Paclitaxel release from micro-porous PLGA disks

Chemical Engineering Science 64 (2009) 4341 -- 4349

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Paclitaxel release from micro-porous PLGA disks Lai Yeng Lee a,b , Sudhir Hulikal Ranganath b , Yilong Fu b , Jasmine Limin Zheng b , How Sung Lee c , Chi-Hwa Wang a,b, ∗ , Kenneth A. Smith a,d a

Molecular Engineering of Biological and Chemical Systems (MEBCS), Singapore-MIT Alliance, 4 Engineering Drive 3, Singapore 117576, Singapore Department of Chemical and Biomolecular Engineering, National University of Singapore, 4 Engineering Drive 4, Singapore 117576, Singapore c Department of Pharmacology, National University of Singapore, Clinical Research Center Bldg MD11, Level 5, #05-9, 10 Medical Drive, Singapore 117597, Singapore d Department of Chemical Engineering, Massachusetts Institute of Technology, 77 Massachusetts Avenue, Cambridge, MA 02139, USA b

A R T I C L E

I N F O

Article history: Received 5 March 2009 Received in revised form 25 June 2009 Accepted 4 July 2009 Available online 23 July 2009 Keywords: Supercritical fluid Foam Polymer processing Microstructure Paclitaxel Drug delivery

A B S T R A C T

Micro-porous biodegradable polymeric foams have potential applications in tissue engineering and drug delivery systems. A two-stage fabrication process combining spray drying and supercritical gas foaming is presented for the encapsulation of paclitaxel in micro-porous PLGA (poly lactic glycolic acid) foams. Encapsulation of paclitaxel in the PLGA polymer matrix was achieved and these foams have potential application as a new type of surgical implant for controlled release of paclitaxel. This technique may also be applied to other hydrophobic drugs which face problems of slow release when encapsulated in a compact polymeric device. The micro-porous structure helps to increase drug release rate due to a shorter diffusion path of the drug in the polymer. The final residual organic solvent content in the polymer was low and well within safety limits due to the high miscibility of supercritical CO2 with the organic solvent. The pore size distribution, the phase behavior, and the in vitro swelling behavior of the foams were characterized. In vitro release results showed a nearly constant release rate for up to 8 weeks. The release profiles from micro-porous foam and from compressed disks were compared to assess the performance of micro-porous foams as sustained release implants. The foams implanted intracranially in mice showed therapeutic concentrations of paclitaxel at distant regions of the brain even after 28 days of implantation. © 2009 Elsevier Ltd. All rights reserved.

1. Introduction Supercritical fluid techniques have been explored for processing of polymeric materials due to their high liquid-like dissolving power and gas-like transport properties (Eckert et al., 1996). Supercritical carbon dioxide (CO2 ) has been selected for many processes in pharmaceutical product fabrication due to its accessible critical temperature and pressure, abundance, and generally environmentally benign nature in comparison to many organic solvents. Different forms of polymeric devices such as particles, fibers and foams have been achieved with various supercritical fluid techniques (Reverchon and Cardea, 2007). With supercritical CO2 as a foaming agent, polymeric foams with micro-porous structures can be obtained. Three-dimensional micro-porous biodegradable polymer structures serve an important role in tissue engineering (Mooney et al., 1996; Lu et al., 2000; Singh et al., 2004) and have potential ∗ Corresponding author at: Molecular Engineering of Biological and Chemical Systems (MEBCS), Singapore-MIT Alliance, 4 Engineering Drive 3, Singapore 117576, Singapore. Tel.: +65 6516 5079; fax: +65 6779 1936. E-mail address: [email protected] (C.-H. Wang). 0009-2509/$ - see front matter © 2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.ces.2009.07.016

