TISSUE ENGINEERINGVolume 1, Number 3, 1995Mary Ann Liebert, Inc.

Fabrication and Characterization of PLA-PGAOrthopedic Implants

C. MAULI AGRAWAL, Ph.D., P.E., GABRIELE G. NIEDERAUER, Ph.D., andKYRIACOS A. ATHANASIOU, Ph.D., P.E.

ABSTRACT

Fabrication methods and property characterization of polyglycolic acid (PGA), polylacticacid (PLA), and their copolymers are reviewed. Both of these aliphatic polyesters belong tothe a-hydroxy group and biodegrade in a physiological environment to monomeric acids,which are readily processed and excreted from the body. The physical and mechanical char-acteristics discussed include molecular weight, crystallinity, stress-strain behavior, perme-ability, and melting/glass transition temperatures. The most common methods of fabricat-ing PLA-PGA materials into medical devices are described.

INTRODUCTION

THE USE OF BIODEGRADABLE MATERIALS for the fabrication of orthopedic implants has seen unprecedentedgrowth in recent years. Out of the relatively small group of materials that are both biocompatible and

biodegradable, the family of polylactic acid (PLA), polyglycolic acid (PGA), and their copolymers has cer-tainly been the most studied. In the field of orthopedics these materials have been used to develop sutures,controlled release systems for drugs and other bioactive agents, scaffolds for regenerating both cartilageand bone, fracture fixation plates and screws for bone, as well as devices for ligament and tendon healing.To accurately match the properties of these devices with the specific requirements of each application, thefabrication process as well as the starting material have to be carefully chosen. The properties of the start-ing material often influence the fabrication parameters and the properties of the final device. At the sametime the fabrication process can significantly alter the starting material properties. Thus, it is important tounderstand the interplay between the molecular structure, material properties, and fabrication processes.

STRUCTURE

Polylactic acid (PLA) and polyglycolic acid (PGA) are both aliphatic polyesters and belong to the a-hy-droxy group. PLA can exist in two stereoisomeric forms: D and L. L-PLA crystallizes in pseudoorthorhom-bic or hexagonal forms1 and has a typical crystallinity of about 37%.2 Also, L -PLA can have two crystallinemodifications, a and /?; a is a helix conformation and (3 is an extended helix conformation.1 PGA has asimple linear structure and typically exhibits a crystallinity of approximately 50%.2 Unlike PLA, PGA doesnot have a methyl group, and this fact contributes to differences in their degradation kinetics. When sub-

Department of Orthopedics, The University of Texas Health Science Center, San Antonio, Texas 78284-7774.

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jected to high temperatures under vacuum, PLA, PGA, and their copolymers thermally degrade to form lac-tides and glycolides.2

To synthesize high-molecular-weight PLA and PGA, ring opening polymerization of the cyclic diestersglycolide or lactide is most commonly used.12 Often catalysts such as antimony, zinc, lead, or antimonyare utilized for the polymerization process.12 However, low-molecular-weight homo- and copolyesters oflactic acid and glycolic acid have also been synthesized by direct polycondensation in the presence of wa-ter without using catalysts.34 Typically, polymerization of PLA and PGA requires 2-6 h at temperaturesof approximately 175°C.5 The purity of the starting materials and processing humidity are critical parame-ters in obtaining quality polymers.5

DEGRADATION

PLA and PGA biodegrade mainly by nonspecific hydrolytic scission, where the polymer chains are es-sentially cleaved by simple hydrolysis of the ester linkages.56 For example, PLA undergoes hydrolytic scis-sion to its monomeric form, lactic acid, which is eliminated from the body by incorporation into the tri-carboxylic acid cycle. The lactic acid is finally excreted by the lungs as CO2 and in urine, with its principalelimination path being respiration.7 PLA implants labeled with radioactive carbon (14C) and placed subcu-taneously in rats for 3 months resulted in no significant radioactivity in feces or urine, confirming that thepolymer is degraded and probably eliminated through CO2 during respiration.8 PGA can be broken downin two ways, by hydrolysis and by nonspecific esterases and carboxypeptidases.9 Like lactic acid, the gly-colic acid monomer is either excreted in urine or enters the tricarboxylic acid cycle.

