Poly(ortho esters): synthesis, characterization, properties and uses

Advanced Drug Delivery Reviews 54 (2002) 1015–1039 www.elsevier.com / locate / drugdeliv

Poly(ortho esters): synthesis, characterization, properties and uses Jorge Heller a , *, John Barr a , Steven Y. Ng a , Khadija Schwach Abdellauoi b , Robert Gurny b a

A.P. Pharma, 123 Saginaw Drive, Redwood City, CA 94063, USA University of Geneva, School of Pharmacy, 30 Quai E. Ansermet, CH1211, Geneva 4, Switzerland

b

Received 22 March 2002; accepted 19 June 2002

Abstract Over the last 30 years, poly(ortho esters) have evolved through four families, designated as POE I, POE II, POE III and POE IV. Of these, only POE IV has been shown to have all the necessary attributes to allow commercialization, and such efforts are currently underway. Dominant among these attributes is synthesis versatility that allows the facile and reproducible production of polymers having the desired mechanical and thermal properties as well as desired erosion rates and drug release rates that can be varied from a few days to many months. Further, the polymer is stable at room temperature when stored under anhydrous conditions and undergoes an erosion process confined predominantly to the surface layers. Important consequences of surface erosion are controlled and concomitant drug release as well as the maintenance of an essentially neutral pH in the interior of the matrix because acidic hydrolysis products diffuse away from the device. Two physical forms of such polymers are under development. One form, solid materials, can be fabricated into shapes such as wafers, strands, or microspheres. The other form are injectable semi-solid materials that allow drug incorporation by a simple mixing at room temperature and without the use of solvents. GMP toxicology studies on one family of POE IV polymers has been concluded, an IND filed and Phase I clinical trials are in progress. Important applications under development are treatment of post-surgical pain, osteoarthritis and ophthalmic diseases as well as the delivery of proteins, including DNA. Block copolymers of poly(ortho ester) and poly(ethylene glycol) have been prepared and their use as a matrix for drug delivery and as micelles, primarily for tumor targeting, are being explored.  2002 Elsevier Science B.V. All rights reserved. Keywords: Poly(ortho ester); Bioerodible polymer; Injectable semi-solid; Drug delivery; Protein delivery; Post-surgical pain; Periodontal disease; Ophthalmic delivery; Micelles; Tumor targeting

Contents 1. Introduction ............................................................................................................................................................................ 1016 2. Poly(ortho ester) I.................................................................................................................................................................... 1016 3. Poly(ortho ester) II .................................................................................................................................................................. 1017 *Corresponding author. Tel.: 1 1-605-366-2626; fax: 1 1-650-365-9452. E-mail address: [email protected] (J. Heller). 0169-409X / 02 / $ – see front matter  2002 Elsevier Science B.V. All rights reserved. PII: S0169-409X( 02 )00055-8

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4. Poly(ortho ester) III ................................................................................................................................................................. 5. Poly(ortho ester) IV ................................................................................................................................................................. 5.1. Polymer synthesis ............................................................................................................................................................ 5.2. Polymer hydrolysis .......................................................................................................................................................... 5.3. Control of mechanical and thermal properties..................................................................................................................... 5.4. Polymer fabrication .......................................................................................................................................................... 5.5. Polymer stability .............................................................................................................................................................. 5.6. Control of erosion rate...................................................................................................................................................... 5.7. Control of drug delivery rates ........................................................................................................................................... 5.7.1. Low molecular weight, water-soluble drugs ............................................................................................................. 5.7.2. Macromolecular drugs ............................................................................................................................................ 6. Semi-solid polymers ................................................................................................................................................................ 6.1. Polymer molecular weight control ..................................................................................................................................... 6.2. Polymer stability .............................................................................................................................................................. 6.3. Drug release .................................................................................................................................................................... 6.3.1. Triethylene glycol semi-solids................................................................................................................................. 6.3.1.1. Post-surgical pain control ........................................................................................................................... 6.3.2. 1,10-decanediol semi-solids .................................................................................................................................... 6.3.2.1. Periodontal disease .................................................................................................................................... 6.3.2.2. Ocular applications .................................................................................................................................... 6.3.2.3. Estrus synchronization in sheep .................................................................................................................. 7. Block copolymers with poly(ethylene glycol) ............................................................................................................................ 7.1. Synthesis ......................................................................................................................................................................... 7.1.1. Release of proteins ................................................................................................................................................. 7.1.2. Micelles ................................................................................................................................................................ 8. Conclusions ............................................................................................................................................................................ Acknowledgements ...................................................................................................................................................................... References ..................................................................................................................................................................................

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1. Introduction

2. Poly(ortho ester) I

Poly(ortho esters) have already been extensively reviewed so no attempt will be made to provide a further exhaustive review. Instead, the reader is referred to a comprehensive article in Advances in Polymer Science that covers all work with poly(ortho esters) prior to and including 1992 [1], and a chapter in Encyclopedia of Controlled Drug Delivery that includes all work up to 1998 [2]. Poly(ortho esters) have been under development since 1970 and during that time, four polymer families have been developed. These are shown in Scheme 1. In this review, we will concentrate on POE IV and on the latest developments in the use and characterization of this material. However, for the sake of completeness, for each poly(ortho ester) family, we will present a brief, updated review with pertinent references, and refer the reader to the original literature for additional details.

