The development of polyanhydrides for drug delivery applications.

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The development of polyanhydrides for drug delivery applications a

J. Tamada & R. Langer

b

a

Room E25-342, Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, MA 02139, USA b

Room E25-342, Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, MA 02139, USA Published online: 02 Apr 2012.

To cite this article: J. Tamada & R. Langer (1992) The development of polyanhydrides for drug delivery applications, Journal of Biomaterials Science, Polymer Edition, 3:4, 315-353, DOI: 10.1163/156856292X00402 To link to this article: http://dx.doi.org/10.1163/156856292X00402

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Review The

development

of

polyanhydrides

for

drug

delivery

applications


J. TAMADA and R. LANGER Room E25-342,Department of ChemicalEngineering,MassachusettsInstitute of Technology, Cambridge, MA 02139, USA Received5 July 1991;accepted 23 September 1991 Abstract-This paper reviewsthe developmentof the polyanhydridesas bioerodible polymers for drug delivery applications. The topics include design and synthesis of the polymer, physical properties, techniques to fabricate the polymer into drug delivery devices, evaluation of biocompatibility, and example applications of the polyanhydrides. Discussionof the interrelationship between the physicalchemicalproperties of the polyanhydrides, fabrication methods, and drug release rates is included. One section is devotedto a case study to provide a historical perspectiveof the developmenta polyanhydridebased drug deliverytreatment from the conception of the idea to the final stages of human clinicaltrials. This section includesan outline of the extensivein vitro and in vivo testing that is necessaryfor development of a new material for biomedical applications. Key words: Polyanhydride; biodegradable polymer; drug delivery;controlled release. INTRODUCTION For many years, the majority of materials used in medical devices were originally designed for other applications. For example, the fabric used in artificial hearts was once used to make women's girdles; dialysis membranes for artificial kidneys were adapted from sausage casings; ethylene vinyl acetate, now used in nonerodible controlled release drug delivery devices, was originally used in wall coatings; polyurethane foams, used in breast implants, were originally used to stuff mattresses [1]. More recently, we and others have suggested moving beyond this practice of using 'off the shelf' materials for biomedical applications to the development of materials specifically synthesized for biomedical devices. This review will discuss the polyanhydrides, a class of bioerodible polymers that were developed specifically for controlled release drug delivery applications. It will describe efforts to understand their mechanical, chemical, and biological properties so that they can be rationally designed for the special needs of specific medical applications. Topics in this review device fabrication include polymer design and synthesis, physical properties, and example applications of the polyanhymethods, evaluation of biocompatibility, drides. One section is devoted to a case study to provide historical perspective of the development a polyanhdride-based drug delivery treatment from the initial conception of the idea to the final stages of human clinical trials. Controlled release of a variety of therapeutic agents is possible through the use of bioerodible polymeric drug delivery systems. In these systems, the drug is incormixture is formulated into porated into a bioerodible polymer. The polymer-drug devices suitable for implantation into the body. The implanted device slowly erodes upon contact with body fluids, releasing the drug to the body. Implanted controlled


316 release systems provide several advantages over conventional oral or injectable drug formulations. Delivery can be localized to the site of implantation, which lowers the drug dosage, thereby reducing potential systemic side effects. Drug delivery rates are steady and controlled, which can better maintain the drug concentration within its over nondegradable window. Bioerodible devices have advantages therapeutic implants. The main advantage is that bioerodible devices are absorbed by the body, which bypasses the need for surgical removal of the device. Additionally, because drug delivery is controlled primarily through the properties of the polymer, the release of both conventional low-molecular-weight drugs and the release of macromolecular drugs, including hormones (e.g., insulin, growth hormone), polysaccharides (e.g., heparin), antibodies, antigens, and enzymes, is possible. However, bioerodible systems pose special challenges. The first challenge is to develop a polymer system which will deliver the drug at a rate and for a long enough period of time to achieve a therapeutic effect. The second challenge relates to the constraints imposed upon materials which are used for biomedical applications. Both the polymer itself and its degradation products must be biocompatible; i.e. must be and and must not nontoxic, noncarcinogenic, nonmutagenic, promote they undue inflammatory or allergic reactions in the patient. Finally, the device must allow the drug to maintain its potency for the duration of the release period. These challenges require extensive testing procedures in order to ensure that the polymer treatment is safe and effective in providing the desired physiological effect. A number of biodegradable polymer systems have been studied for controlled release applications. Degradation can be based on enzymatic or hydrolytic breakdown of the polymer. Generally, the hydrolytic degradation mechanism is considered to be the more desirable option, because there is less variation in breakdown among different implantation sites and different patients. It has also generally been considered desirable for a polymer system to undergo 'surface erosion' kinetics, i.e., for the polymer to erode like a bar of soap from the outside to the inside and to exclude water penetration into the bulk of the matrix. Ideal surface erosion provides release at a rate proportional to the surface area and aids in the delivery of water-labile drugs by minimizing water interaction with the drug prior to release. However, most currently available biodegradable systems, e.g., poly(lactide-co-glycolide) copolymers [2], undergo 'bulk erosion'; water penetrates into the polymeric matrix and degrades it internally as well as externally. Other systems undergo surface erosion, e.g., polyorthoesters [3], but require the use of additives to achieve these erosion kinetics. In the design of a bioerodible polymer system which might perform closely to an ideal surface eroding polymer, we have suggested that polyanhydrides would be promising candidates. The anhydride linkage is water labile, and, by rational selection of monomer units, the polymer can be made sufficiently hydrophobic to discourage water penetration. Work on this concept began in the late 1970's, devoted to the synthesis of polyanhydrides that could be used for drug delivery applications. CLASSESOF POLYANHYDRIDES The polyanhydrides are composed of monomer units connected by anhydride bonds. The anhydride bond is hydrolytically labile and breaks down into two carboxylic acid groups (see Fig. 1). Thus, degradation of the polymer occurs by backbone chain scission across the anhydride bond. The initial long-chain polymer, which is water


317

Figure 1. Structure and degradation of the polyanhydrides. insoluble, is cleaved into shorter and more hydrophilic fragments, which are soluble and can be absorbed by the body. The structures of the monomers determine the properties of the polymer. Hence, careful selection of the monomers is a crucial element in the development of biomedically useful polymers. This section will detail the historical development of different classes of polyanhydrides and give insight into the reasoning involved in monomer selection. Historical

perspective

The first synthesis of polyanhydrides was reported in 1909 by Bucher and Slade [4]. In the 1930's a series of aliphatic polyanhydrides was prepared by Hill and Carothers for use in synthetic fabrics for clothing [5, 6]. However, the hydrolytic instability of the anhydride bond was an undesirable property for textile applications. Despite further development through the late 1950's and 1960's [7-13], polyanhydrides were never commercialized for textile use. In the early late 1970's and early 1980's, we at MIT were seeking an appropriate material for a bioerodible drug delivery system. We proposed that to achieve surface erosion, one would need hydrophobic monomers connected by water labile bonds. The bonds would have to be more labile than ester, amide, and ether linkages. Anhydride bonds possessed one of the few chemistries meeting these needs. We also observed that the degradative characteristics of the polyanhydrides, which are not desired in textiles, make them well suited to be materials in biodegradable drug delivery systems. This revitalized In our efforts, the focus of polymer developthe development of polyanhydrides. ment was directed by the special needs imposed by medical applications. Polyfbis(p-carboxyphenoxy)methane)J The first polyanhydride which embraced the concept of drug delivery was poly[bis(pcarboxyphenoxy)methane)], p(CPM) [14]. Figure 2 shows the structure of the literature monomer unit. This monomer was selected because of its hydrophobicity, data which indicated that its degradative characteristics might be appropriate for drug release, and based on a toxicological assessment that the monomers would be harmless and excretable. An in vitro erosion study of disks of the polymer in pH 7.4 phosphate buffer at 37°C was performed. Figure 3 shows the cumulative erosion of


318

Figure 2. Structure of monomer units used to make polyanhydrides. p(CPM) into buffer. Erosion occurs over a 30-day period, although there is a significant induction period of relatively little erosion at the beginning of release. Release of a model compound, cholic acid, incorporated into the polymer was then Cholic acid release was found to be performed to assess delivery characteristics. released with the monomer over a 6-day period into pH 7.4 phosphate buffer at 60°C. This initial study successfully demonstrated the feasibility of controlled release from polyanhydride matrices. Poly[], 3-bis(p-carboxyphenoxy)propane-co-sebacic

acid]

In order to accommodate the variety of drugs which might benefit from controlled release, an arsenal of bioerodible polymers needed to be developed. Each particular


319

Figure 3. Erosion of p(CPM). Deviceswere compression-molded1.4cm diameter x 0.5 mm thick disks. Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C. The cumulativeerosion was measured by UV absorbance at 243nm [14]. polymer composition would provide release of drug in a specific range of rates and duration. The p(CPM) homopolymer system lacked this capacity. The approach taken to achieve flexibility in the erosion profiles was to use two different kinds of monomers, that is, to formulate copolymers. This approach had been used for other biodegradable systems, such as poly(lactide-co-glycolide) [2]. Changing the ratios of the monomers changes the physical properties, including erosion rates. The use of copolymers has proven to be an effective and simple method of controlling polyanhydride erosion profiles. The polyanhydride copolymer system which currently enjoys the most widespread interest in its medical applications employs the monomers sebacic acid (SA) and 1,3-bis(p-carboxyphenoxy)propane (CPP) [15]. Figure 2 gives the structures of the monomer units. Figure 4 shows the effect of composition on the erosion of the

Figure4. Effect of copolymercompositionon the erosionprofile of p(CPP-SA). Deviceswerecopolymers p(CPP-SA) at 20, 45, 80 and 100 mole percent CPP compressionmolded into 1.4-cmdiameter x I mm thick disks. Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C. The cumulativeerosion is plotted as the appearance of the CPP monomer was measured by UV absorbance at 247nm [15].