The degradation rate of PLA, PGA, and their copolymers depends on numerous factors such as molarratio of the constituents,3610 blood supply at implant site,11 crystallinity of the polymers,31012 thermal his-tory,12 inherent viscosity (average molecular weight),12 geometry and available surface area/porosity for in-teraction with tissue,11"13 uptake of water, and wettability of the polymer.314 Table 1 gives the range ofdegradation rates for PLA, PGA, and their copolymers. An examination of these rates shows that the half-life of these polymers in terms of molecular weight can be varied by altering the PLA to PGA ratios. Thebroad range of degradation times for each copolymer can be attributed to differences in the shape/size, pro-cessing, and sterilization procedures used for the specimens, as well as the molecular weight and distribu-tion of the starting material.

Polymer hydrolysis usually starts in the amorphous portion of the specimen as water can intrude easilyinto these regions. Therefore, a partly crystalline polymer will be preferentially degraded in the amorphousportion, leaving the crystalline regions temporarily intact. This leads to an overall increase of the crys-tallinity during degradation.1516

Researchers have also shown that enzymes play a definite role in the breakdown of lactide/glycolide ma-terials.9'17'18 After testing the effects of 15 different enzymes on the in vitro hydrolysis of PGA, it wasshown that the degradation of PGA in aqueous media is significantly influenced by enzymes with esteraseactivity.17 In addition, degradation has been shown to vary as a function of location in the implant. For ex-ample, degradation of high-molecular-weight PLA was found to proceed more rapidly in the center of the

TABLE 1. DEGRADATION RATES FOR PLA, PGA, AND SELECTED COPOLYMERS

Composition Degradation time Reference

Polylactide (PLA) 6 months-over 4 years 5, 6, 10, 12, 13, 60

Poly-DL-lactide (DL-PLA) 24 weeks-18 months 5, 10, 12Polyglycolide (PGA) 2-5 months 5. 6, 10, 12, 2750PLA:50PGA 7-60 days 5 ,6 ,36 ,61 ,6270 PLA:30 PGA 30 weeks 485 PLA: 15 PGA 90-240 days 3, 5, 10, 1290 PLA: 10 PGA 2 months 5, 10, 12

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implant than at the surface.13 Furthermore, the change in molecular weight of L -PLA ultrahigh strength rodstested in rabbits was found to depend on implantation site and environment. Molecular weight was reducedmost in the medullary cavity, followed by the subcutis, and in vitro degradation in phosphate-buffered salineat 37°C.19 By 8 weeks the molecular weight had decreased 75% from 220 to approximately 50 kDa.

The biodegradation process of aliphatic polyesters is initially manifested by a decrease in molecularweight due to random hydrolytic cleavage of the ester linkage, followed by a decrease in mass and me-chanical strength.20 As mentioned earlier, crystallinity also appears to increase upon degradation as shownin a study of initially amorphous PLA, where percent crystallinity changed from 0 to 49% over a 90-weekperiod.13

Methods for characterizing the degradation of these polymers have been developed both in vivo and invitro. In vitro degradation assessment usually involves exposing polymer specimens to a physiological en-vironment, such as saline or aqueous buffer solution at 37°C. After various exposure times, samples areusually vacuum dried and molecular weight or mass changes, mechanical properties, intrinsic viscosity,crystallinity, and other properties may be determined. For intrinsic viscosity measurements, a 0.5% poly-mer solution in chloroform at 25°C (ASTM Standard D445-88) can be used.14 Molecular weight measure-ments are routinely performed using gel permeation chromatography10-21 or by terminal-carboxyl-groupanalysis.3 An accelerated in vitro hydrolysis test at 80°C has been suggested as a tool for preliminary analy-sis to compare the degradability of various polymers.22 Using this procedure, L-PLA and D-PLA were com-pared, and L -PLA was found to have higher viscosity over time, thus, indicating least molecular weightloss. In vivo degradation has been quantified by measuring the temporal changes in molecular weight or in-trinsic viscosity of PLA-PGA materials by excising implants that have been placed in animals for varioustimes.310 Comparison of in vitro and in vivo degradation has shown that the results are comparable.23