Poly(ortho ester) I, the first such polymer prepared, has been developed at the Alza Corporation and described in a series of patents by Choi and Heller [3–7]. This polymer was originally designated as Chronomer 姠 but the name was later changed to Alzamer  . The polymer is prepared as shown in Scheme 2. When placed in an aqueous environment, the polymer will hydrolyze as shown in Scheme 3. Because ortho ester linkages are acid sensitive and hydrolysis of this polymer produces g -butyrolactone which rapidly opens to g -hydroxybutyric acid, the polymer must be stabilized with a base such as Na 2 CO 3 to avoid an uncontrolled, autocatalytic hydrolysis reaction. The polymer has been used in the treatment of burns [8], in the delivery of the narcotic antagonist naltrexone [9] and in the delivery of the contraceptive steroid levonorgestrel [10]. The polymer has

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Scheme 1.

Scheme 2.

Scheme 3.

also been investigated by Sudmann in a number of orthopedic applications [11–19]. However, the autocatalytic nature of the hydrolysis, as well as the rather low glass transition temperature of the polymer has severely limited its application and this polymer is no longer under development.

3. Poly(ortho ester) II Poly(ortho ester) II was developed at the Stanford Research Institute, now known as SRI International, under contract with the National Institute of Child Health and Human Development. A history of the

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Scheme 4.

development of this polymer has been published [20]. Poly(ortho ester) II is prepared by the addition of a diol to the diketene acetal 3,9-diethylidene 2,4,8,10tetraoxaspiro[5.5]undecane, as shown in Scheme 4 [21,22]. This synthesis represent a very significant improvement over that used to prepare poly(ortho ester) I since it is merely necessary to dissolve the monomers in a polar solvent such as tetrahydrofuran and to add a trace of an acidic catalyst. The reaction proceeds to completion virtually instantaneously, is highly reproducible and polymer molecular weights can be readily adjusted by controlling stoichiometry. Polymer hydrolysis occurs as shown in Scheme 5 [23]. Even though the hydrolysis produces acidic hydrolysis products, unlike POE I, it is not autocatalytic because the initial hydrolysis products are neutral. Details of the hydrolysis mechanism have been published [24]. Even though the synthesis is via a polycondensa-

tion reaction, no small molecule by-products are evolved so that dense, crosslinked materials can be prepared by using triols, or higher functionality hydroxy-containing monomers, as shown in Scheme 6 [25]. The ability to form dense, crosslinked matrices that completely biodegrade to small, watersoluble molecules is unique to poly(ortho esters). Another significant advantage is that mechanical and thermal properties of the polymer can be adjusted by using diols having different degrees of chain flexibility, so that materials that range from hard, glassy materials, to materials that are semisolids at room temperature can be obtained. However, because poly(ortho esters) are extremely hydrophobic materials [26], the amount of water available to react with the hydrolytically labile ortho ester linkage is limited and for this reason, under physiological conditions, the polymer is very stable. Therefore, means of controlling erosion rates of the polymer must be devised if devices having erosion rates that can be varied from a few days to months are desired. Because ortho ester linkages are acid-labile [27], one means of controlling polymer erosion rates is by lowering the pH at the polymer–water interface. A considerable amount of work has been expended in attempts to achieve control of erosion rates by taking advantage of the acid-sensitivity of ortho ester linkages by the addition of acidic excipients, for example suberic acid, to the polymer matrix. This approach was only marginally successful and for that reason devices using acidic excipients have never reached commercialization. Because that work has been reviewed in considerable detail [1,2], it will not be summarized here.

4. Poly(ortho ester) III

Scheme 5.

This polymer was originally developed at the Stanford Research Institute [28] using venture capital

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Scheme 6.

funding and following the transfer of the technology to the University of Geneva [29], was then extensively characterized and investigated by this group in ocular and other applications. This work has recently been reviewed [30]. The polymer is prepared as shown in Scheme 7 [28]. Because this polymer has a very flexible backbone, it is a semi-solid material at room temperature. Such materials offer a number of important advantages. Dominant among these is the possibility of incorporating therapeutic agents by a simple

mixing procedure at room temperature without the use of solvents and the fact that such materials can be easily injected. However, molecular weight must be limited so that injection through a needle size no larger than about 20 gauge is possible. Unlike the synthesis of POE II which is virtually instantaneous after dissolution of the monomers in tetrahydrofuran and the addition of a trace of an acidic catalyst, the synthesis shown in Scheme 6 requires long reflux times and azeotroping out ethanol to drive the reaction to high molecular weight.

Scheme 7.

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Scheme 8.

Also, unlike the synthesis of POE II, it is very difficult to achieve reproducible results. The polymer hydrolyzes as shown in Scheme 8 [31]. A detailed study of the hydrolysis mechanism using model compounds has established that only the 1 and 2 isomeric esters are produced. Thus, the absence of the 6-isomer indicates that cleavage occurs exclusively by an exocyclic cleavage [31]. Even though this material offers a number of significant advantages, and has shown excellent biocompatibility in rabbit eyes [32], difficulties in synthesis and especially difficulties in reproducibly preparing materials of chosen molecular weights, has made practical applications difficult and this material is no longer under development.

5. Poly(ortho ester) IV POE IV is a modification of POE II that allows control over erosion rates without the need to add acidic excipients.