applications in drug delivery systems (Hile et al., 2000; Hile and Pishko, 2004). Porous scaffolds with interconnected pores are useful in tissue engineering to maximize cell seeding, attachment, growth, vascularization and extracellular matrix production. Due to its biocompatibility, poly dl lactide-co-glycolide (PLGA) has commonly been selected for use in these applications (Singh et al., 2004; Lu et al., 2000; Hile et al., 2000; Hile and Pishko, 2004; Mooney et al., 1997; Thomson et al. 1995, 1998). Conventional methods of micro-porous foam formation such as the solvent-casting, particulate leaching technique (Lu et al., 2000) are usually associated with the use of large amounts of organic solvents which may require extensive purification steps to remove the residual solvent. By using supercritical CO2 as the foaming agent, the use of organic solvent may be minimized or even eliminated in the production of PLGA foams (Mooney et al., 1996). The supercritical gas foaming technique is illustrated in Fig. 1: (i) polymer is first placed in a high pressure vessel; (ii) the polymer is contacted with supercritical CO2 , which diffuses into the polymer matrix to produce a solution of CO2 in PLGA; (iii) upon rapid depressurization, the material solidifies to form solid polymeric foam with a micro-porous structure (as shown in Fig. 1(iv)).

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Fig. 1. Schematic representation of the mechanism of supercritical CO2 gas foaming technique. (i) Polymer is first placed in a high pressure vessel; (ii) the polymer is contacted with supercritical CO2 , which diffuses into the polymer matrix to produce a solution of CO2 in PLGA; and (iii) upon rapid depressurization, the material solidifies to form solid polymeric foam with a micro-porous structure (iv).

Fig. 2. Schematic of experimental setup for the supercritical gas foaming system used. (C1) Refrigerating circulator; (C2) circulating water bath; (P1) high pressure liquid pump; (HP) high pressure view cell; (V1) on/off ball valve; (BPR) automatic back pressure regulator with needle valve; and (M) custom-made mold inserted in the high pressure view cell for holding the foaming samples.

2.2. Production of PLGA foams and compressed disks Our previous studies have shown that micro-porous PLGA foams fabricated in this way may be applied to tissue engineering (Zhu et al., 2008) and DNA delivery (Nie et al., 2008). The drug-encapsulating micro-porous foams developed in this study have potential application in the post-surgical treatment of cancer by controlled release of the chemotherapeutic agent. The concept is to implant a controlled release device immediately after the surgery to target any remaining cells and prevent their proliferation (Mu and Feng, 2001). Paclitaxel is a commercially available anti-cancer agent (Mu and Feng, 2001; Singla et al., 2002; Ewesuedo and Ratain, 2004; Wang et al., 2003; Xie and Wang, 2005) but it is also a highly hydrophobic molecule with very low aqueous solubility. Moreover, it has been found to be associated with problems of extremely slow release when encapsulated in compressed disks of PLGA (Wang et al., 2003). This is likely due, at least in part, to the very low diffusivity of paclitaxel in the polymer matrix. The result may be a sub-therapeutic dose. In addition, failure to deliver most of the encapsulated drug prior to the bulk degradation of the polymer leads to the possibility of a sudden burst of toxic drug to the surrounding tissue during the degradation stage. An alternative formulation which releases the drug steadily would be preferred. This work investigates the production and characterization of paclitaxel-loaded micro-porous foams as implantation materials for post-surgical chemotherapy treatment of gliomas as well as some forms of carcinomas. A two-step production process combining spray drying and the supercritical gas foaming technique was devised to encapsulate paclitaxel in PLGA foams for sustained delivery of the chemotherapeutic agent. Finally, the bio-distribution of paclitaxel in the mouse brain after the foams were implanted for 28 days was investigated to demonstrate the in vivo performance of the foams.