As expected, the implant's functional environment affects its in situ degradation characteristics. For ex-ample, the in vivo degradation response of L -PLA plates has been shown to be affected by mechanical load-ing such that the tensile strength of bone plates subjected to mechanical stresses decreased in comparisonto those not loaded.24 Interestingly, no corresponding significant changes were detected in the molecularweight.

To decrease the degradation rate of PLA-PGA devices, coatings have been used to act as barriers to hy-drolysis by Vasenius et al.25 However, they reported that even though the coatings of slowly absorbingpolymers assisted in strength retention in vitrof no significant effects were observed in vivo.

PHYSICAL AND MECHANICAL CHARACTERISTICS

Orthopedic biodegradable materials, such as PLA-PGA, are often intended to act in lieu of metallic im-plants for internal fixation. Obviously, the mechanical properties of such devices play a significant role inthe support of healing bone, such that gross motion at the repair site promoted by insufficient stiffness re-sults in inadequate healing, or that complete elimination of micromotion results in stress-shielding. In manyinstances, PLA-PGA materials are used as porous scaffolds for tissue ingrowth, and, therefore, the stressfields and fluid flow patterns created in the scaffolds, as a result of the functional environment, may bedetrimental to existing or migrating cells due to a mismatch in mechanical properties with surrounding tis-sues. It is thus necessary to critically evaluate the mechanical characteristics of these devices—such as stiff-ness, strength, Poisson's ratio, toughness, permeability—immediately prior to implantation. An additionaldifficulty is created by the fact that these properties are significantly altered with biodegradation, which im-plies that rates at which mechanical properties change in situ must be determined a priori. The mechanicalproperties of PLA-PGA polymers are usually a function of their molecular weight, crystallinity, molecularorientation, porosity, pore size, and pore interconnectivity.

Molecular Weight

One of the most important physical properties of PLA, PGA, and their copolymers is molecular weight,since it greatly influences these materials' mechanical and degradation characteristics. The entire functionalresponse of medical PLA-PGA devices hinges significantly upon the initial value of molecular weight,

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which is also affected by the production or manufacturing methodology, and in vivo temporal variationsdue to the in situ mechanical and physiological environments.

Initial molecular weights of various biodegradable polymers either commercially available or measuredby gel permeation chromatography have been reported to vary widely.26 For L -PLA, values of weight-av-erage molecular weight (Mw) have been reported to range from 50 to 756 kDa. For DL-PLA, M W rangesfrom 21 to 550 kDa.26 PLA-PGA polymers of various molecular weights (ranging from 20 to 550 kDa forPL A and 20 to 145 kDa for PGA) can be purchased from numerous suppliers. Number-average molecularweight values reported in the literature range from 19.6 to 150 kDa for L - P L A and 13.4 to 163 kDa for DL-

P L A . Another measure of the average size of molecular chains in a polymer is its intrinsic viscosity, whichhas also been shown to vary widely from 0.61 to 8.2 dl/g for L -PLA and from 0.25 to 2.01 dl/g for DL-

P L A . 1 2 6 Owing to its insolubility in common organic solvents, such as chloroform or methylene chloride,molecular weight measurements for PGA are relatively uncommon. PGA used for fiber extrusion is solu-ble in hexafluoroisopropanol and has an intrinsic viscosity of 0.6-1.6 dl/g at 0.5% solution, which corre-sponds to an Mw range of 20-145 kDa.27

During fabrication of PLA-PGA implants or other medical devices, the materials are often exposed tosignificant mechanical stresses or fluctuations in temperature and pressure, with sometimes detrimental re-sults as far as molecular weight is concerned. As an example, an almost 50% reduction in initial viscosity-average molecular weight occurs for L -PLA following extrusion and drawing.19 Obviously if the potentialfor significant alterations in molecular weight is expected during production, it is essential that appropriatedesign measures be implemented so as to account for such reductions in the final product's molecular weight.