5.1. Polymer synthesis As already mentioned, POE II is an extremely hydrophobic material and despite the reactivity of ortho ester linkages when exposed to water, polymer hydrolysis is slow and devices fabricated from the polymer are highly stable. This stability is shown in Fig. 1, which shows weight loss of a wafer as a function of time when placed in a pH 7.4 buffer at 37 8C [33]. To make this polymer system adaptable to a wide range of delivery times, it is essential to devise a means of reproducibly controlling erosion rates. This has been achieved by the incorporation of a short segment based on glycolic acid, or lactic acid into the polymer backbone [34]. Such segments act as latent acids, because they readily hydrolyze to glycolic, or lactic acids, which then catalyze hydrolysis of ortho ester linkages in the polymer backbone. Then, by controlling the concentration of such segments in the polymer, rate of erosion can be accurately controlled. Such polymers have been

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Fig. 1. Weight loss as a function of time for a polymer prepared from 3,9-dimethylene-2,4,8,10-tetraoxaspiro[5.5]undecane and 1,6-hexanediol. 0.05 M phosphate buffer, pH 7.4, 37 8C (from Ref. [33]).

designated as autocatalyzed poly(ortho esters). They are prepared as shown in Scheme 9 [34]. The condensation reaction between a diol and a diketene acetal proceed virtually instantaneously

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Fig. 2. Effect of polymerization time on molecular weight of a poly(ortho ester) prepared from 15 mole% trans-cyclohexanedimethanol, 40 mole% 1,6-hexanediol, 40 mole% triethylene glycol and 5 mole% triethylene glycol monoglycolide (from Ref. [35]).

after addition of a trace of an acidic catalyst to a tetrahydrofuran solution of the monomers [35]. The virtually instantaneous rate of reaction is shown in Fig. 2. A detailed characterization of the polymer has been carried out [36].

Scheme 9.

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Scheme 10.

The latent acid diol is prepared by the reaction between a diol and either lactide, or glycolide, as shown in Scheme 10 [36]. This reaction produces a range of products in varying concentrations with n-values varying between 1 and 7, as shown by the gel permeation chromatogram trace in Fig. 3. To control erosion rates, the exact structure of the latent acid diol is not important, and it is the total concentration of the a -hydroxy acid segments in the polymer that determines erosion rates.

5.2. Polymer hydrolysis Polymer hydrolysis occurs as shown in Scheme 11 [37]. The hydrolysis proceeds in three consecutive steps. In the first step, the lactic acid or glycolic acid segment in the polymer backbone hydrolyzes to generate a polymer fragment containing a carboxylic acid end-group that will catalyze ortho ester hydrolysis. A second cleavage produces free a -hydroxy acid that also catalyzes hydrolysis of the ortho ester

links. Further hydrolysis of ortho esters then proceeds in two steps to first generate the diol, or mixture of diols used in the synthesis and pentaerythritol dipropionate, followed by ester hydrolysis to produce pentaerythritol and propionic acid. Fig. 4 shows polymer weight loss and release of lactic acid [37]. The most significant finding of this study is the linearity of weight loss and the concomitant release of lactic and propionic acid. While linear rate of weight loss alone does not necessarily indicate surface erosion [38], the concomitant linear weight loss and release of lactic acid argues convincingly for a process confined predominantly to the surface layers of the polymer matrix. The process is not pure surface erosion because there is a significant drop in molecular weight of the remaining polymer, indicating that some hydrolysis is taking place in the bulk material. The following process would account for the observed facts. Initially, due to the highly hydrophobic nature of the polymer surface, no hydrolysis takes place, and an induction period is observed.

Fig. 3. Gel permeation chromatograph of reaction products between lactide and triethylene glycol.

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Scheme 11.

Fig. 4. The relationship between lactic acid release ( ? ) and weight loss (j) for a poly(ortho ester) prepared from 3,9 diethylidene-2,4,8,10-tetraoxaspiro [5.5] undecane, and a 100 / 70 / 30 mixture of 1,10-decanediol and 1,10-decanediol lactide. 0.13 M, pH 7.4 sodium phosphate buffer at 37 8C (from Ref. [37]).

Gradually, as a result of hydrophilic species generated by the hydrolysis process, sufficient water penetrates the surface layers to induce a more rapid

hydrolysis and once steady state has been reached, constant erosion with concomitant generation of hydrolysis products takes place. In parallel with this process, some water penetrates the bulk material and induces limited polymer chain cleavage. However, due to the limited number of polymer chains cleaved, the material remains very hydrophobic and water concentration in the bulk remains very low. An erosion process that is confined predominantly to the surface layers has a number of important consequences. First, if the drug is well immobilized in the matrix, its release is controlled by polymer erosion so that an ability to control polymer erosion translates into an ability to control rate of drug release. Second, because drug release is controlled by polymer erosion, drug release and polymer erosion take place concomitantly and when drug release has been completed, no polymer remains. Third, because most of the hydrolysis occurs in the outer layers of the device, acidic hydrolysis products diffuse away from the device and do not accumulate in the bulk material. Thus, the interior of the matrix

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does not become highly acidic, as is the case with poly(lactide-co-glycolide) copolymers, or poly(lactic acid), and acid-sensitive drugs can be released without loss of activity. An ongoing study is designed to provide a direct measurement of the pH in the matrix as a function of depth [39]. To do so, pH-sensitive nitroxide radicals having different pKa values are incorporated into the matrix and the EPR signal analyzed. Such data provide a direct measure of the pH in different layers within the device and the mobility of the nitroxide radicals provides a measure of water penetration into the matrix. Preliminary data are consistent with a surface erosion process and indicate that the pH in the outer eroding layer is about 5.4 and the pH in the matrix interior is about 6.4. The exact values may require a slight adjustment pending a determination of the pKa of the nitroxide radicals in water and in the interior of the hydrophobic polymer environment.

Fig. 5. Glass transition temperature of 3,9-diethylidene-2,4,8,10tetraoxaspiro[5.5]undecane, trans-cyclohexanedimethanol, 1,6hexanediol polymer as a function of mole% 1,6-hexanediol (from Ref. [23]).