320 polymer into pH 7.4 phosphate buffer at 37°C, as measured by cumulative CPP appearance. A typical 1 mm thick, wafer-shaped polymeric device composed of a homopolymer of sebacic acid would be mostly eroded in a few days, and completely eroded in 1 or 2 weeks (not shown). As the CPP content of the copolymer increases, 20 : 80 (20 mol% CPP, the erosion period is extended. A device of p(CPP-SA) 45 : 55 80 mol?lo SA) would take a few weeks to .erode; a device of p(CPP-SA) would take a few months. Extrapolations of erosion data indicate that a device made of the homopolymer of CPP would take at least 2 years to completely erode. Hence is a versatile system, which can potentially it was established that the p(CPP-SA) be used for a variety of therapeutic applications. The erosion rates of other related polyanhydride homopolymers and copolymers have been characterized. The monomer 1,6-bis(p-carboxyphenoxy)hexane (CPH), a close relative of the CPM and CPP monomers, has been used in both homopolymeric and copolymeric polyanhydrides [15]. The effect of the number of methylene groups in the backbone linkage is seen by comparing the erosion rates of the series p(CPM) (n = 1), p(CPP) (n = 3), and p(CPH) (n = 6), where n is the number of methylene of the polymer, i.e. the groups in the monomer. The greater the hydrophobicity slower rate. backbone the the erosion the longer linkage, A number of other systems have been synthesized, including poly(terephthalic acid), abbreviated as p(TA), and poly(isophthalic acid), abbreviated as p(IPh), polyacid and 1,3-benzenedicarboxylic mers of 1,4-benzenedicarboxylic acid, respectively acid used as more Dodecanoic has been a slightly [15]. hydrophobic alternative to sebacic acid in copolymerization. Aliphatic-aromatic

polyanhydrides

Although the p(CPP-SA) copolymer system has shown tremendous potential for application to many drug delivery needs, methods to engineer other classes of polyanhydrides with improved properties are under intensive development. For example, we sought a way to improve the degradation kinetics to provide a more linear release into mechanisms of polymer erosion determined that difprofile. Investigations ferences in the rate of erosion of the SA and CPP monomers contributed to the nonlinearity of the release profiles. The aliphatic SA monomers were found to erode significantly more quickly than the aromatic CPP units (see 'Comparison of release of comonomers' is a random copolymer (see below). The polymer p(CPP-SA) below). Therefore, the chain contains a distribution of CPP-CPP, 'Spectroscopy' and SA-SA linkages. Erosion is relatively rapid initially as the SA-SA CPP-SA, and SA-CPP bonds degrade. After the SA is depleted, a partially eroded device bonds is left and the device remnants erode very slowly. containing only CPP-CPP 45 : 55 composition as a steep initial This is manifested in Fig. 4 for the p(CPP-SA) slope in the cumulative erosion profile of the CPP release which curves over and flattens out. To give more linear release, but still offer a range of erosion rates, a class of polyanhydrides was prepared from aliphatic-aromatic dicarboxylic acid monomers, the alkanoic carboxyphenoxy alkanoic acids (see Fig. 2) [16]. The polycarboxyphenoxy acid monomers have an aromatic acid on one side and an aliphatic acid on the other, linked together by a non-labile bond. With the aliphatic-aromatic homopolymer, there can no longer exist regions of the polymer enriched in aromatic or aliphatic


321

Figure 5. Erosion of aliphatic-aromatic polyanhydrides. (0) p(CPA), (*) p(CPV), and (0) p(CPO). Deviceswere compression-molded1.4-cmdiameter x 1 mm thick disks. Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C. The cumulativeerosion was measured by UV absorbance at 235nm. Erosion rate can be controlled by the length of the aliphatic segmenton the monomer [16]. groups. As shown in Fig. 5, these polymers give linear release profiles, without rapid initial depletion of one monomer in favor of the other as seen for copolymers. Erosion rates were controlled by the length of the aliphatic chain on the alkanoic showed acid portion of the monomer. Release of a model compound p-nitroaniline excellent correlation between erosion of p(CPV) and drug release for So7oloaded compression molded disks into pH 7.4 phosphate buffer at 37°C (Fig. 6). Crosslinked

polyanhydrides

The development of an erodible, biocompatible properties could be useful for many applications,

material with good mechanical e.g., for use as a temporary bone

Figure 6. Polymer erosion and release of p-nitroanilinefrom p(CPV) homopolymer. (8) p-nitroaniline, (0) polymer erosion. Release from compression molded disks containing 5% (w/w) p-nitroaniline in phosphate buffer pH 7.4 at 37°C [16].


322

Figure 7. Release of p-nitroaniline from p(FA-SA) 50:50 copolymer. Devices were compressionmolded 1.4-cm diameter x 1.5 mm thick disks loaded with 507op-nitroaniline. Release was into 0.1 M pH 7.4 phosphate buffer at 37°C. The cumulative erosion is plotted as the appearance of the FA monomer as measured by UV absorbance at 245nm. Release of p-nitroaniline was followed by UV absorbance at 380nm [17]. discussed in the previous sections are generally replacement. The polyanhydrides brittle, lacking the compressive mechanical strength necessary for load-bearing that could be applications. Accordingly, it was sought to develop polyanhydrides crosslinked and thus perhaps improved in mechanical strength. As a prototype, we studied the synthesis of polyanhydrides of fumaric acid with aliphatic diacids [17]. The fumaric acid has a double bond which could be used for attachment of side chains or for crosslinking. The erosion profile and release of from a copolymer of p(FA-SA) 50 : 50 into pH 7.4 buffer at 37°C is p-nitroaniline shown in Fig. 7. As with other copolymer systems, changing the composition of the copolymer changes the erosion rates. The hydrophilic p(FA) homopolymer degrades by rapidly, completely disappearing in 2 days. Increasing polymer hydrophobicity of increasing the SA content increases the erosion period, until the homopolymer p(SA) erodes over a week or two. Infrared and NMR analysis confirmed the existence that the double bonds remained of double bonds in the polymer, demonstrating intact through the synthesis. The potential for crosslinking was demonstrated. Chemical crosslinking was in both in bulk and For was conducted solution. solution crosslinking, p(FA-SA) was with dissolved in tetrahydrofuran. or added, along Styrene methyl methacrylate as an accelerator. benzoyl peroxide or 2-butanone peroxide, and dimethyltoluidine as solution and were also tested Alternatively, azobisbutyronitrile divinylbenzene Bulk was crosslinking agents. crosslinking accomplished by mixing p(FA-SA) 50 : 50 copolymer with equimolar amounts of either styrene or methyl methacrylate, accompanied with the addition of either benzoyl peroxide or 2-butanone peroxide, either and also accompanied with the addition of an accelerator, dimethyltolidine, with or without cobalt naphtonate. The crosslinked polymer formed by these methods was insoluble in common solvents. DSC analysis was used to verify the

323 crosslinking. The lack of a melting point of the insoluble material was interpreted to be evidence of crosslinking of the polymer. These preliminary studies exhibited difficulties in yield and molecular weight of the crosslinked material. Nonetheless, these studies demonstrate that it may be possible to develop a new class of polyanhydrides with better mechanical strength.


Poly(anhydride-co-imide) The polyanhydrides in the previous sections do not have good fiber-forming abilities, and so would not be good candidates for sutures or other applications requiring tensile strength. However, this disadvantage may be eliminated by integration of imide bonds, which have excellent mechanical and thermal qualities [18]. Amino acids were converted into dicarboxylic acids by derivatization of their amino terminus with trimellitic anhydride. The resulting dicarboxylic acid can be converted into the mixed anhydride prepolymer by reflux in acetic anhydride. The amino acid prepolymer can be mixed with conventional prepolymers of sebacic acid and sub(see Fig. 8). This technique allows the integration of jected to melt-polycondensation essentially any amino acid into the backbone, and creates a class of polymers which

Figure 8. Reaction scheme for synthesisof poly(anhydride-co-imide)polymers [18].

324 can degrade at the anhydride linkages and which have improved mechanical from the imide linkages.

strength


FORMULATIONMETHODS This section outlines several approaches to formulating a polymer-drug mixture into implantable delivery devices. The first, compression molding, uses high pressure to form the devices; the second, melt casting and injection molding, uses high temperature ; the third, solvent casting, uses a solvent. The choice of method depends on properties of the drug, the polymer, and the desired drug release profile. Drug properties which impact the choice of fabrication method include hydrophobicity, diffusivity, stability, and tendency to interact with the functional groups on the polymer. Polymer properties, such as melting point, crystallinity, or brittleness, dictate the conditions for practical and reproducible fabrication of the device. Sometimes injection, rather than implantation, is the more desirable mode of introduction of the controlled release formulation to the body. Accordingly, there has been a major directed effort towards production of microspheres of drug-loaded of the polyanhydrides is polymer. A review devoted to microsphere preparation given in Mathiowitz and Langer [19]. These microspheres can be suspended in solution and injected into the body and thus obviate the need for surgical implantation. Microsphere systems impose additional challenges compared to flat, wafer-shaped implants. Because the microspheres can be small, less than 50,um in diameter, controlling the particle size of the drug is vital in order to achieve desirable release kinetics. If the drug particle size becomes comparable to the microsphere size, some particles will contain essentially only drug, while other particles will contain only the high polymer. This would not be conducive to steady release. Additionally, surface area to volume ratio of small spheres promotes rapid release of the drug. The diameter of the microspheres used in release must be carefully characterized to obtain consistent and meaningful release kinetics. This section outlines several methods of microsphere solvent removal, and spray preparation-hot-melt, have been used for polyanhydrides. drying-that Drug incorporation For all formulation procedures, the drug can be integrated into the polymeric device in two different ways [20]. One method involves trituration of the drug into a powder followed by physical mixing of the drug and polymer powders. This produces a device with small particles of drug distributed throughout the device. Alternatively, for drugs which are soluble in methylene chloride or chloroform, the polymer and drug can be codissolved in the solvent. The solvent is evaporated, and the result is a solid solution of drug and polymer. The cosolution method gives a device with intimate, homogeneous mixing between the polymer and the drug. The trituration method is more universally applicable, requiring only that the drug be formed into small particles. However, it may be difficult to fabricate devices with reproducible drug release profiles by this method. Sieving of the drug to an appropriate size range is critical for reproducibility. Additionally, the burst effect caused by rapid dissolution of the particles of drug on the surface of the device can be problematic for some applications. The cosolution method generally gives excellent reproducibility because of the homogeneity of the polymer-drug mixture.