Crystallinity

Adjacent sections of polymeric molecular chains can sometimes pack into a stable crystalline arrange-ment, which represents an ordered structure. Most polymers display low crystallinity, never reaching 100%,and are thus either amorphous or semicrystalline. As a general rule, the greater the crystallinity of a poly-mer the greater are its stiffness and density. A measure of crystallinity can be obtained from density mea-surements. The density of PGA has been reported as 1.5-1.64 g/cm3,27 while densities of 1.29 and 1.25g/cm3 have been reported for the crystalline and amorphous phases of L -PLA, respectively.1

There is no linear proportionality between the glycolic acid/lactic acid ratio and the crystallinity of thecorresponding copolymer. PGA is highly crystalline, but crystallinity rapidly decreases in PLA-PGA copoly-mers. The degree of crystallinity depends upon the molecular chemistry and chain structure, temperature,and the rate of cooling during solidification from a melt. Molecular structure is important because sidebranches and cross-linking hinder the mobility of chains. Molecular chemistry plays a significant role indetermining crystallinity because monomer units containing large or complex chemical species render crys-talline order very difficult. Elevated temperatures and a slow rate of cooling enables the chains to be mo-bile and realign themselves in a more ordered solid structure. Thus, the crystallinity of PLA-PGA poly-mers can be altered as a result of fabrication processes where heat is used. For instance, a 74% crystallinityhas been reported for L -PLA screws fabricated from a 65% crystalline batch of L - P L A . 2 8 The percent crys-tallinity of a polymer sample can be estimated by X-ray diffraction, by infrared spectroscopy, or by mea-suring its heat of fusion using a differential scanning calorimeter and relating this measured value to theheat of fusion for a known degree of crystallinity. PGA sutures typically are 46-52% crystalline,2 while thecrystallinity of L -PLA has been reported to range from 15 to 74%.2'26*28

DL-PLA is predominantly amor-phous. Copolymers of L-PLA and PGA are usually amorphous when the PGA content is in the range of25-70%. The same is true for DL-PLA and PGA copolymers when PGA content is 0-70%.2

Stress-Strain Properties

The two physical characteristics discussed above, i.e., polymer molecular weight and crystallinity, aredirectly responsible for rendering PLA-PGA polymers strong and functionally capable. Usually as molec-ular weight and crystallinity increase so does the implant's strength. It should be noted that mechanicalstrength does not follow a linear relationship with an increasing ratio of glycolic acid to lactic acid. Themechanical properties (tensile strength and modulus) of DL-PLA are dramatically improved with increas-

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ing molecular weight.26 Tensile strength of L -PLA can also be improved with increasing crystallinity.1 Ithas been suggested that L -PLA or DL-PLA should possess a molecular weight of at least 100 kDa to achieve"good mechanical properties1' necessary for orthopedic applications such as nails.26 In the same study, therelative brittleness of L -PLA compared to DL-PLA was also presented.26 Ductility, like most other me-chanical properties, can be manipulated by altering fabrication techniques or starting with different molec-ular weights. For example, hot drawn melt-spun and solution-spun fibers have shown 12-26% elongationsat break,1 while braided melt-spun L -PLA fibers have been reported to exhibit initial elongations of ap-proximately 15-35%.29