5.3. Control of mechanical and thermal properties One of the most desirable attributes of POE IV synthesis, shown in Scheme 9, is the ability to vary independently thermal and mechanical properties and erosion rates. Mechanical and thermal properties can be varied by choosing the appropriate R-group in the diol and in the latent acid. The use of rigid diols, produces materials having high glass transition temperatures and hence high fabrication temperatures, while the use of flexible diols produces materials having low glass transition temperatures and hence low fabrication temperatures. The use of very flexible diols produces materials that are semi-solid materials at room temperature and these will be covered later in this chapter. The effect of diol structure on the glass transition temperature is shown in Figs. 5 and 6. Fig. 5 shows the effect of varying the ratio of a rigid diol, transcyclohexanedimethanol and a flexible diol, 1,6-hexanediol on the glass transition temperature [23]. Clearly, by an appropriate choice of diol ratio, any glass transition temperature between 110 and 20 8C can be selected. Fig. 6 shows the effect of chain length of a diol on the glass transition temperature [40]. Figs. 5 and 6 were generated with polymers having no latent acid. However, the latent acid diol

Fig. 6. Effect of diol chain length on the glass transition temperature of polymers prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane and a,v -diols (from Ref. [40]).

does have a significant effect on the glass transition temperature, as shown in Fig. 7 [36]. Thus, both diol structure and latent acid diol structure must be considered when designing polymers having desired thermal and mechanical characteristics.

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Fig. 7. Glass transition temperatures for poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane and noctanediol, n-decanediol and n-dodecanediol, each with 5, 10 and 20 mole% of the corresponding lactide.

5.4. Polymer fabrication Polymer fabrication is a crucial component in the development of drug delivery devices and poorly fabricated devices, even with the best of polymers, can exhibit very poor drug release kinetics and a significant burst and non-linear kinetics are a common consequence of a poorly fabricated device. The most desirable situation is one where finely micronized drug particles are individually surrounded by polymer and there is minimal particle-toparticle contact. Although this is not always the case, drug loading should, in general, be no higher than about 30 wt%, although devices with good release characteristics were obtained with drug loading as high as 70 wt%. Poly(ortho esters) are excellent thermoplastic materials that can be easily fabricated by extrusion, injection molding or compression molding. The polymer is very thermally stable as shown in Fig. 8 [35]. The polymer is also soluble in tetrahydrofuran, ethyl acetate and methylene chloride, so that microparticles using conventional microencapsulation techniques can easily be prepared. In developing

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Fig. 8. Effect of thermal processing on molecular weight of a poly(ortho ester) prepared from 3,9 diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, trans-cyclohexanedimehanol, 1,6-hexanediol, triethylene glycol and triethylene glycol glycolide (15 / 40 / 40 / 5). Processing parameters shown in legend (from Ref. [35]).

microencapsulation methods, it is essential to fabricated dense microspheres since porous microspheres will, in general, show an initial drug burst and drug release kinetics that are not always linear. A highly desirable fabrication method is a solventless thermal fabrication method. One such procedure is an extrusion method where finely ground polymer and a micronized protein are intimately mixed and then extruded into thin strands at temperatures low enough so that the stability of the incorporated therapeutic agent is not compromised. The ability to control both mechanical and thermal properties poly(ortho ester) makes an extrusion fabrication method possible, and the method has been applied to fabricating thin strands with incorporated proteins without loss of activity [42].

5.5. Polymer stability As shown in Fig. 9, poly(ortho esters) have excellent stability, and when stored under anhydrous conditions, they are stable at room temperature. The particular polymer used in this stability study is a hydrophilic polymer containing 40 mole% latent acid that has been ground to produce microparticles, thus, greatly increasing surface area. This is a very rapidly

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cause irradiation generates free radicals that in a solid matrix can be long-lived, stability studies were extended to 3 months to determine if post-irradiation chain cleavage takes place. As shown in Fig. 10 [35], the polymer is stable after the initial drop in molecular weight and EPR studies have shown that free radicals dissipate in less than 24 h.

5.6. Control of erosion rate

Fig. 9. Stability of a polymer prepared from 3,9 diethylidene2,4,8,10-tetraoxaspiro[5.5]undecane, trans-cyclohexanedimethanol, triethylene glycol and triethylene glycol glycolide (35 / 25 / 40) stored at room temperature and under anhydrous conditions.

eroding polymer that would completely erode in a matter of a few days if placed in an aqueous buffer. The polymer is also relatively stable when sterilized by irradiation [35]. As shown in Fig. 10, there is a decrease in molecular weight after irradiation, but the decrease is of the same order of magnitude as that observed with other bioerodible polymers. Be-

Fig. 10. Effect of g -irradiation at 24 kGy on a poly(ortho ester) prepared from 3,9 diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, trans-cyclohexanedimethanol, 1,6-hexanediol, triethylene glycol and triethylene glycol glycolide (15 / 40 / 40 / 5). Polymer stored post-irradiation at 5 8C in a dessicator (from Ref. [35]).

From a consideration of the hydrolysis mechanism, it can be inferred that two factors will affect erosion rates. One is the concentration of the latent acid in the polymer backbone and the other is the hydrophilicity of the matrix that can be controlled by using triethylene glycol as one of the diols used in the synthesis. Polymer hydrophilicity has a major influence on erosion rate of the polymer, as illustrated in Fig. 11, which shows the rate of erosion of two polymers having the same concentration of latent acid, but differing degrees of hydrophilicity [35].

Fig. 11. Weight loss as a function of time for a poly(ortho ester) prepared from 3,9 diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, and a mixture of (j) 25 mole% trans-cyclohexanedimethanol, 45 mole% 1,6-hexanediol, 5 mole% triethylene glycol and 25 mole% triethylene glycol glycolide, (d) 75 mole% trans-cyclohexanedimethanol and 25 mole% trans-cyclohexanedimehanol. 0.13 M, pH 7.4 sodium phosphate buffer at 37 8C (from Ref. [35]).