325

Figure9. Comparisonof the trituration method to the cosolutionmethod for the incorporation of BCNU into p(CPP-SA) 20: 80 compressionmolded disks. Releasewas performed in pH 7.4 phosphate buffer at 37°C [20]. However, because it requires that the drug be soluble in chloroform or methylene method. Figure 9 chloride, it is not as universally applicable as the trituration vs. compares release of BCNU which has been incorporated by the trituration cosolution method for compression molded devices. Release from the cosolution method is slower and more controlled, with less of an initial burst. Compression

molding

Compression molding is a simple and flexible method of fabrication. The polymerdrug mixture is ground or spray dried into a fine powder, placed in a piston-type mold, and compressed into a flat wafer with a hydraulic press. Compression of the molding is usually done 5-10°C above the glass transition temperature polymer, which allows low temperature fabrication of devices of certain polymer 20 : 80. Typical pressures are 30000 psi. Comcompositions, such as p(CPP-SA) pression molding has the advantage of wide applicability to a variety of compounds. Because it is possible to fabricate devices at room temperature for some polymers, include it may alleviate problems of drug-polymer interaction. Disadvantages limitation to producing flat devices with the mold geometry, and, depending on the in drug which can result in poor reproducibility drug, poor matrix uniformity, release profiles. Nonetheless, overall, compression molding has been the most popular formulation method for polyanhydrides. Melt molding Heating the polymers above their melting temperature gives a viscous liquid, which is easily formable into the desired geometry. Injection molding or, alternatively, simply molding the melt under low pressure in conventional molds, gives a dense

326 and uniform polymeric matrix. There is generally good reproducibility from devices made by melt methods. The disadvantage is that melting the polymer requires in the range of 80°C for low-melting polyapplication of elevated temperatures, such 20: as p(CPP-SA) 80, as high as 150-200°C for the higher melting anhydrides as 50 : 50. The elevated temperatures such p(CPP-SA) greatly polyanhydrides increases the probability of interaction of the drug with the polymer (see 'Drugpolymer interactions' below). It may result in deactivation of heat sensitive drugs within the device. e.g., high temperatures can cause protein denaturation


Solvent casting Most of the polyanhydrides are soluble in chloroform and methylene chloride. This gives the opportunity to solvent cast films of a polymer-drug mixture. In solvent casting, the polymer is dissolved at approximately 10% (w/v) in the solvent. Solvent soluble drugs are codissolved with the polymer. Solvent insoluble drugs must be added as a fine powder. The solution is poured into a flat, open mold, which is chilled over dry ice. The solvent is allowed to evaporate slowly from the device, usually at -20°C, producing a thin, flat film. Solvent casting has the advantage that it requires no introduction of heat to the are sufficiently soluble to allow room temperature system. Most polyanhydrides dissolution, which is advantageous for heat sensitive drugs. However, some drugs, such as proteins, may undergo deactivation in a dry solvent medium. A major disadvantage is that solvent casting can be difficult to control, and often results in fragile and porous, nonuniform films. Additionally, there is the potential that the drug particles will settle to the bottom of the solution, producing films with more drug on one side of the device than on the other. Hot-melt

microencapsulation

The hot-melt microencapsulation process to produce microspheres is analogous to the melt molding process to form flat devices. In the hot-melt microencapsulation such process, the drug and polymer are suspended in a polymer immiscible-solvent, a silicone or olive oil. The mixture is heated to 5 ° C above the melting point of the polymer and stirred continuously to form a suspension of the polymer in the oil. The liquid is cooled until the polymer solidifies into microspheres. The microspheres are then washed with petroleum either, dried, sieved, and stored under nitrogen at - 20°C as a dry powder. Hot melt microencapsulation has been used to incorporate model hydrophilic, and protein compounds-acid and insulin, hydrophobic, orange, p-nitroaniline, 21:79 microspheres respectively-into p(CPP-SA) [21]. The method produced dense spheres, with smooth, even surfaces. Reproducible size distribution of the microspheres was achieved by adjustment of the stirring speed. Comparison of polymer erosion for different microcapsule particle sizes (Fig. 10) demonstrates that smaller particles eroded more rapidly, as would be expected from the higher surface area to volume ratio. The hydrophilic drug model, acid orange, was incorporated as sieved particles of less than 50,um. Release of the acid orange was more rapid for smaller microsphere size as seen in Fig. 11. It can also be seen from Fig. 11 that the release profiles of the dye from p(CPP-SA) 20:80 exhibited a burst which increased with size of the effect, microspheres decreasing


327

Figure 10. Effect of microspheresizeon the erosion of p(CPP-SA) 20: 80 microspheresprepared by the hot-melt microencapsulation method. Numbers indicate the size of the microspheres in microns. Microsphereswereloaded with 5 Voacid orange. Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C. Releasemeasured by UV absorbance at 247 nm [21].

Figure 11. Effect of microspheresize on the releaseof acid orange from 5%-loaded p(CPP-SA) 20: 80 microspheresprepared by the hot melt microencapsulationmethod. Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C [21]. microspheres. This is expected because as the size of the particle decreases, the relative size of the drug particle to microsphere size increases and increases the relative burst effect. Release of acid orange preceded the polymer erosion as measured by CPP appearance in the buffer. However, the correlation between drug release and CPP appearance in the buffer is closer for low loadings of the drug (2%) than for high loadings (27%) as seen in Fig. 12. It is hypothesized that at high drug loadings, the drug itself can provide pores and channels for release of the drug, whereas at low loadings, the release is controlled more by the polymer erosion.


328

Figure12. Effect of loadingon the releaseof acid orange from p(CPP-SA) 20 : 80microspheresprepared by the hot melt microencapsulationmethod. Microspheresize was 800-1000,ummicrospheres. Release was into 0.1M pH 7.4 phosphatebuffer at 37°C. Polymer erosion was measuredby total weightloss [21]. Release of the hydrophobic p-nitroaniline demonstrated close agreement between compound release and polymer erosion as measured by CPP appearance (Fig. 13). The hydrophobic character of the drug, the initimate mixing of drug and polymer, and low porosity of the particles, resulted in no burst of the model compound. of insulin as one ,11m particles showed release of the insulin over an Incorporation 8-day period, correlating reasonably well with polymer erosion as measured by CPP appearance

(Fig. 14).

Solvent removal microencapsulation Some drugs are heat sensitive and are therefore not amenable to incorporation into microspheres by the hot-melt encapsulation method. A common room temperature technology for microsphere production that has been used for other biodegradable However, this method employs an aqueous polymers is solvent evaporation. suspension, which is unsuitable for polyanhydrides because of their high hydrolytic activity. Initial trials using this method indicated that the polymer was significantly degraded by the encapsulation process [22]. An adaptation of this method that is suitable for the hydrolytically labile polyanhydrides has been developed. In this known as solvent the is method, removal, drug dispersed or dissolved in a polymer solution of a volatile organic solvent, usually methylene chloride or chloroform. The mixture is suspended in an organic oil which is immiscible with the volatile organic solvent such as silicon oil. One to 5% of a surfactant such as Span 85 is added and the mixture is stirred. Petroleum ether is introduced and stirring continues. The organic solvent is extracted into the oil, creating microspheres. The microspheres are under dry filtered, washed with petroleum ether, dried and stored at - 20°C nitrogen.


329

Figure 13. Release of p-nitroaniline from microspheresprepared by the hot melt microencapsulation method. (a) 50-104, (b) 425-500, and (c) 850-1000,um.Releasewas into 0.1M pH 7.4 phosphate buffer at 37°C. Polymer erosion was measured by weight loss [21]. This method was used to make microspheres of p(CPP-SA) 20: 80 and 50: 50 and p(CPP-DD) 20: 80 and 50: 50 [23-25]. Molecular weight was found to affect the ability to make good, reproducible microspheres, with low-molecular-weight easier to Cross sections of the polymers being process. microspheres revealed a porous structure with increased density toward the external surface of the sphere. These microspheres lost their physical integrity during erosion studies and occasionally


330

Figure 14. Release of insulin from p(CPP-SA) 20: 80 microspheresprepared by the hot-melt microencapsulationmethod. Microspheresize 850-1150,um,15% insulin loading. Polymer erosion measured by weight loss [21 ] . left empty shells. Crystalline polymers, such as p(SA) did not produce smooth 50 : 50 produced microspheres, while amorphous polymers, such as p(CPP-SA) microspheres with smooth external surfaces. Spray drying microencapsulation Spray drying is a reproducible, rapid, and easy to scale up method for preparing microspheres. In this method, the polymer is dissolved in a solvent such as chloroform or methylene chloride along with the drug, either in a dissolved or dispersed form. The solution is sprayed through an atomizer. As the particles fall toward the bottom of the spray dryer, they are simultaneous dried by an upward flow of nitrogen. 20 : 80 and 50 : 50, p(FA-SA) Spray drying was tested for p(SA), p(CPP-SA) 50 : 50, and p(CPV) [25, 26]. The more amorphous 20 : 80 and 50 : 50, p(CPH-SA) 50 : 50, p(CPH-SA) 50 : 50 and p(CPH)-were not amenpolymers-p(CPP-SA) able to spray drying, and produced aggregates and uneven particle morphologies. It was proposed that the low glass transition temperatures of the polymers allowed them to fuse together during spray drying. The other polymers which were tested were well-suited to spray drying, forming uniform microspheres 1-5,um in diameter. However, cross sectioning revealed that microspheres of p(FA-SA) copolymers displayed highly porous structures. The morphologies of the other polymers varied from dense to porous, and smooth to rough, with the type of drug incorporated. PREPARATION AND PHYSICAL PROPERTIES Preparation of methods of synthesis for the polyanhydrides have been developed, A review of the both and solution methods. including melt-polycondensation methods of synthesis of polyanhydrides is given by Leong [27]. The solution A number