In a review paper covering studies performed on various biodegradable polymers and composites be-tween 1980 and 1988 it was reported that PLA (including L-PLA and DL-PLA) has been shown to exhibittensile and flexural strengths of 11.4-72 and 45-145 MPa, respectively, with tensile and flexural moduliof 0.6-̂ 4- and 2.4-10 GPa.30 Fiber reinforcement of PLA employed by a number of researchers was reportedto result in significant increases in tensile and flexural characteristics; tensile and flexural strengths increasedup to 200 and 412 MPa, respectively, with tensile and flexural moduli increasing up to 29.9 and 124.4 GPa,respectively.30 Percent elongations at yield and break for L-PLA have been reported to be between 1.8 and3.7 and 2.0 and 6.0%, respectively.26 Unreinforced PGA has tensile strength of 57 MPa and tensile modu-lus of 6.5 GPa, but is exceedingly brittle, failing at only 0.7% elongation.3^elf-reinforced PGA rods werereported to have significantly increased flexural strength (370 MPa) and shear strength (250 MPa), but by5 weeks in distilled water the flexural strength dropped to only 5% of its original value.32 Using a sinter-ing technique to achieve self-reinforcement33 or a drawing technique,27 it has been shown that PGA can at-tain elongations of 7-8 or 15-35%, respectively.

Strength and stiffness characteristics of PLA-PGA copolymers, which are expected to depend on theirglycolic acid to lactic acid ratio, are generally lower than their homopolymers and have been reported tovary significantly. For example, a 90% PGA-10% PLA copolymer had tensile and flexural strengths of 45and 150 MPa, respectively, which were reduced to 4 and 7% of these initial values after 4 weeks in dis-tilled water.34 Reinforcement of this copolymer increased the initial tensile and flexural strengths; for ex-ample 1% volume of carbon fiber bundles increased the tensile and flexural strengths to 90 and 190 MPa,respectively.34 Self-reinforced PLA-PGA rods were reported to exhibit ductile behavior, while melt moldedPLA-PGA rods exhibited brittle characteristics.34

Permeability

Permeability is a significant mechanical characteristic of biodegradable, porous structures, which is of-ten overlooked. By definition, permeability is the ability of an object to be traversed by another substance.In orthopedic applications, where PLA-PGA scaffolds are used to encourage repair of musculoskeletal tis-sues, permeability denotes the ease or difficulty of body fluids (e.g., vascular supplies, synovial fluid) tomove through the pores of the biomaterial. Obviously, this property can have profound effects in modulat-ing the repair process, since it regulates the volume and flow rate of nutrient-carrying fluids, which are nec-essary for initiating or maintaining cellular activities in and near the repair site. A large permeability canalso be essential in aiding degradation through rapid evacuation of byproducts. It can additionally assist inproviding a buffering action to maintain the pH, which can be severely affected by the acidic byproductsof degradation, especially in the vicinity of PLA-PGA implants. The degree of permeability can also af-fect the release of bioactive agents carried by the implant. A suitable means to measure this mechanicalproperty is by applying Darcy's Law, which can be conveniently expressed by

hAt

where k is the permeability constant, Q is the quantity of discharge, L is the length of the sample in the di-rection of flow, A is the cross-sectional area of the sample, h is the hydraulic head, and t is the time. Thismechanical property can be expected to be related to the material's porosity, pore size, and pore intercon-nectivity.

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To overcome the inherent difficulty of water entering the usually air-filled pores of PLA-PGA foams, atwo-step immersion in ethanol and water was recently reported.35 For L -PLA porous discs, void volumefilled by water was reported to have increased from 23 to 79% after 1 h prewetting in ethanol. For PLA-PGAcopolymers the increase was from 59 to 97%. Of course, porous materials such as the ones used by Mikoset al.35 or Athanasiou and co-workers36 are expected to exhibit significant values of hydraulic permeabil-ity. Such values can easily be measured using a direct permeation experiment and applying Darcy's Law.