2. Materials and methods 2.1. Materials PLGA with a range of copolymer compositions (varying lactic acid: glycolic acid ratio) was used in this study. PLGA 50:50 (Cat no. P2191; MW = 40–75 kDa)), PLGA 65:35 (Cat no. P2066; MW = 40–75 kDa), PLGA 85:15 (Cat no. 430,471; MW = 40–75 kDa), poly ethylene glycol (PEG, MW = 8 kDa) and phosphate buffered saline (PBS, pH = 7.4) were purchased from Sigma Aldrich (St Louis, MO, USA). Paclitaxel was a generous gift from Bristol Myers Squibb. Compressed carbon dioxide (CO2 ) (Air Liquide Paris, France) was purchased from Soxal (Singapore Oxygen Air Liquide Pte Ltd.). Dichloromethane (DCM) (DS1432, HPLC grade), N, Ndimethylformamide (DMF) (DR0456) and acetonitrile (ACN) (AS1122, HPLC grade) were purchased from Tedia (Fairfield, OH, USA). Millipore water (Millipore Corporation, Billerica, MA) was used throughout the study.

Blank PLGA foams were fabricated from PLGA pellets (in the original form as purchased without further purification). The experimental equipment (with CO2 as the supercritical fluid) is shown in Fig. 2. CO2 was liquefied in a heat exchanger before it was introduced into the high pressure vessel by a high pressure liquid pump. Known amounts of PLGA was placed in a custom-made mold and contacted with supercritical CO2 at pressures ranging from 8 to 15 MPa and contact times ranging from 1 to 24 h. After which, the vessel was depressurized through a back pressure regulator (BPR) at 14 to 16 of the fully open position (i.e. a depressurization rate of around 0.05 MPa/s). Temperature was maintained at 35 ◦ C. Paclitaxel-loaded foams and compressed disks were prepared from drug-loaded PLGA microparticles obtained by spray drying (Buchi 191 Mini Spray Drier, Flawil, Switzerland). Typically, 2% (w/v polymer in DCM) solution was prepared for the production of microparticles by spray drying at 700 N L/min, 100% aspiration ratio and inlet temperature of > 50 ◦ C. Microparticles of 100 mg and a contact time of 120 min with supercritical CO2 were used for preparation of the drug-loaded foam. For paclitaxel-loaded compressed disks, 150 mg of microparticles were subjected to a compressive stress of 2 ton/m2 for 10 min. Cylindrical disks, 3 mm in diameter and 1 mm high, were subsequently cut out of the foams and compressed disks for in vitro studies. The masses of the foam disks and the compressed disks were approximately 3 and 8 mg, respectively. The micro-porous foams were solid and could be handled with surgical instruments such as pincers without observable deformation in shape or size. 2.3. Pore size and surface morphology of foams Scanning electron microscopy (SEM, JEOL JSM-5600 LV, Japan) was used to study the porous structure and surface morphology of the polymeric foams. A very thin slice of the foam was fixed on SEM copper studs with carbon tape. The samples were then coated with platinum (Autofine Coater, JEOL JFC-1300, Japan) at a current of 40 mA for 40 s. The pore size and size distribution of the microporous foams were determined by measuring the Ferets' diameter (the greatest distance possible between any two points along the boundary of a region of interest) of the pores from the SEM images (∼70–100 samples for each measurement) with the SMILEVIEW (JEOL, Japan) program. Field emission scanning electron microscopy (FESEM, JEOL, Japan) was used to study the cross-sections of the foams at higher resolutions to determine the presence of crystalline paclitaxel during in vitro release. 2.4. Differential scanning calorimetry Differential scanning calorimetry (DSC, 2920 modulated, Universal V2.6D TA instruments) was used to study the glass transition

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temperatures (Tg ) of the polymeric foams and the physical state of paclitaxel in the polymer matrix (Dubernet, 1995). Approximately 5 mg of sample was weighed and loaded onto standard aluminum pans (40 mg) with lids. A blank aluminum pan was used as reference in all of the analyses. The samples were purged with pure dry nitrogen at a flow rate of 5 ml/min. PLGA foams and original pellets were analyzed at a constant temperature ramp of 10 ◦ C/min from 20 to 100 ◦ C.