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Polymer hydrophilicity and latent acid concentration are interrelated, and highly hydrophobic matrices are relatively insensitive to the concentration of latent acid, while hydrophilic matrices are very sensitive to the concentration of latent acid. This sensitivity is shown in Fig. 12. With this particular system, there is no erosion for 30 days in the absence of latent acid, but as little as 1 mole% latent acid changes erosion times to about 3 weeks.

5.7. Control of drug delivery rates In discussing drug delivery, it is of interest to consider two types of drugs, low molecular weight water-soluble drugs and macromolecular drugs.

5.7.1. Low molecular weight, water-soluble drugs This represents the most difficult situation because water-soluble drugs will increase the hydrophilicity of the matrix and thus accelerate erosion rate. Further, if there is matrix swelling, release by diffusion is possible. However, as shown in Fig. 13, this is not the case with poly(ortho esters) and the erosion-controlled release of 5-FU can be demonstrated by the concomitant 5-FU release and device

Fig. 12. Weight loss of a polymer prepared from 3,9 diethylidene2,4,8,10-tetraoxaspiro [5.5] undecane and (m) trans-cyclohexanedimethanol, triethylene glycol (70 / 30 and (j) trans-cyclohexanedimethanol, triethylene glycol, triethylene glycol glycolide (70 / 29 / 1). Extruded strands 1 3 10 mm, 0.01 M phosphate buffered saline, pH 7.4, 37 8C.

Fig. 13. Polymer weight loss (d) and 5-fluorouracil (5-FU) release (j) from a polymer prepared from 3,9-diethylidene2,4,8,10-tetraoxaspiro [5.5] undecane and 1,3-propanediol / triethylene glycol diglycolide (90 / 10). Drug loading 20 wt%, 0.05 M phosphate buffer, pH 7.4, 37 8C (from Ref. [44]).

weight loss [41]. In this particular case, 5-FU material balance is a little low, but there is little doubt that the predominant mechanism controlling drug release is erosion and not diffusion. Further, as shown in Fig. 14, rate of drug release is proportional to drug loading. These data are consistent with a surface erosion process and provide additional evidence for the hydrolysis and erosion study described previously [37].

5.7.2. Macromolecular drugs As already mentioned under fabrication, solventless incorporation of peptides and proteins into thin, extruded strands, is a particularly attractive method, if the extrusion can be carried out at temperatures lower than the temperature at which the protein begins to denature. Such a study has been carried out with a model protein FITC-BSA that was mixed with polymer and extruded into 1 mm strands cut to 10 mm lengths. The extrusion was carried out at 75 8C. The strands were then placed in a pH 7.4 phosphate buffer at 37 8C and weight loss and FITC-BSA release determinations carried out. Results of this study are shown in Fig. 15 [42].

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Fig. 14. Effect of loading on amount of 5-fluorouracil (5-FU) released from a poly(ortho ester) prepared from 15 mole% transcyclohexanedimethanol, 40 mole% 1,6-hexanediol, 40 mole% triethylene glycol and 5 mole% triethylene glycol glycolide. (j) 5.5 wt% 5-FU (14 mg), (h) 11.6 wt% 5-FU (28 mg), (d) 23.6 wt% 5-FU (56 mg), 0.01 M phosphate buffered saline, pH 7.4, 37 8C (from Ref. [44]).

Three major features are notable. First, there is no burst despite the fact that 15 wt% of a water-soluble material has been incorporated, second, there is a significant lag-time, and third, the rate of BSA release is linear and polymer weight loss and BSA release occur concomitantly. These are important findings, but the long lag-time is not optimal in most applications. For this reason, means of reducing the length of the lag-time were investigated. We have found that the addition of very small amounts of poly(ethylene glycol), molecular weight 2 kDa, markedly changes kinetics of BSA release [42]. Because poly(ethylene glycol) had to be physically mixed into the powdered polymer prior to extrusion, we have chosen a solid poly(ethylene glycol) and hence the choice of 2 kDa. The change in kinetics after addition of 1 wt% poly(ethylene glycol) is shown in Fig. 16. These data indicate that poly(ethylene glycol), even in very small amounts is effective in reducing the lag-time. As could be anticipated, the incorporation of a hydrophilic moiety onto the polymer will accelerate erosion rate and hence release of FITC-BSA.

Fig. 15. Release of FITC-BSA (d) and weight loss (j) from a poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, 1,4-pentanediol and 1,6-hexanediol glycolide (100 / 95 / 5). Strands, 1 3 10 mm, extruded at 70 8C. 0.01 M phosphate buffered saline, pH 7.4, 37 8C. FITC-BSA loading 15 wt% (from Ref. [42]).

Fig. 16. Effect of 2 kDa poly(ethylene glycol) (PEG) on release of FITC-BSA from a poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, 1,3-propanediol and triethylene glycol glycolide (100 / 85 / 15). (j) pure polymer, (d) 1 wt% PEG, (Strands, 1 3 10 mm, extruded at 70 8C, 0.01 M phosphate buffered saline, pH 7.4, 37 8C. FITC-BSA loading 15 wt% (from Ref. [42]).