Melting and Glass Transition Points

Glass transition temperature and melting temperature, obtained from a specific volume versus temperatureplot, represent properties of the amorphous and crystalline phases of the polymer, respectively. For L-PLA,

glass transition and melting temperatures have been reported to be in the range of 54-59 and 159-178°C, re-spectively.1'2'2637 Glass transition temperature for the amorphous DL-PLA is in the range of 50-53°C.26 ForL-PLA and DL-PLA, there is a slight tendency for glass transition temperature to increase with molecularweight; no such relationships are shown between molecular weight and melting point, or degree of crys-tallinity.26 PGA exhibits a glass transition temperature of 36°C, while its melting temperature ranges from 210to 226°C.2'26'27 PLA-PGA copolymers have glass transition temperatures of 37-55°C.2-38

FABRICATION

PLA, PGA, and their copolymers are thermoplastic in nature, which implies that upon heating they softenand melt. In addition, they are usually soluble in several organic solvents. As a result, these polymers arehighly conducive to being formed into intricate final shapes using a variety of techniques. The choice of aparticular technique is often dictated by the physical and thermal properties of the specific material beingprocessed as well as the desired form and properties of the final product.

Fiber Fabrication

One of the most widely used forms of PLA-PGA polymers are fibers, which are used extensively to fab-ricate sutures and also to reinforce composites. Fibers can be produced by solution-spinning, melting fol-lowed by extrusion or dry spinning, or a combination of these processes.837 '3940 Eling et al.1 comparedmelt spun and solution spun L -PLA fibers. Solution spun fibers were fabricated by extruding an L - P L A so-lution in toluene at 110°C while melt spun fibers were made by melt extrusion at 185°C. Above this tem-perature it was not possible to melt spin because of the low viscosity of the melt. The authors determinedthat in general solution-spun fibers show improved tensile properties compared to those melt-spun. For ex-ample, solution-spun and melt-spun L -PLA fibers exhibited tensile strengths of 1.0 and 0.5 GPa, respec-tively. This difference was attributed to a lower number of entanglements in the solution-spun fibers. Hyonet al.37 determined that melt-spun acetylated L -PLA fibers have tensile properties comparable to those ofconventional crystalline polymers. Dauner et al.40 reported that the surface of melt-spun fibers is smoothercompared to those solution-spun.

The molecular weight of a polymer plays an important role in determining the parameters for fiber fab-rication. As the molecular weight increases the viscosity of the melt or solution for spinning increases cor-respondingly, thereby influencing the fabrication process. In a study on L -PLA, Suuronen39 noted that whilepolymers with a molecular weight of 180 kDa can be melt spun, suitable molecular weights for solution-spinning are in the range of 250 to 530 kDa.

The tensile properties of fibers can be significantly improved by hot drawing the fibers after spinning.1

Recently, Andriano et al.14 published an extensive study on the processing and characterization of PLAfibers. They examined DL-PLA polymers with different D/L ratios and determined that the tensile strength,elastic modulus, and birefringence increased up to a draw ratio of 6.7 and declined thereafter. These prop-erties are a strong function of both the draw ratio as well as the drawing temperature. The drawing processresults in improved alignment of the molecular chains leading to enhanced mechanical properties in the lon-gitudinal direction. In addition, the hot drawing process can also cause an increase in crystallinity.

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Once produced, fibers can be used as sutures in the form of monofilaments or can be braided to formmultifilament sutures. Frazza et al.27 presented an overview of PGA sutures with regard to their fabrica-tion, characterization, and in vivo evaluation. Sutures and fibers of PLA-PGA can also be used to reinforcebiodegradable composites, weave fabrics, or meshes, and fabricate three-dimensional scaffolds for tissueengineering.

Hot Molding and Machining

PLA-PGA plates, rods, and screws used in orthopedics are often fabricated using injection molding tech-niques.41 The physical and mechanical properties of the final product are in part determined by the mold-ing parameters. These properties include molecular weight, percent crystallinity, chain orientation, and stiff-ness. Leenslag et al.42 have described the use of 200°C and 18,000 kg "pressure" for fabrication of L -PLA

implants. Plates and screws can also be machined from larger blocks of material. For cases where the useof heat and solvents has to be restricted for a variety of reasons, PLA-PGA can be compression moldedunder high pressure with no heat application.43 In such cases the low-molecular-weight fractions in the poly-mer plastically deform and fuse or coalesce together to yield the final molded product.