331 methods include dehydrative coupling using bis(2-oxo-3-oxazolidinyl)phosphinic chloride and phenyl N-phenylphosphoroamidochloridate [28], dehydrochlorination involving reaction of acyl chloride with carboxylic acids [28], and a one-step polymerization using phosgene or diphosgene coupling agents [29]. The most commonly used method to produce polyanhydrides is melt polycondensation [30]. The melt polycondensation method has the advantages that it can produce relatively highmolecular-weight polymers and does not require solvents during polymerization, whereas the solution methods yield low-molecular-weight polymers and introduce a solvent impurity to the product. Melt polycondensation involves the a series of steps. In the first step, the dicarunits monomer are boxylic prepared and purified. Prepolymers are synthesized by reflux of the diacid with acetic anhydride for several hours. This forms the mixed anhydride prepolymer. The crude prepolymer is recrystallized from dry toluene, and then immersed in a 1 : mixture of dry petroleum ether and diethyl ether to extract traces of acetic anhydride and toluene. This prepolymer purification step has been found to be critical to further synthesis into high-molecular-weight polymers. The purified prepolymer is placed into a temperature-controlled vessel under high vacuum, where polymerization takes place. The acetic anhydride, which is produced by the polymerization reaction, is removed by vacuum. The optimum temperature for polymerization was determined to be 180°C. After polymerization is complete, the polymer is allowed to solidify at room temperature and is transferred to dry bottles and stored under dry nitrogen at -20°C. The average molecular weight of the polymer can be controlled to some extent by the length of time of synthesis, and the maximum attainable molecular weight is affected by the polymer composition [30]. Molecular weight increased as the polymerization reaction time increased from 10 to 90 min in the melt polycondensation synthesis. However, after 45 min, the molecular weight increase was largely due to a small amount of a high molecular weight (M = 1000 000) fraction, with the bulk of the material still in the 50 000 molecular weight range. Thus increased reaction time created increased polydispersity. Different compositions were found to influence the molecular weight attained under identical reaction conditions in melt polycondensation. Generally, copolymers with high sebacic acid content were easy to polymerize to high molecular weights, M,, 140 000. As CPP content increased, the molecular weight decreased to 40000 for p(CPP-SA) 50: 50. It is The crude polymer must be purified before use in in vivo applications. believed that residual acetic anhydride and other impurities can cause inflammatory reactions. To purify the polymer, the polymer is dissolved in methylene chloride or chloroform (10% w/v). Petroleum ether or hexane is added to the solution, which causes the polymer to precipitate. The precipitate is washed three times with diethyl ether to remove residual acetic anhydride. The purified polymer is dryed under vacuum and transferred to dry bottles under nitrogen for storage at -20°C. Stability in the product shelf life and in the proStability is an important consideration The of stability polyanhydrides of SA, CPM, and CPH, cessability of the polymer. acid (PDP) in both the solid and solution state has and 1,4-phenylenedipropionic been studied [31]. Upon storage under dry argon or vacuum at 21 ° C, the aromatic


332 their original and p(CPH) homopolymers-maintained polyanhydrides-p(CPM) molecular weight for the study period of 12 months. The aliphatic polyanhydridesa decrease in M, from 137 500 to 78 000 for p(SA) and p(SA) and p(PDP)-showed 43 000 to 23 000 for p(PDP) homopolymers after storage under dry conditions at 21° C for 12 months. Temperature was found to affect the stability under dry conditions. The homopolymer p(SA) of an original MWof 137 000 decreased in molecular weight to 65 800, 78 000, 96 600 and 118 400 at 37, 21, 9, and -10°C, respectively after 60 in chloroform solution was of polyanhydrides days of storage. Depolymerization 20 : 80 and 50 : S0, p(CPH), p(CPM), and p(PDP) also studied for p(CPP-SA) under dry nitrogen at 37°C. The aromatic homopolymers p(CPH) and p(CPM) showed no decrease in intrinsic viscosity or molecular weight over a 3-day period, while the copolymers with high aliphatic content showed fairly sharp decreases in intrinsic viscosity, indicating molecular weight loss. affected the high-molecularGPC analysis revealed that the depolymerization weight fraction of the polymers to a greater extent. Thus the Mn of the polymers is less affected than The loss of molecular weight tends to be rapid initially, but eventually slows to an apparently stable MW plateau. Thus, one would infer that if a high-molecular-weight and low-molecular-weight polymer of the same composition the molecular weight of the higherwere stored under the same conditions, molecular-weight polymer would decrease more rapidly, and eventually, the molecular weights of the two polymers might stabilize to similar values. The authors via intra- or intermolecular anhydride interchange proposed a self-depolymerization resulting in ring formation as the mechanism in both the solution and dry condition NMR data and experiments involving addition of radiolabelled depolymerization. water appeared to support this hypothesis [31, 32]. From these studies, it can be concluded that the storage stability of the polyanhydrides tested is excellent at -20°C if the polymers are kept under vacuum under dry argon or nitrogen. The polymers can be maintained for many months under these conditions with little loss in molecular weight. It should be noted that significant molecular weight loss will occur in the polymers within days or weeks if not environment. Therefore, careful storage of both the kept in a dry, low-temperature raw polymer and finished devices is critical to ensure the integrity of the devices. As a practical consideration, it is recommended that the surface area of the polymer be minimized during storage. It has been observed that polyanhydride powders tend to lose molecular weight, even under storage at dry conditions at -20°C. Therefore, processing steps, such as purification, grinding, or spray drying should be performed immediately before the polymer is fabricated into devices. From the stability studies in solution, it can be concluded that the solution stability of polyanhydrides is sufficient to allow solution processing, so long as the process can be completed in several hours and does not extend to many days. Molecular

weight

Molecular weight of the polymer may affect its hydrophobicity, crystallinity, and other properties, which in turn affect the erosion and drug release rates. Most have been determined by molecular weight measurements for the polyanhydrides gel permeation chromatography (GPC) relative to polystyrene standards. More recently, Mark-Houwink parameters have been determined for the most commonly

333 used copolymer, p(CPP-SA) 20: 80, in chloroform [33]. These parameters are used in a universal calibration curve which can be used to determine the molecular weight of the polymers by GPC. The universal calibration curve is given by the expression


where h is the retention volume and [1]]is the intrinsic viscosity. The Mark-Houwink parameters were calculated from the relationship

where MW was approximated to be close to the viscosity average molecular weight, in the Mark-Houwink equation. Viscosity versus MW data were acquired for 20: 80 in chloroform solution in the range of M, of 245 000-14000. p(CPP-SA) The intrinsic viscosities were in the range of 1.25-0.22 dl/g. The Mark-Houwink parameters calculated from the data were K = 3.88 x 107 and a = 0.658. The effect of molecular weight on device fabrication is currently under study. Molecular weight has been found to influence fabrication methods, particularly microsphere preparation. High-molecular-weight polymers showed a tendency to form rods rather than spheres, although this is dependent upon details of the processing methods [23]. It has been observed on occasion that some high-molecularweight copolymers of p(CPP-SA) (M, 100000) seem to have lower solubility in chloroform, and are therefore more difficult to process. However, a detailed study of the effect of molecular weight on solubility has not been performed. The effect of molecular weight on polymer erosion rates has not yet been fully established. Preliminary studies suggest that the effect of molecular weight on the overall erosion and drug release is minimal. A higher molecular weight polymer will exhibit a slightly longer lag period at the beginning of the erosion process, but is not a primary variable in affecting the subsequent erosion rates. Crystallinity

and thermal properties

Differential scanning calorimetry (DSC) has been used to characterize the thermal properties of several polyanhydride [34]. Glass transition (Tg) and compositions and heats of fusion for copolymers of p(CPP-SA), melting (Tm) temperatures and p(FA-SA) are given in Table 1. Tg and Tm are important p(CPH-SA), parameters because they dictate which device fabrication methods are practical for a given system. Tg determines the minimum temperature necessary for compression molding, and Tm determines the minimum temperature necessary for injection molding or melt pressing. It has been hypothesized that the crystallinity of the polymer will affect erosion rates. Crystalline regions are expected to erode more slowly than amorphous regions, and the type of crystals that form the crystalline regions may also affect the erosion rate. Additionally, crystallinity of the polymer has been found to affect device morphology. For example, the use of p(SA), a highly crystalline polymer, to make microspheres by both the solvent removal and spray drying methods resulted in microspheres with crenulated and irregular external surfaces. Use of amorphous 50 : 50, yielded microspheres with smooth external polymers, such as p(CPP-SA) surfaces.

334


Table 1. Thermal properties of p(CPh-SA), p(FA-SA), and p(CPP-SA)

of p(CPP-SA), of various copolymer compositions crystallinity studied a combination of DSC and X-ray and has been by p(CPH-SA), p(FA-SA) diffraction [34]. Fig. 15 shows the effect of copolymer composition on the percent crystallinity of various copolymers. As compositions of copolymers shifted towards a high content of one of the monomers, polymer crystallinity was generally high, 6070%, except for the case of high CPH content, where crystallinity was around 20% for the homopolymer. As the composition shifted toward equimolar content of the monomers, crystallinity decreased, with only 5-1007o crystallinity for p(CPP-SA) 50 : 50, and 35% crystallinity for p(CPH-SA) 50 : 50. Crystallinity of the p(FA-SA) copolymers was high, greater than 400%, for all compositions. The crystallinity of p(FA-SA) copolymers exhibited a minimum at equimolar ratios of SA to FA and maxima at extremes of composition corresponding to homopolymers of FA or SA. The


335

Figure 15. Percent crystallinityof p(CPP-SA), p(CPH-SA) and p(FA-SA) copolymers [34].

X-ray powder diffraction patterns of the homopolymer p(SA) revealed a distinct structure characterized by four peaks at 4.41, 4.10, 3.72, and 3.42 A, corresponding to calculated crystallite sizes of 6.8, 10.4, 13.6, and 3.7 A as calculated from the Scherrer and Bragg approximation. Powder diffraction of the homopolymer p(CPP) displayed a well-defined pattern at 4.64, 4.15, 4.0, 3.58 and 3.24 A. Crystallite sizes of 7.9 and 9.4 A were calculated for the first two diffractions; the other peaks were too wide to determine crystallite size. The homopolymer p(FA) showed diffractions at 4.09, 3.70, and 3.14 A with crystallite sizes of 10.37, 9.10, and 20.47 A, 0respectively. The homopolymer of p(CPH) did not show characteristic crystalline diffraction patterns. Unlike the other homopolymers, p(CPH) is highly amorphous. Powder diffractions of copolymers p(CPP-SA), and p(FA-SA) p(CPH-SA), of the displayed the structures of crystallites characteristic of the homopolymer monomer at the higher concentration. Fig. 16 shows the crystalline pattern of the p(CPP-SA) copolymer system at various ratios of CPP to SA. For the p(CPP-SA) of SA, the pattern looks similar to that of p(SA). The system, at high ratios crystalline pattern disappeared at molar ratios between 30: 70 and 70: 30. For the p(CPH) system, the structure of p(SA) is seen in copolymers up to a 30 : 70 molar a broad halo appears, on top ratio of CPH to SA. At higher CPH concentrations, 20: 80 displayed of which some fine diffraction persists. The pattern of p(FA-SA) the crystalline pattern of p(SA), while p(FA-SA) 90: 10 displayed the pattern of p(FA). Intermediate ratios of FA to SA resulted in new diffractions at 4.28, 3.95, and 3.6 A, indicating new crystal structures in the equimolar mixture. The method of device fabrication was found to affect the crystallinity of the resultant device. Formulation methods which involved solvents, such as microsphere production by solvent removal [23] or by spray drying [26], resulted in devices with lower crystallinity than the original polymer. The proposed explanation is that the rapid solidification of the device gave inadequate time for crystal formation in the polymer.