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6. Semi-solid polymers

6.1. Polymer molecular weight control

To prepare semi-solid materials, it is necessary to use highly flexible diols, and in order to allow direct injection of such materials, their viscosity must be limited by preparing polymers having molecular weights no higher than about 5 kDa. While it is, of course, possible to prepare a great variety of semisolid polymers by using a range of highly flexible diols, in order to limit toxicology studies preparatory to regulatory approval, this work was confined to two diols, triethylene glycol and 1,10-decanediol. Semi-solids based on triethylene glycol produce somewhat hydrophilic materials while semi-solids based on 1,10-decanediol produce hydrophobic materials. The most significant advantage of semi-solid materials is that therapeutic agents can be readily incorporated by mixing at ambient temperature and without the use of solvents. Such conditions are mild enough to incorporate even the most sensitive therapeutic agents. Mixing can be accomplished on a small scale by using a mortar and pestle, but on a somewhat larger scale it is better carried out using a three-roll mill [43].

Polymer molecular weight control can be achieved by using an excess of diol relative to the diketene acetal, but can also be accomplished by using a chain-stopper. In this approach, a calculated amount of a monofunctional alcohol is used [44]. As shown in Scheme 12, when n-decanol is used as a chainstopper in combination with 1,10-decanediol, both polymer ends have n-decanol residues so that a chain-stopped material is somewhat more hydrophobic relative to a stoichiometry-controlled material that has terminal hydroxyl groups. The use of chain-stoppers allows excellent and reproducible molecular weight control by varying the ratios of 1,10-decanediol to n-decanol as shown in Fig. 17 [44]. The existence of terminal methyl groups has been established by 1 H NMR studies [44]. Because the principal means of administration of semi-solid materials is by injection, preparation of materials having reproducible viscosities is important. Synthesis reproducibility as measured by Brookfield viscosity, for a number of different preparations is shown in Fig. 18 [45]. Clearly, the

Scheme 12.

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Fig. 17. Effect of n-decanol on the molecular weight of a poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, 1,10-decanediol and 1,10-decanediol lactide (100 / 70 / 30 (from Ref. [44]).

Fig. 19. Room temperature stability for a polymer prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, triethylene glycol, and triethylene glycol glycolide (60 / 50 / 50). Material stored under anhydrous conditions (from Ref. [45]).

semi-solid polymer after storage at room temperature under anhydrous conditions for 9 months [45]. As with the solid polymers, within experimental error, there is no change. Fig. 20 shows the effect of irradiating the polymer at a dose of 22.9–25.6 kGy [45]. Within experimental error, no changes in molecular weight could be detected. This is consistent with previous finding for POE III and indicates that low molecular weight

Fig. 18. Variation in Brookfield viscosity for nine typical preparations at 25 8C. Injectable formulation prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, triethylene glycol, and triethylene glycol glycolide (60 / 50 / 50). Formulation contains 20 wt% monomethoxy poly(ethylene glycol), molecular weight 550 (from Ref. [45]).

synthesis is sufficiently reproducible to assure that the same viscosity materials can be repeatedly prepared.

6.2. Polymer stability Fig. 19 shows changes in molecular weight of a

Fig. 20. Stability of a polymer prepared from 3,9-diethylidene2,4,8,10-tetraoxaspiro[5.5]undecane, triethylene glycol, and triethylene glycol glycolide (60 / 50 / 50) irradiated at 24 kGy (from Ref. [45]).

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polymers, unlike their high molecular weight analogues, do not significantly change molecular weight on irradiation.

6.3. Drug release In developing semi-solid formulations for drug release, two formulations are under development. One formulation is based on triethylene glycol, and the other on 1,10-decanediol.

6.3.1. Triethylene glycol semi-solids Semi-solid materials based on triethylene glycol are somewhat hydrophilic and are designed for relatively short delivery times. 6.3.1.1. Post-surgical pain control Post-surgical pain originates at the site of surgery, a relatively small area, so that treatment by the systemic administration of analgesic agents with the well-known side-effects such of drowsiness, constipation and others does not represent optimal treatment. Instead, the administration of a sustained release formulation at the site of surgery is more appropriate. Poly(lactic acid) microspheres have been evaluated as a sustained delivery system for local anesthetics such as butamben, tetracaine and dibucaine [46,47]. The use of poly(lactic-co-glycolic acid) copolymers was also investigated as a delivery system for the release of bupivacaine and dexamethasone using a rat sciatic nerve blockage model [48,49]. However, in many applications, such as treatment of post-operative pain, an analgesic activity of only a few days would be desirable. Because erosion times of poly(lactic acid) is measured in months, and even years and the erosion time of poly(lactic-co-glycolic acid) copolymers is measured in weeks, these erosion times are clearly not optimal for short term therapy. Because the erosion time of poly(ortho ester) can be as short as a few days days and injectable semi-solid formulations can be prepared, such polymers are currently under investigation as delivery systems for analgesic agents, such as bupivacaine, or mepivacaine. Because anesthetic and analgesic agents such as bupivacaine are cardiotoxic, it is important to control the amount released at the post-operative site to achieve analgesia while blood levels remain well

Fig. 21. Release of bupivacaine from a polymer prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, triethylene glycol, and triethylene glycol glycolide (60 / 50 / 50) containing 50 wt% monomethoxy poly(ethylene glycol) and 10 wt% bupivacaine free base. Phosphate buffer, pH 7.4, 37 8C (from Ref. [45]).

below the cardiotoxic level. Fig. 21 [45] shows the desired short time release using a formulation comprised of 40 wt% polymer, 50 wt% monomethoxy poly(ethylene glycol) M Wt 550 and 10 wt% bupivacaine free base. The poly(ethylene glycol) derivative was added to improve ease of injection. The in vivo functionality of a bupivacaine delivery system has been investigated using a rat sciatic nerve model. In this model, the sciatic nerve in the hind limb of the rat is exposed and the formulation containing bupivacaine placed in the close proximity of the nerve. The hind limb is then electrically stimulated. No response is indicative of total nerve block while dragging of the hind limb is indicative of motor block. Results of that study are shown in Fig. 22 [50]. Clearly, sustained motor block has been achieved while blood levels remained far below the recognized cardiac and CNS toxic level of 2–4 mcg / ml [51]. Additional in vivo data are presented in the following chapter. Another important application of semi-solid materials is in the delivery of peptides and proteins. Clearly, an appealing feature of semi-solids is that therapeutic agents can be incorporated by a simple, room-temperature mixing procedure without the need for solvents. Such a procedure should not com-

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as good as when a hydrophobic, solid polymer was used. Nevertheless, good linear kinetics were obtained. Although no weight-loss determinations were made, visually, there was no semi-solid materials remaining after all the FITC-BA has been released.