2.5. Encapsulation efficiency and in vitro release High pressure liquid chromatography (HPLC, Shimadzu with UV detectors, Japan) was used to determine the encapsulation efficiency and in vitro release profile of the paclitaxel-loaded foams and disks. An ACN:water (1:1 v/v) solution was used as the mobile phase for HPLC analysis at a flowrate of 1 ml/min and detection was carried out at a wavelength of 227 nm using a C18 column. Standard concentrations of paclitaxel in the mobile phase were used to prepare a calibration curve for the determination of paclitaxel concentration. Measurements of the encapsulation efficiency and of the release rate from paclitaxel-loaded samples were performed in triplicate. For encapsulation efficiency, 3 mg of sample was dissolved in 1 ml DCM to extract the paclitaxel embedded in the polymer matrix. After the DCM had evaporated, the sample was re-dissolved in the mobile phase, filtered with 0.22 m syringe filters, and analyzed using HPLC. For release rate measurements, Paclitaxel-loaded foams and compressed disks were weighed and each was soaked in 5 ml of PBS and placed in a shaking water bath (120 rpm) maintained at 37 ◦ C to mimic physiological conditions. At predetermined time intervals, the PBS was removed for sampling and fresh PBS was added. The paclitaxel content in the supernatant was extracted with DCM, re-dissolved in mobile phase, filtered and analyzed using HPLC.

2.6. Residual organic solvent The spray dried microparticles were placed in a freeze dryer (Alpha 1-4; Martin Christ, GmbH, Germany) for 24 h to remove residual DCM. Residual content of DCM in samples of paclitaxelloaded microparticles and foams was determined using Gas chromatography—mass spectrometry (GC/MS HP 6890 series with an auto-sampler). DMF was used as the mobile phase for the analysis. Known amounts of drug-loaded samples were dissolved in DMF to extract the residual DCM. The extracted solution was subsequently filtered with syringe filters (0.22 m) and used for GC/MS analysis. Standard concentrations of DCM in DMF were prepared as a calibration curve for the determination of DCM concentration.

2.7. Animal care and maintenance All animal experiments were performed with approval from the SingHealth's Institutional Animal Care and Use Committee (IACUC) and experimental practices were conducted in accordance with the National Advisory Committee for Laboratory Animal Research (NACLAR) guidelines in facilities licensed by the Agri-food and Veterinary Authority of Singapore (AVA). BALB/c male mice of about 5 weeks were housed in cages of five and left to acclimatize for a week in the Department of Experimental Surgery of Singapore General Hospital prior to the experiment. Free access to food and water was given. The anesthesia administered during surgical procedures was 50 mg/kg Ketamine and 5 mg/kg Valium.

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2.8. In vivo bio-distribution of paclitaxel in mouse brain Animals were randomized into two control groups, namely sham (no implants) and placebo (blank PLGA implant) containing four animals each and two experimental groups, namely F100 (100 g/mouse paclitaxel implant) and F200 (200 g/mouse paclitaxel implant) containing five animals each. Both the foams (F100 and F200) were made from PLGA 50:50 with 10% (w/w) paclitaxel loading. The animals were anaesthetized and immobilized in a mouse adaptor on a stereotactic frame. A midline incision was made on the scalp and a burr hole was placed 2 mm lateral to the sagittal sinus at the midpoint between the bregma and lambda. The formulation was then implanted into the burr hole and the wound closed using suturing. At 28 days post-implantation, the animals were euthanized by carbon dioxide inhalation. The mice brains were then removed from the skulls and frozen at −80 ◦ C for further analysis. Later, each brain was sectioned coronally into 11 sections using an acrylic mouse brain matrix yielding 1 mm thick coronal brain sections. Weight measurement of each section was then taken. Quantification of paclitaxel in the brain sections was performed by liquid chromatography tandem mass spectrometry (LC-MSMS) analysis.