336

Figure 16. X-ray crystal patterns for the p(CPP-SA) copolymer system [34]. Spectroscopy Infrared spectroscopy readily confirms the identity of the polymer as a polyanof a hydride. A doublet occurring between 1670 and 1800 cm-1 is characteristic can also be NMR observed carboxylic anhydride [15]. Polyanhydride linkages by spectroscopy. In addition to confirming the structure of the polymer, NMR spectroscopy analysis allows the determination of the comonomer sequence distribution of polyanhydride copolymers [33]. NMR studies can be used to determine if the polymer is blocklike (i.e. -A-A-A-A-A-B-B-B-B-B-), alternating (i.e. -A-B-A-B-A-B-A-B-), random (i.e. probability that A or B is next to a given monomer is equal to its mole fraction in the polymer) or some combination of these. Proton NMR spectra of in deuterated chloroform reveals two doublets at 8.1 and 8.0 ppm p(CPP-SA)


337 (J = 8.7 Hz), and two triplets at 2.6 and 2.4 ppm (J = 7.4 Hz). The homopolymer p(SA) has only one triplet at 2.4 ppm, and the homopolymer p(CPP) has only one doublet at 8.1 ppm. Thus the downfield doublets at 8.1 and 8.0 ppm correspond to the diads CPP-CPP and CPP-SA respectively, and the upfield triplets at 2.6 and 2.4 ppm correspond to the diads SA-CPP and SA-SA, respectively. Integration of the areas under the appropriate peaks for copolymers of different compositions CPP-SA and CPP-CPP provides the relative fraction of SA-SA, (or SA-CPP), diads in the polymer. If the reactivity ratio of the monomers is unity, that is, if the probability that SA reacts with SA is the same as the probability that CPP reacts with SA and vice versa, then the expected diad ratios can be easily calculated. For example, a random X : Y, where X is the mole fraction CPP copolymer of composition p(CPP-SA) and Y is the mole fraction SA, would show the following behavior.

where f (A-B) for a random

is the fraction of A-B diads. Furthermore, copolymer is given by average block length SA sequences average block length CPP sequences

the expected block length

= X-1 = = Y-1.

(1 -

y)-l

(6) (7)

of p(CPP-SA), with Y ranging from 0.49 to 0.96, and Compositions with Y ranging from 0.12 to 0.83, were tested. Examination of the p(CPH-SA), effect of melt polymerization reaction time, 20, 40 and 90 min, did not reveal any effect of reaction time on the conversion or the statistical fractions of the different types of bonds in the polymer, only on the molecular weight of the polymer. Figure 17 shows a comparison of the experimental diad distribution and that expected for

Figure 17. Comparisonof copolymersequencedistribution with that predicted for a random copolymer for p(CPH-SA) and p(CPP-SA). For p(CPH-SA): (8) SA-SA, (0) SA-CPH, (A) CPH-CPH. For p(CPP-SA): (0) SA-SA, (0) SA-CPP, (A) CPP-CPP.


338

Figure 18. Comparison of block length as measured by NMR to that predicted by a random copolymer distribution for p(CPH-SA) and p(CPP-SA). For p(CPH-SA): (*) SA-SA, (s) CPH-CPH. For p(CPP-SA): (0) SA-SA, (D) CPP-CPP. a random copolymer system from Eqs (3) to (5). Comparison of the experimental probabilities indicate agreement between expected and experimental probabilities and p(CPH-SA) are well within experimental error, confirming that p(CPP-SA) random copolymers, at least at the range of ratios studied. A comparison of experimental and theoretical average block lengths for these copolymers is shown in Fig. 18. The sequence distribution of monomers in the copolymer can help in understanding several effects. First, the segment length affects how crystalline the polymer is likely to be. The long block lengths of a particular monomer that are expected for a high fraction of the monomer in the copolymer suggest a more crystalline structure a hypothesis that was verified experimentally. at the extremes of composition, distribution can help explain erosion behavior. If the difthe Secondly, sequence ferent types of bonds have different reactivities, then the appearance of monomers relative to each other would be affected. This is seen in the more rapid erosion of SA and than CPP in the copolymer device seen in Fig. 3 that occurs as the SA-CPP bonds behind (see 'Comparison SA-SA bonds are broken, leaving the CPP-CPP of Release of Comonomers' below). Mucoadhesive

properties

20 : 80 microspheres with mucin gels has been The interaction of p(CPP-SA) characterized by quantification of the mucoadhesive forces between polymer and gel [35]. The study was done to examine the potential of polyanhydride devices which could be delivered orally and attach to the intestinal wall. The adhesive interactions between polymer microspheres and mucus relate to the retention of oral delivery microtract. It was found that larger p(CPP-SA) systems in the gastrointestinal spheres exhibit stronger mucoadhesive forces. Mucoadhesive forces of the order of 120,uN were developed on microparticles of diameter larger than 900,um in contact with the intestinal mucosa of Sprague-Dawley rats. Although more work is required to assess the mucoadhesive properties of the polyanhydrides, these studies provide a preliminary basis for evaluation of the potential of orally delivered polyanhydrides.


339

DSC and X-ray diffraction have been used to study the effect of the incorporation of drug into the polymer for several formulation methods and substances. In general, it would be expected that hydrophilic compounds, which are not compatible with the polymer, would be dispersed as particles throughout the polymer. For these dispersed systems, X-ray diffraction studies should show the characteristic patterns of the compound, and the DSC should exhibit two endotherms, one for the melting of the drug and one for the melting of the polymer crystalline regions. Hydrophobic compounds potentially can form a solid solution with the polymer. For a solid solution, DSC measurements would be expected to display a single endotherm for the polymer with melting point depression, and X-ray diffraction would not display a distinct pattern from the drug. Studies on several systems have confirmed this hypothesis. X-ray diffraction studies indicated that acid orange, a hydrophilic dye, is integrated as crystals and spray drying the polymer during solvent evaporation dispersed throughout [23, 26]. On the other hand the same formulation methods used with methyl red, a hydrophobic dye, gives no crystalline diffraction nor any endotherm related to mixture decreased, the melting of the dye. The melting point of the polymer-drug solution had been formed. Incorporation further suggesting that a drug-polymer of p-nitroaniline by the solvent removal method gave a product with characteristics of BCNU (1,3The incorporation of both dispersed and dissolved p-nitroaniline. a hydrophobic chemotherapy agent, by the bi s [2-chloroethyl] -I -nitro sourea), trituration and the cosolution method has been compared [36]. Compression method, molding of particles of BCNU with polymer powder, the trituration produced a wafer with two endotherms, one corresponding to that of BCNU and to that of the polymer. Thus a dispersed device had been one corresponding formed. On the other hand, incorporation of BCNU into the polymer by cosolution of drug and polymer in methylene chloride, followed by spray drying and compression molding, gave a single endotherm on the DSC, at a lower melting temperature than that of pure polymer. This suggests formation of a solid solution of drug and polymer. Amine, alcohol, and thiol groups have potential to form covalent bonds with the anhydride bond, which would be deleterious to drug activity, drug release, and polymer integrity. A study was performed to examine covalent drug-polymer interactions for drugs with amine groups, which may form amide bonds with the 20: 80 with different para substitute anhydride bond. Injection molding p(CPP-SA) that interaction of the drug with the polymer may anilines at 10% loading showed indeed be problematic for some compounds [37]. Although IR spectroscopy of showed no apparent amide 20: 80 withp-nitroaniline injection molded p(CPP-SA) or p-phenylformation, inj ection-molded polymer with p-bromoaniline, p-anisidine, enediamine, showed amide bands of moderate intensity, indicating an undesired interaction. covalent drug-polymer However, applying the same spectroscopic molded to samples of the polyanhydrides with the same drugs compression analysis did not produce any signs of the covalent reaction. Therefore, temperature seems to interactions. It may be possible to prevent play a critical factor in drug-polymer methods which do not undesired covalent interactions by choosing formulation elevated require temperatures.

340 IN VITROCHARACTERIZATIONOF EROSION Morphological

changes


Most biodegradable polymer systems undergo significant bulk erosion. It has been shows a uniform pore structure, reported that eroded poly(lactide) indicating erosion has occurred almost uniformly from throughout the polymer [38]. The are a sharp contrast to this. Photomicroscopy of device cross polyanhydrides sections has revealed the existence of what is termed an 'erosion zone' in the polymer. Erosion occurs primarily in a relatively sharp, slowly advancing front from the exterior to the interior of the device [21]. Thus drug release is expected to occur primarily from the exterior to the interior of the polymer, similar to the expected behavior of an ideal surface eroding device.

Figure 19. Effect of pH on the erosion of p(CPP) homopolymer. Devices were compression-molded 1.4 cmdiameter x I mm thick disks. Release was into 0.1M phosphate buffer at various pH values at 37°C [15]. Erosion profiles The effect of polymer composition on the erosion profiles of the polyanhydrides has been discussed in previous sections. Erosion rates depend on polymer composition, formulation method, and device size and geometry. More hydrophobic polymers erode more slowly as was seen in Fig. 4 for copolymers of p(CPP-SA) in a series of increasing hydrophobicity as CPP content increased and in Fig. 5 for homopolymers p(CPA), p(CPV), and p(CPO) as the hydrophobicity increased with backbone chain length. Smaller microspheres erode more quickly, because of increased surface area to volume ratio as was seen in Fig. 15 for microspheres of p(CPP-SA) 20: 80 produced by hot-melt microencapsulation.