Fig. 22. Bupivacaine blood levels and duration of motor block using a polymer prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, triethylene glycol, and triethylene glycol glycolide (60 / 50 / 50) in a rat sciatic nerve model (from Ref. [50]).

promise the integrity of even the most fragile therapeutic agents. This work is currently underway. A preliminary result showing the release of FITC-BSA is shown in Fig. 23 [50]. In this particular case a hydrophilic polymer was used to release FITC-BSA, a watersoluble material. Thus, immobilization of BSA is not

Fig. 23. Release of BSA from a semi-solid poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, and a 95 / 5 mole% mixture of triethylene glycol and triethylene glycol glycolide. BSA loading 4 wt%, 0.1 M phosphate buffer, pH 7.3, 37 8C (from Ref. [50]).

6.3.2. 1,10 -decanediol semi-solids Semi-solid materials based on 1,10-decanediol are very hydrophobic, and their intended application is in delivery times significantly longer than those achievable using semi-solid materials based on triethylene glycol. This material is under investigation in the treatment of periodontal disease, in ocular applications and in estrus synchronization. 6.3.2.1. Periodontal disease Destructive periodontitis is associated with pathogenic flora that resides in deep periodontal pockets [52]. While conventional non-surgical treatments such as mechanical scaling and root planing are in many cases successful, there are situations where refractory periodontitis is encountered and those cases can best be treated by the systemic administration, or via rinses of an antimicrobial agents such as tetracycline. However, the concentration of an antimicrobial agent in the periodontal pocket can be significantly enhanced by placing a sustained delivery device into the periodontal pocket and a number of such devices are on the market [53]. In this application, an injectable 1,10-decanediol formulation with 1,10-decanediol lactide latent acid semi-solid formulation containing tetracycline was used. Such a formulation can be injected directly into the periodontal pocket and was investigated in human volunteers. The in vitro tetracycline free base release is shown in Fig. 24 [45,54]. As the figure shows, excellent linear release for the desired length of time has been achieved. It is this formulation that was used in the human volunteer study. Preliminary findings indicated that the formulations were very well tolerated with no pain during application and no signs of irritation or discomfort after treatment. Further, as shown in Fig. 25 [55], concentrations of tetracycline in the periodontal pocket were well above the minimum inhibitory concentration (MIC) of 1 mg / ml even at day 11.

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at day 11, two out of 24 sites were positive for tetracycline. Results of this study suggest that the injection of a tetracycline formulation into a periodontal pocket is a promising treatment modality, but retention of the device must be significantly improved before such a system can become a successful treatment of periodontitis.

Fig. 24. Cumulative release of tetracycline free base from a polymer prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, decanediol and decanediol lactide (100 / 50 / 50). In PBS at pH 7.4 and 37 8C. Drug loading 20 wt% (from Ref. [55]).

6.3.2.2. Ocular applications The 1,10-decanediol semi-sold material with 1,10decanediol lactide as the latent acid has been investigated in ocular applications. In a biocompatibility study using rabbits, the polymer was injected subconjunctivally, intravitreally, intracamerally and suprachoroidally. In all of these applications, the polymer exhibited outstanding biocompatibility. Additionally, the polymer exhibits surprisingly long lifetimes and it appears that devices having a lifetime of a few months are feasible [56]. Final results of this study are not available at this writing. 6.3.2.3. Estrus synchronization in sheep The objective of estrus synchronization in sheep is to make all female sheep fertilizable at chosen times. While a number of non-degradable devices are on the market either as vaginal sponges or ear-implants, a bioerodible device would clearly be a significant labor-saving device. An injectable 1,10-decanediol formulation containing fluorogestone acetate is currently under investigation as a 14 day delivery device and in vivo testing is in progress [57].

7. Block copolymers with poly(ethylene glycol)

Fig. 25. Average tetracycline concentration in gingival crevicular fluid of devices that have been retained in the periodontal pocket after the placement of a delivery system prepared from 3,9diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, 1,10-decanediol, and 1,10-decanediol lactide (100 / 70 / 30) containing 10 wt% tetracycline free base. Fraction of retained devices as shown above bars (from Ref. [55]).

Block copolymers of poly(ortho esters) and poly(ethylene glycol) were prepared and are currently under investigation as a matrix for drug delivery and as micelles.

7.1. Synthesis AB and ABA block copolymers can be prepared as shown in Schemes 13 and 14.

However, as also shown in Fig. 25, retention of the device in the pocket was not satisfactory and at day 4, nine out of 24 sites were positive for tetracycline, at day 7, 6 / 24 sites were positive for tetracycline and

7.1.1. Release of proteins As discussed under 5.5.2, the addition of very small amounts of poly(ethylene glycol), molecular

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Scheme 13.

weight 2 kDa, was effective in reducing the lag-time in BSA delivery from extruded strands. However, poly(ethylene glycol) had to be physically mixed into the powdered polymer prior to extrusion, and the addition of very small amounts of an excipient represents an additional complications. For this reason, we have investigated the use of an AB-block copolymer containing 6 wt% of a 2-kDa poly(ethylene glycol) in the polymer chain as a matrix for BSA release. Release of FITC-BSA from this AB block copolymer is shown in Fig. 26 [42]. Clearly, there has been a very significant improvement, and the use of AB and ABA block copolymers comprised of poly(ethylene glycol) and poly(ortho esters) is now under active development as release matrices for the delivery of proteins and other therapeutic agents.