3. Results and discussion Fig. 3 shows micro-porous structures of PLGA 50:50 foams obtained at different operating conditions. By carefully manipulating the operating pressure and contact time during the supercritical gas foaming process, mean pore sizes ranging from 20 to 500 m can be achieved. Goel and Beckman (1994, 1995) described the relationship between pressure and pore density by using classical nucleation theory. The energy barrier for homogenous nucleation of a gas bubble in the polymer solution is given by (Goel and Beckman, 1994; Reverchon and Cardea, 2007):

Ghomo =

163 3P 2

(1)

where Ghomo represents the energy barrier for homogenous nucleation, P is the magnitude of the pressure difference between the gas nucleus and the polymer/gas solution;  is the surface tension of the polymer/gas interface. From Eq. (1), it can be deduced that for a higher pressure drop the energy barrier for nucleation is lower, and hence the homogenous nucleation rate and the cell density should both increase. Fig. 4 shows the pore size and size distribution of PLGA 50:50 foams achieved at varying conditions of pressure and contact time. It may be observed that, as operating pressure increases, the corresponding pore size tends to decrease. Similar trends of pressure vs. pore size have been reported for supercritical gas foaming of polystyrene (Reverchon and Cardea, 2007; Sumarno et al., 2002; Arora et al., 1998) and cellulose acetate (Reverchon and Cardea, 2007). It should also be noted that thermodynamic equilibrium seems to require a contact time of about 12 h.

Fig. 3. Scanning electron micrographs (at 200× magnification) of blank PLGA 50:50 foams fabricated at varying operating pressure and contact time to yield different pore sizes.

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Fig. 4. Average pore size and size distribution of PLGA 50:50 foams achieved at operating pressure of 10, 12 and 15 MPa for varying contact times.

PLGA 85:15 pellets used in this study have Tg values of 42.1 and 46.1 ◦ C, respectively (see “first cycle” data). The Tg values for PLGA 50:50 and PLGA 85:15 foams were 45.0 and 53.3 ◦ C, respectively, and were higher than the Tg s for the original pellet form. Fleming and Kazarian (2005) examined the effect of high pressure CO2 processing on polymer materials and reported on the CO2 -induced crystallinity of semi-crystalline polymers. The shift in glass transition temperature after the foaming process is most likely to be due to the rapid removal of CO2 from the PLGA/CO2 solution during the venting process of the supercritical gas foaming. Presumably, the rapid departure of the CO2 induces high rates of strain which in turn produce polymeric orientation. This more ordered polymeric structure is retained as the PLGA solidifies following the venting process. To investigate the effect of CO2 on the Tg of PLGA polymer, repeated DSC experiments were performed on PLGA foams and pellets. The Tg of foams was first determined by DSC analysis from 20 to 100 ◦ C at a constant heating rate of 10 ◦ C/min. The polymer was then maintained at 100 ◦ C for 20 min before slow cooling to 20 ◦ C at a constant cooling rate of 2 ◦ C/min. This was to allow the polymer to relax to its natural conformation without quenching. After cooling, the sample was maintained at 20 ◦ C for 10 min before a second Tg analysis was performed on the cooled sample. A similar analysis was performed on PLGA pellets. After the foam was heated far above its Tg and allowed to cool back to 20 ◦ C, the Tg of the polymer decreased to 43.7 and 50.4 ◦ C for PLGA 50:50 and PLGA 85:15 foams, respectively. The Tg s were very close to the values for PLGA pellets with similar treatment (second cycle of Tg analysis). Clearly, the CO2 induced a more oriented structure in PLGA during the foaming process, leading to the higher Tg value for the foam. This process may be reversed by heating the foam samples to a temperature above its Tg , and subsequently cooling the samples back to 20 ◦ C as illustrated in Fig. 5. 3.2. PLGA foams for controlled release Paclitaxel is a highly hydrophobic molecule with a low diffusivity in the polymer matrix. The release of paclitaxel from compressed disks has been observed to be very slow (Wang et al., 2003). In this study, a highly porous polymer matrix was used to circumvent the problems associated with slow release. Micro-cellular polymeric foams have mean pore and wall (between pores) sizes of approximately 50 and 10 m, respectively. The interconnectivity between pores enhances water penetration and the shorter diffusion path in the polymer increases the release rate of hydrophobic molecules. Fig. 6a and b shows the microparticles of paclitaxel-loaded PLGA achieved by spray drying. High loading efficiency is a very important factor in most pharmaceutical processes as the active ingredient is costly. The loading efficiency for the supercritical gas foaming method was determined by loading efficiency (%) =