341 Additionally, the conditions of the release medium can affect erosion rates. The erosion rates of polyanhydrides are pH dependent [15]. Figure 19 shows the erosion rates of p(CPP) from compression molded disks into 0.1 M phosphate buffer at pH 7.4, 8.0, 9.0 and 10.0. Erosion was more rapid in more basic solutions. At pH 1, less than one percent of the p(CPP) eroded in 5 weeks. Elevated temperature of the release medium results in faster erosion of the polymer. Studies done at 37 and 60°C for the erosion of p(CPM) indicate much faster erosion at the higher temperature [14].


Comparison

of release of comonomers

With a few exceptions, most work on biodegradable copolymer systems has studied only drug release or total mass loss of the system. To gain a more fundamental of polymer degradation, we have recently developed an high perunderstanding formance liquid chromatography measure the assay (HPLC) to simultaneously of both CPP and SA monomers in solution [39]. concentration 20 : 80 and 50:50 copolymer The results for the erosion from p(CPP-SA) devices are shown in Figs. 20 and 21. The sebacic acid is released from the device more rapidly than the CPP. The discrepancy is more pronounced for higher CPP contents. For p(CPP-SA) 50 : 50, all the SA was eroded within 4 days, while the device remains physically intact, with a significant fraction of the initial CPP content, for a period of several weeks. Spectroscopic studies strongly suggest that diads degrades more slowly than the the anhydride linkage between the CPP-CPP diads. Thus the SA and the CPP anhydride linkage between SA-SA and CPP-SA which is bonded to the SA erodes rapidly, leaving behind a shell containing only the oligomers. slowly degrading CPP-CPP-containing Drug release device Drug release is affected by a number of factors. Polymer composition, fabrication method, size and geometry of the device, particle size of incorporated

Figure 20. Comparison of the erosion of SA with CPP for p(CPP-SA) 20:80. Deviceswere compression-molded 1.4-cm diameter x 1.2 mm thick disks. Release was into pH 7.4 phosphate buffer at 37°C [39].


342

Figure 21. Comparison of the erosion of SA with CPP for p(CPP-SA) 50:50. Devices were compression-molded 1.4-cm diameter x 1.2 mm thick disks. Release was into pH 7.4 phosphate buffer at 37°C [39]. of drug, drug loading, and drug hydrophilicity all drug, method of incorporation affect the rate of drug release. Generally hydrophilic drugs, which are dispersed as particles within the devices, will display a burst effect. That is, drug on the surface of the device will dissolve very rapidly. The extent of the burst depends on the particle size of the drug compared to that of the device; the finer the drug particles, the smaller the burst effect. This was seen in Fig. 12 for the release of acid orange from microspheres of p(CPP-SA) 20 : 80. Drugs which are soluble in methylene chloride or chloroform can be incorporated as a solid solution with the polymer, and hence display little or no initial burst effect, as was seen in Fig. 13 for release of p-nitroaniline from p(CPP-SA) 20: 80 microspheres. The correlation between drug and polymer release is better for small drug particle sizes or drug-polymer solid solutions than for large particles of dispersed drug, as was seen in Fig. 9 in the comparison between the trituration and solution methods of drug incorporation. Generally, the hydrophilic drugs studied released more rapidly than hydrophobic drugs (compare the release of hydrophilic acid orange in Fig. 12 to hydrophobic p-nitroaniline in Fig. 13), although this may be due to differences in drug incorporation method and not necessary drug properties. CLINICAL DEVELOPMENTOF THE POLYANHYDRIDES:A CASE STUDY The application of polyanhydrides which has advanced the furthest clinically is for the treatment of human brain tumors. Glioblastoma multiforme is a type of brain cancer that afflicts 14000 patients in the US each year. Patients with advanced forms of the disease have a life expectancy of less than one year after diagnosis, despite surgical removal of the tumor, radiation therapy, and conventional systemic The tumor invariably grows back. Additionally, chemotherapy. systemic chemotherapy has side effects which greatly compromise the quality of life for the patient. The chemotherapy 1,3-bis[2-chloroethyl]- I -nitroagent used, BCNU (carmustine, 15 min; hence very little sourea), has an in vivo systemic half life of approximately In an experimental active drug contacts the tumor site with systemic administration.


343 therapy, BCNU is incorporated into the polymer, which is then implanted into the site of the excised tumor. Bioerodible controlled release devices have several advantages over systemic treatments of this type of disease. Implantation directly into the brain bypasses the blood brain barrier, thus minimizing systemic drug exposure and maximizing tumor exposure. The polymer protects the labile BCNU from degradation before it is released, allowing more active BCNU to contact the tumor site. The to the tumor site, where it is needed. device provides a high local concentration release occurs over an extended period of several days to several weeks, Finally, for the a greater opportunity drug to destroy remaining tumor cells. giving This section is a case study of the development of a polymeric drug delivery system from a laboratory experiment to human clinical trials. Development of this type of between workers in the fields of polymer therapy involves extensive collaboration chemical materials and science, chemistry engineering, biology, and medicine. A number of challenges must be addressed to bring the project to fruition. A system must be designed to release the BCNU at the desired rate for the desired period of time. Extensive in vitro and in vivo assays in animals must be performed to assess of the polymer treatment. The safety of the treatthe safety and biocompatibility ment to the patient group must be determined. The next sections will summarize the variety and nature of testing that is involved in the introduction of a new biomaterial to human trials. In vivo drug release One of the first goals of a controlled release treatment is to develop a system which would reproducibly release the drug over the desired period of time at the desired the optimum drug dosage was rate. Because of the novelty of the treatment, unknown. However, the 200 mg standard dose for systemic intravenous administration of BCNU provided an upper limit to the total drug release. Dosage could be controlled by the number of implanted wafers or the drug loading. It was believed that BCNU would show efficacy if released over a period of several days. Preliminary 20: 80 degraded in vitro over results with blank polymer indicated that p(CPP-SA) a 3-week period, and this composition was chosen for the clinical formulations. Once the polymer composition was selected, a suitable fabrication method needed to be established. Because of the instability of BCNU, methods which involve the application of heat, such as melt casting or injection molding, were deemed to be unsuitable for the application. Some early studies tested the trituration method of fabricating drug-loaded devices, in which ground drug and polymer are physically mixed together and then compressed into wafers. The resultant wafers showed an unsuitably large burst effect, excessively rapid release of the drug, and poor reproducibility between batches. The cosolution method, in which BCNU and polymer were codissolved into methylene chloride, produced devices with better drug release The common laboratory approach of milling of the profiles and reproducibility. for clinical scale promixture, however, proved to be inappropriate polymer-drug duction of wafers. Spray drying of the drug-polymer methylene chloride solution produces uniform spheres, which compress readily into the final wafer formulation. In the final dosage form for the first clinical trials, the appropriate percentage of BCNU and polymer were codissolved in methylene chloride, spray dried, and cominto 200 mg wafers of approximately pressed at 30 000 psi at room temperature 1.4 cm in diameter and 1.0 mm in thickness [36]. ..

344


In vitro toxicity studies In vitro testing of the polymer provides a relatively rapid assessment of potential deleterious effects that the polymer may present to a human subject. In vitro studies were performed to assess cytotoxicity, mutageneity, teratogeneity, and effect of polymer on cell growth [37]. In vitro testing of the cytotoxicity and mutageneity of the degradation products of the p(CPP-SA) 45 : 55 copolymer system was determined by a forward mutation assay in Salmonella typhimurium using 8-azaguanine resistance as a genetic marker, both with and without the addition of mammalian metabolism enzymes. The polymer did not induce mutation, and exhibited slight, but not significant, toxicity. Teratogenic potential was assessed by the attachment efficiency of ascitic mouse ovarian tumor cells to plastic surfaces coated with concanavalin A. Results indicated that the polymer degradation products could be In tissue culture studies, bovine aortic endothelial cells considered nonteratogenic. and smooth muscle cells were grown on the surface of p(CPP-SA) 45 : 55, 50 : 50, and p(TA). There was no evidence of toxic effects on these cells p(TA-SA) as measured by their morphology or ability to proliferate. These first tests were not exhaustive, but they covered the more important aspects of acute toxicity. The vivo biocompatibility studies. encouraging results led to the next step-in In vivo biocompatibility The initial in vivo biocompatibility into testing was performed by implantation rabbit eye corneas and subcutaneously into rats [37]. Polymer pellets of p(CPP) and p(TA) 50:50 were implanted into pouches within the rabbit corneal stroma to determine local tissue response. The host response to the polymer exhibited no observable inflammatory characteristics. The inertness is comparable to that of and ethylene vinyl acetate, both biocompatible poly(hydroxyethyl methacrylate) materials currently used in human implant therapies. Subcutaneous implantation of rats was also performed. The p(CPP) in the abdominal region of Sprague-Dawley rats were sacrificed 6 months after implantation and histology of the tissue surrounding the implant site was studied. After 6 months the implantation site was only cell infiltraslightly encapsulated by layers of fibroblastic cells. No imflammatory tion was seen in the tissues adjacent to the implant. With the encouraging results of the preliminary in vivo biocompatibility studies, a more detailed in vivo toxicity study was executed [40]. This investigated the in vivo 20: 80 systemic and local effects of implantation of high doses of the p(CPP-SA) copolymer, dosages much greater than the expected human levels. It was anticipated that the total amount of polymer that would be administered to a human patient would be less than 1500 mg of polyanhydride. Assuming the weight of an average human to be 70 kg, the dose level of the polymer would be less than 30 mg/kg. Dose levels of polymer of 800 and 2400 mg/kg, 25 and 80 times the anticipated human dose on the basis of weight polymer per body weight, were tested in Sprague-Dawley rats. These dose levels required implantation of 200-mg devices of 1.4 cm in diameter and 1.2 mm thick. Because the size of the polymer devices to achieve these high dose levels was so large compared to the size of the rat brain, the polymer was implanted subcutaneously. Twenty-four rats were randomly assigned to one of three groups: a control group, which received no implants, Group A which were implanted with one 200 mg