7.1.2. Micelles When amphiphilic polymers are dispersed in water, they spontaneously self-assemble into micellar

structures [58]. Such structures have a number of important applications, but perhaps the most important application is tumor targeting using the enhanced permeability and retention effect [59]. Briefly, this effect is based on differences in permeability between normal vasculature and that feeding tumors. Because vasculature feeding tumors in newly formed, it has an incompletely formed endothelium and is thus more permeable than normal vasculature. Thus, macromolecules, or micelles having dimension in the order of 50 nanometers can not escape normal vasculature, but are able to escape tumor vasculature and accumulate in the tumor, and because tumors have poor lymphatic drainage, such agent will be retained in the tumor. In the actual application of micelles in tumor targeting, a hydrophobic anticancer agent is physically entrapped in the hydrophobic core of the micelle and the formulation injected intravenously. The physical entrapment of drugs is preferable to one where the drug is chemically attached to the hydro-

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Scheme 14.

phobic portion of the micelle, because the drug has not been chemically modified thus making regulatory approval easier. Micelles based on poly(ethylene glycol) and poly(ortho esters) are of particular interest because poly(ortho esters) are highly hydrophobic materials so that the entrapment of highly hydrophobic drugs such as doxorubicin, or taxol, should be very efficient. Additionally, poly(ortho esters) are pH-sensitive and because the pH in the interior of tumors is about 5.5, once internalized, the micelles should rapidly degrade and deliver their payload. Micelles having a structure shown in Scheme 15 are currently under investigation [60]. In this particular composition, a poly(ortho ester) segment Fig. 26. Release of FITC-BSA from an AB-block copolymer containing 6 wt% 2 kDa poly(ethylene glycol). Poly(ortho ester) prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro[5.5]undecane, and a 85 / 15 mole% mixture of 1,3-propanediol and triethylene glycol glycolide. Strands, 1 3 10 mm, extruded at 70 8C, 0.01 M phosphate buffered saline, pH 7.4, 37 8C. FITCBSA loading 15 wt% (from Ref. [42]).

Scheme 15.

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based on trans-cyclohexanedimethanol was chosen because this diol will produce the most hydrophobic structure. Ongoing work has shown that the entrapment efficiency of ABA block copolymers is higher than that AB block copolymers based on other biodegradable segments. Optimization studies are currently ongoing, but taxol has been entrapped at a loading as high as 40 wt% which is significantly higher than that achieved with other bioerodible micelles using doxorubicin. Unfortunately, no direct comparison with taxol is available. The stability of these micelles measured by micelle size as a function of pH and in the presence of BSA has been investigated and is shown in Fig. 27. Thus, at the pH encountered in tumor cells, the micelles will rapidly loose their micellar structure and deliver the entrapped drug, in this case taxol, to the tumor. When micelles are injected, they undergo a very significant dilution. For this reason, they must have a very low critical micelle concentration (CMC). We have determined that this value is in the order of 10 24 g / l which is low enough to assure that upon injection, micellar structure will be maintained. Light scattering measurements have shown that micelles in

the desired range of 50–80 nanometers can be prepared. Thus, these micelles appear to be a promising delivery system and are currently under active development.

8. Conclusions Poly(ortho esters) have evolved through a number of families to the current version, POE IV and this polymer has significant potential for producing useful, commercially relevant bioerodible drug delivery products. The success of POE IV is largely based on the versatility of the synthesis that allows excellent control over erosion rates by selecting the proper ratios of diol to latent acid diol, and excellent control of mechanical properties by appropriate selection of diols used in the synthesis. Also of great importance is ease of synthesis and synthesis reproducibility, thermal stability that makes compression, extrusion or injection molding possible and ability to withstand radiation sterilization without excessive loss of molecular weight. Most importantly, the polymer can be stored under anhydrous conditions at room temperature for long periods of time. In vitro and in vivo studies, have shown that the polymer erodes to completion and drug depletion coincides with total polymer erosion. GLP toxicology studies for a triethylene glycol semi-solid polymer have been completed and an IND for post-operative pain control filed. An initial human clinical trial has been completed and a Phase I human clinical trial using injectable, semi-solid formulation is now in the advanced planning stages.

Acknowledgements

Fig. 27. Stability of micelles prepared from poly(ethylene glycol)–poly(ortho ester)–poly(ethylene glycol) (5000–4000–5000) ABA block copolymers. (m) pH 7.4, 10 mM phosphate buffer, 37 8C. (j) pH 5.5, 10 mM citrate buffer, 37 8C. (d) pH 7.4, 10 mM phosphate buffer containing bovine serum albumin, 37 8C.

It has been over 30 years since poly(ortho esters) have been under development, and a large number of individuals have made major contributions to the development of POE I, POE II and POE III. Because this review has concentrated on POE IV, the authors wish to acknowledge the contributors to the development of this polymer system, at AP Pharma, at the University of Geneva and at the University of Gent. At AP Pharma, major contributions were made by Dr

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Hui-Rong Shen, and Dr Ellen Qi Li. At the University of Geneva, major contributions were made by Dr Suzanne Einmahl and Dr Alexandra RothenWeinhold. At the University of Gent, the micelle work was carried out by Professor Etienne Schacht and Dr Veska Toncheva.

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