Microparticle Formation

The use of microparticles for the delivery of drugs and proteins is steadily expanding. Although they arenot currently used much for orthopedic applications the recent growth in the field of tissue engineering ofbone and cartilage is likely to foster the increased use of microparticles in this area. Microparticles of PLAand PGA polymers can be readily formed because the polymers are soluble in organic solvents. Three dif-ferent techniques are commonly used for microparticle fabrication:5 phase separation, solvent evaporation,and fluidized bed coating.

Phase separation is useful for microencapsulating water soluble compounds in PLA-PGA polymers. Thisprocess involves dissolving the polymer in an organic solvent followed by coacervating it with the use ofa nonsolvent such as silicone oil.

Solvent evaporation is perhaps the technique most commonly used for microsphere formation and worksbest for encapsulating water-insoluble compounds. The polymer is first dissolved in an organic solvent andthe compound is added to this solution under agitation. The solution is then emulsified in a solution ofpolyvinyl alcohol or another surfactant in water under continuous stirring. The resulting microspheres areextracted by filtration or by removal of the volatile solvent under vacuum.

Grandfils et al.44 described the use of solvent evaporation for fabrication of DL-PLA microspheres forembolic material. They examined the use of different tensioactive agents (polyvinyl alcohol, gelatin, andmethylcellulose), stirring speed, and dispersed phase viscosity on the size of the microspheres. It was de-termined that the viscosity of the dispersed phase was the best parameter to alter to manipulate the size ofthe particles. In another recent study, Hafeli et al.45 used a solvent evaporation technique to encapsulateFe3O4 particles in PLA to produce magnetic microspheres that could be guided to a specific location in vivowith the aid of an external magnetic field.

As in the two techniques described above, the fluidized bed technique uses a solution of the polymer inan organic solvent. This solution is then processed through an air suspension coating system to produce mi-crospheres.44

Gel Casting

Coombes and Heckman4647 described a gel casting technique to produce microporous implants ofPLA-PGA polymers. This process involves first dissolving the polymer in a solvent such as acetone. Thesolution is then poured into a mold and allowed to stand at room temperature until it forms a gel. The gelis extracted and processed through several stages of solvent exchange in mixtures of acetone, ethanol, andwater to yield a microporous solid implant. Bioactive factors may be incorporated in these implants byadding them to the starting solution.4849 In one embodiment of this type of implant, Agrawal et al.49 useda 50:50 PLA-PGA copolymer to fabricate implants with bone morphogenetic protein and soybean trypsininhibitor. Studies examining protein release kinetics determined that a high percentage of the proteins in-

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corporated in these devices were released in the first 48 h even though the release continued thereafter formore than 70 days. The advantage of this technique is that it uses low heat (<45°C), so the probability ofdenaturing bioactive agents is low. In addition, implants of complicated shapes can be fabricated.

Solution Casting

PLA-PGA polymers can be dissolved in appropriate organic solvents and molded into different shapesby extracting the solvent using either evaporation or the addition of a nonsolvent such as ethanol, where-upon the polymer precipitates. Schmitz and Hollinger11 have described such a technique. They solubilizeda 50:50 copolymer of PLA-PGA in chloroform, precipitated it with the addition of methanol, and com-bined it with demineralized freeze-dried bone. The doughy composite was then forced in molds and sub-jected to temperatures of 45-48°C for up to 24 h. The resulting implants were surgically placed in the cal-varia of rabbits.

Heckman et al.50 have reported the solution casting of PLA implants with bone morphogenetic proteinto treat fracture nonunions in a canine model. In a series of reports published by Athanasiou et al.36'51'52

the fabrication and use of solution cast implants for treating cartilage defects have been described. Theseimplants were fabricated from a 50:50 copolymer of PLA-PGA, which was dissolved in acetone and pre-cipitated in ethanol. The resulting "gummy" precipitate was packed into Teflon® molds, placed in a vac-uum, and subjected to a specific temperature regime. The resulting microporous structure is shown inFigure 1.