345 an 800 mg/kg dose), and Group B, which were polymer device (approximately with 200 three implanted mg polymer devices (approximately a 2400 mg/kg dose). was assessed by a number of methods. Thirty-six clinical chemistry Biocompatibility and hematology parameters were monitored throughout the study. Blood samples were taken at 0 weeks (start of study), 2, 3, 5, and 8 weeks (termination of study). Of the 25 blood clinical chemistry parameters studied, the implantation of polymer disks had no statistically significant effects on any plasma levels. Of the 11blood hematological effects studied, the implantation of one matrix had no statistically significant effects on any plasma levels. With three implants, 9 of the 11blood hematology were not significantly affected. Two parameters, and parameters neutrophil levels did have mean effects over but the levels meastime, lymphocyte significant ured were still in the normal range for Sprague-Dawley rats of the age under study. Histological evaluations were performed on sacrificed animals after the 8-week study period. After 8 weeks the amount of polymeric material remaining in the animals was minimal or nonexistent, albeit small amounts of degraded polymer in three of the Group B rats and two of the Group A rats was found. Normal histology was found in all of the 26 tissues studied for the Controls, Group A, and Group B rats. The primary cellular reaction to the implanted polymer was an accumulation of mononuclear macrophages, present in five of the Group B rats, five of the Group A rats, and none of the Control group rats. There was no indication that the material was cytotoxic. Fibrosis and fibroplasia was present in a few of the implanted rats, but was not a prominent feature. Heterophils, lymphocytes, and plasma cells were rarely encountered, and were not part of the cellular response at the 8-week termination point. The next step was to perform more detailed biocompatibility studies at the site of 20: 80 the application, the brain. Brain biocompatibility was tested for p(CPP-SA) to into rats The results were Surgicel implanted Sprague-Dawley [41]. compared (oxidized regenerated cellulose) and Gelfoam (absorbable gelatin sponge), two The p(CPP-SA) 20: 80 was currently approved materials for brain implantation. tested in 42 hemispheres; the Surgicel, 35 hemispheres; the Gelfoam, 35 hemispheres. The local tissue response of intracranial implantation provided a measure of the immediate effects of implantation to a specific site. The results were categorized into three stages: the acute, subacute, and chronic. During the acute phase (3-6 days), characterized by the predominance of polymorphonuclear leukocytes, the polyanhydride induced the most pronounced reaction, followed by Surgicel, and then Gelfoam. In the subacute phase (6 to 10 days), characterized by the predominance of histiocytes, Surgicel produced the most pro20: 80. In the chronic nounced response, followed by Gelfoam, and the p(CPP-SA) erosion of the implants, phase (15 to 36 days), characterized by near-complete and Surgicel caused the most pronounced reaction, followed by the p(CPP-SA), then Gelfoam. The p(CPP-SA) was completely degraded in 36 days. The inflammatory reaction elicited by the copolymer was comparable to that of Surgicel, but greater than that of Gelfoam. The copolymer elicits a transient, well-demarcated tissue reaction, which subsides as the material is degraded, similar to the response to Surgicel. 50 : 50 were performed on of p(CPP-SA) Studies of the brain biocompatibility twenty adult New Zealand white male rabbits [42]. Each rabbit received a 2 x 2 x 2 mm polyanhydride implant in one frontal lobe and a Gelfoam sponge in the other


346 frontal lobe. Groups of four animals were sacrificed 1, 3, 7, 21, and 60 days postoperatively, and the brains were processed for histological evaluation. Gross and microscopic examination of the brain tissue in contact with the implants indicated that the tissue reaction of the polyanhydride was comparable to that of Gelfoam. 50 : 50 implant on day 7, and There was slightly more necrosis with the p(CPP-SA) a better-demarcated around in reaction Gelfoam general. However, after day gliotic 7 there was little difference in the extent of necrosis between the two materials. There was a comparable fibrous reaction at the craniotomy site for the Gelfoam and There were no hematomas, cysts, or abscesses. polyanhydride. In summary, evaluation of tissue response of the polymer into both rat and rabbit brains indicated the the p(CPP-SA) similar to that polymers had biocompatibility of Gelfoam and Surgicel, two currently approved materials used in cranial surgical procedures. Gelfoam has been used as the standard for evaluation of hemostatic agents and has been used extensively in neurological surgery with minimal adverse effects. As a final safety assessment, the polymers were tested by implantation into the brains of fifteen male Cynomologous monkeys, the test that would most closely simulate conditions in humans [43]. Five monkeys received polymer disks weighing 660 mg with 12.5 mg of BCNU; five monkeys received blank polymer disks without BCNU; five control monkeys received surgery but no implant. Blood samples were obtained from each monkey for hematology and blood chemistry analyses. On day 9 after surgery, non-contrast and contrast-enhanced CAT studies were performed, and on the 12th day, MRI studies were performed. Three monkeys in each group were sacrificed on days 16 to 18 and the remaining two on days 71 to 81. Complete autopsies were performed. There were no adverse systemic or neurological changes in any of the groups, and no differences were found in blood chemistries of hematological studies. It was concluded that the animals showed no neurological deficits from the treatment and tolerated the implants well. In vivo distribution The in vivo spatial distribution and local concentrations of drug in the brain are important variables in evaluating the possible effectiveness of the implant therapy. Accordingly, the distribution of BCNU in normal rabbit brains has been studied [44]. In this study, a comparison between drug distribution following direct injection of BCNU to polyanhydride implants was made. BCNU with 3H labelled on the methylene hydrogens of the chloroethyl groups was prepared to facilitate measurement of the concentration of BCNU or its metabolic products by radiographic 20: 80 of 3 mm techniques. Compression molded polymer devices of p(CPP-SA) diameter x 1 mm thick, 12 mg in weight, containing 2.5, 5, or 10% BCNU, were prepared by both the trituration and cosolution methods. BCNU was introduced to the rabbit brains by device implantation or by stereotactic injection of the labelled BCNU directly into the brain. Three, 7, 14, and 21 days later, the animals were sacrificed and the brains were prepared for quantitative autoradiography. Radiolabelled compound was found within a local region of the brain for conof BCNU-loaded siderably longer after implantation polymers than with direct To the the injection. quantify exposure area, percent of the brain in which the exceeded the two standard deviation units was at least radioactivity background by

347 followed for the different time points. After BCNU-polymer implantation approximately 50% of the area of the brain sections was exposed to radiolabelled compound at 3 days, 15% at 7 days, and < 10% at 14 and 21 days. The direct injection control animals showed almost 0% area exposed to radiolabelled content greater than back3 days after implantation ground after only 3 days. The local brain concentration was found to be approximately 6.5 mM adjacent to the wafers, and approximately 200,uM as far as 10 mm from the implant site. Pharmacologic studies demonstrated that approximately 25% of the tritium label was associated with intact BCNU 3 days after implantation.


Efficacy The positive results from the animal safety studies, which indicated that the polymer would not be harmful to the patient, led to the next step in treatment developmentefficacy studies [45]. A rat model, in which 9-L glioma tumor cells were implanted into the rats' brains, was used to determine efficacy of the implants. Fischer 344 rats were divided into groups of 12 rats each. The control and safety assessment groups underwent tumor cell implantation surgery 4 days prior to treatment but received no actual tumor cells. Of these groups receiving the sham operation, one received no treatment, one received systemic intraperitoneal (IP) BCNU, and the third received BCNU-polymer implants. The test groups received an implantation of the 9-1 glioma tumor cell line 4 days prior to treatment. Of the groups receiving the tumor cells, one had no treatment, one had polyanhydride implants with no BCNU, one had systemic IP BCNU (15 mg BCNU/kg), and one had BCNU-polymer implant (15 mg BCNU/kg). Animals which survived the 125-day course of the study and who were found to have no residual tumor at necropsy were considered to be cured. Survival studies comparing no treatment, polymer alone, injected BCNU, and polymer plus BCNU indicated that significant extension of survival was achieved with the drug-loaded implanted devices. The animals receiving no treatment and polymer without BCNU died at 10.9 (mean) and 11.6 days after tumor implantation with zero cures, and the animals treated with IP BCNU survived 27.3 days after implantation with zero cures. The group treated with the polyanhydride with BCNU survived 62.3 days with two cures. Toxicity of the BCNU in the control animals (animals without tumor implant) showed a longer mean survival of 92.7 days with the implanted BCNU-polyanhydride device, compared to 63.3 days with the IP BCNU. The study suggested that the intracranial delivery of BCNU using polymeric carriers would be superior to systemic administration of the same dose in effectiveness. It also indicated that polymer implants may have a lower toxicity than standard systemic chemotherapy for the same total dose of BCNU. Human

trials

Because of the positive results from the extensive safety studies, strong indication of improved efficacy over systemic therapy, and the suggestion of lower systemic effects than standard systemic chemotherapy, the decision was made to continue the polyanhydride implant therapy to the human clinical trial stage [46]. The initial stage of human clinical studies involves Phase 1, in which assessment of the safety of the treatment to the patient is made, and Phase 2, in which determination of the appropriate dosage is made. A combined Phase 1/2 human clinical trial has been


348 completed on the pilot study of 21 patients, who entered the study between September 1987 and July 1988. The clinical group consisted of patients diagnosed with Grade III or IV anaplastic astrocytoma who were undergoing a second resection, after apparent failure of traditional therapy. All treated patients had an indication for reoperation, presence of a unilatural single focus of tumor, a Karnofsky performance score of at least 60, one course of external beam radiation therapy, and no chemotherapy during the 6 weeks prior to enrollment. The first treatment group of five patients received 1.93 ?Io BCNU loading, for a maximum dose of 31 mg. The second treatment group of five patients received a 3.85To BCNU loading for a maximum dose of 62 mg. The third treatment group of 11patients received 7.35 % BCNU loading for a maximum 102 mg dose of BCNU. The polyanhydride wafers used in the study were 1.4 cm in diameter x 1.0 mm in thickness, prepared by compression molding of a chloroform solution of polymerBCNU that had been spray dried into a powder. They were sterilized by 2.2 Mrads of gamma irradiation. After a baseline examination, patients underwent a craniotomy for maximal resection of tumor. The tumor was removed and wafers were placed on the resection surface to cover as much tissue as possible. Up to eight wafers were used, according to the size of the resection area. Patients were followed blood score, hematology, Karnofsky performance by neurological examination, and CT and MRI were done 1 or 2 and days post-operatively, urinalysis. chemistry, at 14 and 49 days. again From the standpoint of the safety of the treatment, the initial results were with the encouraging, polymer implants being well-tolerated by the patients [46]. No showed a patients significant reduction in blood cell counts that would indicate systemic exposure to BCNU. Blood chemistry and urinalysis did not show evidence of renal or hepatic injury. Traditional systemic administration of BCNU causes enough toxicity so that long recovery intervals are needed between courses. The usual standard single intravenous dose of BCNU is 200 mg. The patients treated with eight wafers received intracranial doses of about 30, 60, and 100 mg of BCNU, depending on treatment group. The BCNU wafers were free from the usual systemic toxic side effects, which was remarkable considering the large amounts of drug released to the tumor site. There was some local tissue response to the implants as measured by CT scan, although these responses were similar to that observed with other local treatments, such as interstitial intratumoral and hyperthermia. radiation, interleukin-2, local tumor destruction was associated with some localized tissue Additionally, necrosis. Some patients required reoperation to remove the necrotic material. A few of the patients who underwent autopsy or reoperation 13 to 23 weeks after implantation, showed persistent disks in situ. There did not appear to be any correlation between the persistence of polymer remnants and specific neurological effects in individual patients. Median post-implant survival times were 65, 47, and 23 weeks for treatment groups 1, 2 and 3, respectively, with corresponding mean survival periods of greater than 65, 64, and 32 weeks. Overall median survival time was greater than 46 weeks post-implant (some patients still living), compared to the expected 36 week median survival after second surgery. In the first two groups at the lower BCNU doses, 8 of 10 patients survived beyond the expected 36 weeks. In the third group, 4 of 111 patients survived beyond the expected 36 weeks. Based on this study, the maximum