Fig. 5. Determination of glass transition temperatures (Tg ) of (i) PLGA 50:50 and (ii) PLGA 85:15 before and after foaming.

3.1. Tg analysis of PLGA foams with varying copolymer compositions Blank micro-porous PLGA foams with varying lactic to glycolic acid ratios, namely PLGA 50:50 (A), PLGA 65:35 (B) and PLGA 85:15 (C) were fabricated at a pressure of 12 MPa and a contact time of 4 h. As the lactic acid content in the copolymer increases, its glass transition temperature (Tg ) increases correspondingly. From the DSC results in Fig. 5a, it can be seen that the original PLGA 50:50 and

Drug loading in foams (mg/mg) ×100% Drug loading in microparticles (mg/mg) (2)

The drug loading efficiency for the drug-loaded foams achieved ranged from 80.7% to 99.9%. The foam samples fabricated in this study, A2% (PLGA 50:50 with 2% drug loading), C2%(PLGA 85:15 with 2% drug loading), AP2% (PLGA 50:50 and 5% PEG with 2% drug loading) and CP2% (PLGA 85:15 and 5% PEG with 2% drug loading) are shown in Fig. 6(c)–(f), respectively. The pressure and contact time required to achieve a micro-porous structure with a mean pore size of approximately 50 m is only 8 MPa and 2 h when microparticles were used for foaming. This is attributed to the small size of the microparticles as compared to the original polymer pellets used in foaming pure PLGA, thereby allowing better diffusion of CO2 into PLGA. Therefore, if microparticles are used as

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Fig. 6. Scanning electron micrographs of paclitaxel-loaded formulations: (a) PLGA 50:50 (A2%) microparticles; (b) PLGA 85:15 (C2%) microparticles; (c) PLGA 50:50 (A2%) foam; (d) PLGA 85:15 foam (C2%); (e) PLGA 50:50 (AP2%) foam; and (f) PLGA 85:15 (CP2%) foam.

the precursor for supercritical gas foaming, relatively low contact times are required for the production of micro-porous foams with small pores. 3.3. In vitro release profiles The in vitro release profiles of paclitaxel from microparticles and nanoparticles of PLGA and poly l lactide (PLA) usually follow a diffusion mechanism with relatively fast release during the initial period followed by a much slower release rate after the first 2 weeks (Mu and Feng, 2001; Xie and Wang, 2005). Paclitaxel in PLGA disks obtained by using compression molding have shown extremely slow cumulative release profiles (Wang et al., 2003). Foam and compressed disks with similar cylindrical dimensions (3 mm diameter by 1 mm height) and drug loading (2% (w/w)) were compared in this work. Fig. 7a shows the cumulative release profile of paclitaxel from PLGA 50:50 foams and compressed disks. Drug release from PLGA 50:50 foams showed a continuous and nearly linear release of approximately 50% by 8 weeks. The release from micro-porous foams has a complex mechanism. In general, initial drug release is mainly by diffusion of drug through the polymer matrix into PBS.