Solvent-Casting Paniculate Leaching

Several studies have recently reported the fabrication of biodegradable foams or scaffolds using a par-ticulate leaching technique.53'54 Mikos et al.53 described a process wherein different copolymers ofPLA-PGA are dissolved in chloroform. Sieved particles of sodium chloride were added to these solutions,which were then vortexed and allowed to dry by evaporation. The salt to polymer ratio was 9:1. The resid-ual solvent was extracted by vacuum treatment. The sodium chloride crystals were leached out by immer-sion in deionized water at 25°C for 48 h to yield porous membranes. Three dimensional scaffolds were

FIG. 1. Scanning electron micrograph of microporous PLA-PGA polymer implant showing an average pore size of150 ^m and a porosity of 60% by volume (magnification 80 X).

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formed by laminating these membranes. The growth of cells on these scaffolds has been investigated invitro.53*54

Fiber Bonding

A unique technique for forming three-dimensional scaffolds has been developed by Mikos et al.55 L -PLA

polymer was dissolved in chloroform and added to a Petri dish containing a nonwoven mesh of PGA fibers.After the solvent was removed by evaporation, the resulting PLA-PGA composite was subjected to a tem-perature of at least 195°C for 90 min. The composite was removed from the oven and immersed in liquidnitrogen followed by air drying and vacuum treatment. Next the L -PLA matrix of the composite was re-moved by dissolution in methylene chloride utilizing the fact that PGA is insoluble in this solvent. Thisprocess yielded a scaffold of PGA fibers bonded together by the heat treatment. The seeding and prolifer-ation of cells isolated from rats on these scaffolds have been investigated by the researchers.55

Composites

The mechanical properties of PLA-PGA materials are usually inferior to those of bone and hence theuse of these polymers in bulk form may not be adequate for fracture fixation. To address this problem re-searchers have attempted to increase the mechanical strength of such biodegradable devices using fiber re-inforcement. As mentioned previously, Daniels et al.30 compiled an extensive review of such implants. Thereinforcing components include PLA, PGA fibers, carbon fibers, and ceramics. Several studies have re-ported the use of self-reinforced PLA-PGA and PGA rods.3256"58 Self-reinforced PGA rods were fabri-cated by Tormala et al.33 by sintering together bundles of PGA sutures (Dexon®) at temperatures of205-232°C for time periods of 3 to 7 min under high pressure. In another study by the same group of re-searchers melt molded, self-reinforced, and carbon fiber-reinforced PLA-PGA composites were compared.34

It was determined that self-reinforcement significantly improved the flexural strength of the rods, while thecarbon-reinforced rods exhibited high tensile strengths. In all cases, reinforced materials were determinedto be superior to the melt-molded devices.

Suuronen et al.28 reported on the use of melt extruded L-PLA plates. The molecular orientation in thesedevices was further increased by die-drawing at 150°C to a draw ratio of 4. The enhanced orientation ofthe molecular chains in the draw direction results in self-reinforcement of the plates. A technique for rein-forcing PLA matrices using a PGA fabric has also been described.59 Such a configuration results in en-hancing the properties of the resulting devices in more than one direction.

In summary it is important to note that the properties of PLA-PGA devices are a strong function of boththe starting material and the fabrication process. Although, PLA-PGA polymers offer the convenience ofbeing formed into devices using a variety of techniques, these techniques have the potential of significantlyaltering the properties of the starting material. Thus, it is imperative to bear this point in mind while choos-ing a particular fabrication technique. In addition, the properties of the device should be carefully evalu-ated after fabrication and prior to use.

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Address reprint requests to:Dr. C. Mauli Agrawal

Department of OrthopedicsThe University of Texas Health Science Center

7703 Floyd Curl DriveSan Antonio, TX 78284-7774

252

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