349 tolerated dose was assessed to be at the 3.85?1o BCNU loading. However, the initial study involves too few patients to be conclusive and all the results must be considered preliminary. Currently, the therapy is undergoing Phase 3 trials in a placebocontrolled, randomized clinical trial at over 32 medical centers in the US and Canada. At the time of this publication, the implants have been placed into more than 250 patients. OTHER APPLICATIONS


Studies using polyanhydrides to deliver a variety of active agents are underway several animal models for various diseases. Alzheimer's

in

disease

Currently, Alzheimer's disease afflicts over one million people in the United States. Studies have suggested that the loss of basal forebrain cholinergic neurons in the cerebral cortex is an important feature of Alzheimer's disease. The degeneration of neurons in this region of the brain correlates with diminuition of cortical cholinergic activity. It is hypothesized that the resultant low local levels of acetylcholine are associated with clinical memory and learning deficits. Treatments using cholinergic agonists of acetylcholine precursors or acetylcholinesterase inhibitors to increase stimulation of cholinergic neural receptors have been disappointing. It was suspected that the methods of delivery of these agents, either systemically or by direct infusion into cerebral ventricles, may have caused detrimental stimulation of cholinergic receptors outside of the basal forebrain region. Thus it has been proposed that localized controlled delivery of agents to stimulate the cholinergic receptors in the basal forebrain area may provide benefits unattainable by conventional delivery methods. Howard et al. have performed in vivo investigations of the controlled, localized into the release of bethanechol, an acetylcholinesterase-resistant cholinomimetic, basal forebrain area of a rat model of Alzheimer's disease [47]. The delivery system was bethanechol-loaded 50 : 50 microspheres prepared by the solvent p(CPP-SA) removal method. Microspheres ranged in size from 1 to 5,um. Twenty rats received bilateral fimbria-fornix and lesions, which resulted in cholinergic denervation impairment of spatial memory. After a 1-week recovery period, rats were trained for 2 weeks on a radial-armed maze. Performance assessment of these untreated rats indicated poor spatial memory, with the rats entering maze arms with near random frequency. After the 2-week training period, the rats were divided into treatment groups. Five rats received bilateral intrahippocampal implants of saline, five rats received blank polymer, and ten rats received bethanechol-loaded polymer. Spatial on the radialmemory performance was assessed for 40 days after implantation armed maze. The control groups treated with saline and blank polymer showed no improvement over their baseline untreated performance, but the bethanechol implanted groups improved significantly. Improvement was rapid, reaching statistical significance during the first 10-day period. The improved performance was maintained throughout the entire 40 days of the experiment. These encouraging results indicated that lesion-induced memory deficits in rats were successfully loaded with bethanechol and that the of p(CPP-SA) reversed with implantation polyanhydride treatment has potential as a neurosurgical method of treatment of disorders. Alzheimer's disease and other neurodegenerative

350


Osteomyelitis Osteomyelitis is a bone infection which can be difficult to treat by conventional means. Current treatment for osteomyelitis involves systemic antibiotics, implantation of nondegradable carriers, and surgical debridement. Systemic treatment requires high serum levels of drug to deliver enough antibiotic to the site of infection, and accordingly, can result in serious toxicity risks to various organ systems in the bone cement polybody. Local drug releasing systems using the non-degradable methylmethacrylate (PMMA) loaded with antibiotics have been developed, with good efficacy results. However, the PMMA must be removed using a second The bioerodible polysurgical procedure several weeks after initial implantation. anhydride system offers the advantage of obviating the necessity for the second surgical procedure. This concept was tested in a rat model of osteomyelitis [48]. 50: 50 powder with 10% Matrices were prepared by dry mixing of p(CPP-SA) (w/w) gentamicin, the antibiotic, followed by compression molding into cylindrical devices of 3 mm in diameter and 5 mm in length. Similar PMMA bone cement followed by devices were prepared by loading PMMA with 1007o gentamicin, fabrication into rods of 3 mm in diameter and 5 mm in length. Osteomyelitis was induced by injection of a suspension of 106 colony-forming units of Staphylococcus aureus into the right and left proximal tibiae of 20 Sprague-Dawley rats. A 2 x 3 mm preformed plastic cylinder was placed into the bone to act as a foreign body to encourage infection. After 3 weeks the foreign body implant was removed and the animals were divided into four groups. Group I received no treatment; Group II PMMA matrices loaded with 10% (w/w) gentamicin; received the nondegradable 50 : 50 polymer implants without drug; Group IV III received Group p(CPP-SA) received p(CPP-SA) 50: 50 implants loaded with 10% (w/w) gentamicin. After 3 additional weeks, the rats were sacrificed and examined. The tibiae of both PMMA/gentamicin and polyanhydride/gentamicin treated animals appeared to have resolving infections. Control wound sites showed evidence of chronic osteomyelitis. Quantitative culture of the amount of colony forming units found at the implant site were taken. The polyanhydride/gentamicin implants were of bacteria over the control groups, and perfound to reduce the concentration formed as well or better than the current therapy using the nonerodible polymer, PMMA, impregnated with gentamicin. Clinical trials using a biodegradable polymer to release gentamicin for the treatment of osteomyelitis will begin this year. Insulin release One major advantage of bioerodible controlled release systems over most other types of delivery systems is that the release of drug is theoretically controlled primarily by the properties of the polymer and not of the drug. This gives biodegradable polymers the potential to release macromolecules, such as proteins. Such technology would have tremendous impact on the effectiveness of protein-based drugs. Critical challenges remain in the development of formulation procedures which will maintain the protein in its active form device fabrication. In vivo investigations of the release of insulin, a 5600 dalton molecular weight protein, were performed to assess the feasibility of releasing an active protein from polyanhydride microspheres. Two methods, hot melt [21] and solvent removal [22], were tested for microencapsulation their ability to give physiological therapeutic effects for extended periods.


351 Insulin was incorporated into p(CPP-SA) 20: 80 microspheres by hot melt microand into p(CPP-SA) 50:50 by solvent removal. In the hot melt encapsulation microencapsulation method, 1-,um insulin particles were mixed with p(CPP-SA) 20 : 80 at 15% loading of insulin. The mixture was suspended in an organic oil, heated to approximately 80°C (5°C above the melting point of the polymer), and stirred continuously. The emulsion was cooled until the polymer microspheres solidified. In the solvent removal method, insulin, sieved to less than 50,um, was 50: 50. The hot melt method incorporated at 5 and 100% loadings into p(CPP-SA) produced microspheres with a denser structure but cannot be used on the 50 : 50 copolymer because of the high melting point. The solvent removal method can be used on the more hydrophobic p(CPP-SA) 50 : 50 polymer, but produces porous spheres. The physiological effectiveness of the drug-polymer implants was tested on rats by injection of diabetic rats. Diabetes was induced into Sprague-Dawley at 65 dose. Three rats were left uninjected as streptozotocin mg/kg body weight normal controls; five rats comprised a diabetic control group; six rats were treated with 40 to 50 mg of p(CPP-SA) 20 : 80 hot-melt-produced 850 to microspheres, at 15% insulin five rats received 200 of mg p(CPP-SA) 1000/mi diameter, loading; 50 : 50 solvent-removal-produced microspheres, 50 to 425 ,um diameter, at 5% insulin 50 : 50 solvent-removalloading; and five rats received 100 mg of p(CPP-SA) 10% 50 at to 425,um diameter, loading. Initially, it was produced microspheres, in to the as a saline suspension through an 18 gauge inject microspheres attempted imneedle. However, the microspheres did not suspend well, and subcutaneous a 1 cm incision in the dorsum of the rats was of the plantation microspheres through used to introduce the microspheres to the animal. The implants of the hot melt method were implanted on the 4th and 22nd day. Implants of the micropheres prepared by the solvent removal method were implanted after 7 to 10 days. The effect of insulin was assessed by measuring blood glucose levels. The results indicated that at least partial activity of insulin was maintained by both encapsulation processes. Normalglycemia was achieved in the rats for a 3- to Near microspheres. 4-day period using the hot-melt produced insulin-polymer normal glucose levels were maintained for 1 day for the 5% loaded solvent-removalproduced microspheres, and 4 days for the 10% loaded solvent-removal-produced microspheres. Therefore, it is possible to incorporate proteins into polyanhydrides and release them in a biologically active form. Additionally, release occurs over an extended period. CONCLUSIONSAND FUTURE WORK Efforts continue to develop new and better polymers with improved mechanical, chemical, and pharmacological properties and to fully understand the physical and chemical behavior of these systems. Future studies must assess any loss of activity into the delivery device and of sensitive drugs (e.g. proteins) during incorporation after implantation into the body. Mathematical models which can accurately predict and the release rates of incorporated the degradation profiles of polyanhydrides drugs are needed. Nonetheless, critical steps toward the development and clinical of polyanhydrides have been taken. It is expected that this foundaimplementation both as new tion will enable many new potential uses for the polyanhydrides,

352 biomaterials (e.g., bone plates, vascular grafts) and as drug delivery systems (e.g., release of long acting anesthetics or anti-arrhythmia drugs). We also believe that the studies conducted will provide a framework for how new degradable polymers can progress from a conceptual stage to clinical implementation.


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