Fig. 7b summarizes the performance of PLGA 50:50 foam and compressed disks. The initial amount of drug in the compressed disk is much higher than in the foam, due to its higher density. After 8 weeks, less drug was released from the compressed disks than from the foam despite the presence of more drug in the former. The amount of drug remaining in the compressed disk is also much higher than the amount of drug remaining in the foam disk after 8 weeks of release, which suggests inefficient release from compressed disk formulations in contrast to foam disks. PLGA implants are designed to undergo biodegradation, thereby eliminating the need for removal of the implant from the patient after the drug delivery treatment. This is meant to reduce the trauma on the patient receiving the treatment. PLGA 50:50 implants typically degrade by 10–18 weeks (Hile et al., 2000; Wu and Wang, 2001; Lu et al., 1999). From Fig. 7a and b, it was observed that, by week 8, compressed disks release only approximately 10% of the drug. This raises toxicity concerns for the compressed disks formulations, as the remaining 90% of the drug may be released in a short time following the degradation of PLGA 50:50. This large burst of drug at the later stage of the treatment may pose a potential risk to the patient. Sporadic toxicity in rats each implanted with 10 mg PCPP-

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Fig. 7. (a) Comparison of release profiles of paclitaxel from PLGA 50:50 foams and PLGA 50:50 compressed disks over a period of 8 weeks. (b) Comparison of release performance of paclitaxel from PLGA 50:50 foams and PLGA 50:50 compressed disks over a period of 8 weeks (without the addition of PEG).

SA implants containing 20% (w/w) taxol has been reported (Walter et al., 1994). The implants showed a very high initial burst of about 30% in the first four days which could be the reason for the toxicity effects in the animals. On the other hand, PLGA 50:50 foam disks release more than half the encapsulated drug by week 8. Furthermore, the total amount of drug released from foam is also higher than that from the compressed disks. This demonstrates the enhanced drug delivery achieved with micro-porous foams. The diffusion coefficient of paclitaxel (Dpaclitaxel ) in extracellular space (ECS) is 9×10−10 /m s2 (Fung et al., 1998). The dimensions of the drug-encapsulating disks are 3 mm diameter by 1 mm width. Assuming a diffusion path (L) of 0.5 mm (half the width of the disk) for paclitaxel from the liquid-filled pores of the foam to the surrounding fluid medium, the Fourier number, (Dpaclitaxel · t/L2 ) ∼ 1 when t∼300 s. Thus, the characteristic time to establish quasi-steady behavior is very much less than the duration of the experiment. The diffusion path (L) is likely to be an underestimate of the actual diffusing length of paclitaxel in the liquid-filled pores as not all the pores in the micro-porous polymer are interconnecting.

If the process were controlled by diffusion in the pores, the flux would be approximately Dpaclitaxel · CS /L. The solubility of paclitaxel (CS ) in water is 30 g/ml (Swindell et al., 1991) and the concentration in the ambient is essentially zero, so the driving force would be approximately CS . Given the approximate path length and the diffusivity that have already been cited, the estimated flux is found to be approximately 5×10−8 kg/m2 s, which is two orders of magnitude greater than the experimental mass flux of 7×10−10 kg/m2 s. This suggests that the rate-limiting step for the drug release lies in the diffusion of drug through the polymer matrix. To further increase the release rate of paclitaxel from PLGA foams, the thickness of the walls between the pores could be reduced by controlling the foaming conditions. Addition of a hydrophilic additive such as PEG may enhance water penetration into the polymer matrix and increase the drug release rate. PEG is non-toxic and is commonly employed to improve drug release of hydrophobic drugs from biodegradable microspheres (Wang et al., 2003). The effect of addition of PEG on the release rate of paclitaxel from PLGA 50:50 foams and compressed disks is also illustrated in Fig. 7a. For both foam and compressed disk formulations, the addition of PEG (5%) to the polymer matrix increases the release of paclitaxel only slightly. The release profiles of paclitaxel from PLGA 50:50 and PLGA 85:15 were compared to determine the effect of the copolymer composition. Fig. 8 shows the results for PLGA 50:50 and PLGA 85:15 foams at 2% and 10% drug loading. For both drug loadings, it was observed that the release rate from PLGA 50:50 was higher. PLGA 50:50 is more hydrophilic than PLGA 85:15 due to a higher content of glycolic acid, and also exhibits more pronounced swelling in vitro. Therefore, faster drug release is achieved as compared to the more hydrophobic PLGA copolymers. A slower fractional release is observed at high drug loading (10%) as compared to lower drug loading (2%) This may be explained by super-saturation of paclitaxel in PLGA, especially after the PLGA has been exposed to PBS (Lee et al., 2008) At higher drug loadings, drug crystals may form inside the polymer matrix (Polakovic et al., 1999; Lee et al., 2008). This was investigated by comparing samples with 2% (A2%) and 10% (A10%) drug loading. Samples were removed from PBS after 3 days and washed thrice with deionized water to remove any released drug that might be adhering to the surface of the polymer matrix. The foam samples were then vacuum-dried and the cross-section of the foams was examined at high resolution by field emission scanning microscopy (FESEM). Fig. 9 shows the cross-section of paclitaxel-loaded PLGA 50:50 foams at 2% and 10% drug-loadings. At 10% drug loading, the cross-section

