25 1100 1400 1.3 -- --
 DSC Measurements
 In the DSC measurements the 5-10 mg sample was heated with a rate of
 10.degree. C./min in a calorimeter chamber. In order to get a similar
 thermic history for all the samples, the samples were heated above their
 melting temperature to temperature of +80.degree. C. and cooled down to
 about -50.degree. C. The Tm and Tg values were determined from the curve
 recorded from the second heating, and they are presented in the Table 1.
 In FIG. 4 is presented the DSC curve of the product prepared in Example 3,
 which is typical for all the polymers according to the invention.
 Presentation of the Characteristic Monomer Ratio-molecular Weight
 Dependence for .epsilon.-CL/D,L-LA Copolymers
 When the monomer ratio-molecular weight value pairs corresponding to the
 polymer products from Examples 1-14 (Table 1) and being suitable for
 implantation use are presented graphically in a right angle co-ordinate
 system, one can see that the hit points are located on the area between
 two curves (the shaded area in FIG. 1). It is apparently possible to
 adjust the properties of the implant material according to the invention,
 of which melt viscosity, degradation rate in the body and mechanical
 properties in a solid state are the most significant, by means of FIG. 1
 based on Examples 1-4 by selecting suitable combinations of comonomer
 ratio and polymerization parameters (of which the concentration of the so
 called initiator compound can be mentioned here). It is self-evident to
 anyone skilled in the art that the method to regulate material properties
 described in the invention can also be used with other monomer
 combinations suitable for the production of biodegradable implant
 materials.
 Examples 15-26
 Dependence Between Biodegradability and Monomer Composition of
 .epsilon.-CL/D,L-LA Copolymers Plasticizable in the Range of Temperatures
 from 37 to 55.degree. C.
 The biodegradability of typical materials prepared in the manner described
 above was tested by hydrolysis experiments in a buffered aqueous solution
 at a temperature of 37.degree. C. (Examples 15-26; Table 2). The change of
 sample weight was followed with time so that their degradation rate was
 demonstrated. The hydrolytic degradation is the first, and thus the
 limiting, stage in regard to biodegradation, and therefore reflects quite
 well the overall total degradation rate in vivo, too. Based on
 experimental results FIG. 2 illustrates the depedence between monomer
 composition and degradation rate in polymers according to the invention.
 Degradation rate of a polymer in certain conditions is affected not only
 by the composition and consequent differences in hydrophilic character but
 at least by the average molecular weight, and this is why also samples
 having the same composition may have different degradation rates. The
 values presented in Table 2 can be regarded as typical examples for the
 materials according to the invention. Naturally they do not, however,
 represent the only possible property combinations of the implant materials
 produced with the method according to the invention.
 TABLE 2
 Composition
 Example .epsilon.-CL/lactide Molecular weight
 nr. (M-%) (g/mol) Decomposition rate
 15 0/100 10 000 1 d
 16 10/90 10 000 2 d
 17 10/90 30 000 1 week
 18 20/80 10 000 2 d
 19 20/80 30 000 2 weeks
 20 40/60 10 000 3 d
 21 40/60 50 000 1 month
 22 60/40 10 000 5 d
 23 60/40 80 000 6-10 months
 24 80/20 20 000 3 weeks
 25 80/20 300 000 1 year
 26 100/0 80 000 over 1 year