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Fig. 9. FESEM micrographs of the interior of paclitaxel-loaded PLGA 50:50 foams (a) and (c) sample A2%; (b) and (d) sample A10%.

3.4. Residual DCM content in paclitaxel-loaded foams The production of blank PLGA foams by the supercritical gas foaming technique does not require the use of an organic solvent. However, DCM was used as a solvent for paclitaxel and PLGA during the spray drying process when encapsulating paclitaxel in PLGA microparticles. To ensure that the residual DCM content in the formulations was within safety limits, the residual DCM content in the samples was examined. Fig. 10 shows the residual DCM content in the polymer at different stages of the production process for sample A 10%. At the end of the spray drying process, a relatively high residual DCM content of approximately 300 ppm was detected. After freeze drying (24 h) to remove DCM from the spray dried microparticles, the DCM content had been reduced to approximately 200 ppm. Foaming (2 h contact time) the microparticles with supercritical CO2 reduced DCM content to approximately 150 ppm. This suggests that supercritical CO2 removes DCM more efficiently and in shorter times

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of the polymeric foam exhibit signs of drug re-crystallization in the polymer matrix, as shown in Fig. 9b and d. In contrast, this phenomenon was not observed at lower drug loading of 2% as shown in Fig. 9a and c. This suggests that at 2% drug loading, the drug release is mainly affected by diffusion, polymer swelling and polymer degradation. At 10% drug-loading, the drug release is also rate-limited by the dissolution of the drug crystallites into the polymer matrix, and hence the lower fractional rate of drug release. This inference is further bolstered when the paclitaxel release from 5% drug loaded foams reported by Ong et al. (2009) are compared to the 2% and 10% drug-loaded foams in this study. Clearly, after 35 days of release, the % cumulative release from the 5% drug-loaded foams falls in between that of the 2% and 10% drug-loaded foams. Furthermore, our study confirms the role of drug-loading in controlling the release rate, which has not been investigated by Ong et al. (2009) and thus gives an additional insight.

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Fig. 10. Residual DCM content (ppm) in (a) spray dried microparticles; (b) sample (a) after 1 day of freeze drying; (c) sample (a) after supercritical gas foaming process; and (d) sample (b) after supercritical gas foaming.

than freeze drying. This is due to the excellent miscibility of supercritical CO2 with DCM as well as with other organic solvents. The amount of organic solvent left will be very low after processing with supercritical CO2 . The lowest residual DCM was achieved by freeze drying the microparticles before foaming, as shown in Fig. 10. The residual DCM content of the paclitaxel-loaded PLGA foams were well below the allowable limits of 600 ppm by US Pharmacopeia (2002).

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Distance from foam implantation site (mm) Fig. 11. In vivo bio-distribution of paclitaxel in the mouse brains released from F100 foam (100 g paclitaxel/mouse paclitaxel) and F200 foam (200 g paclitaxel/mouse) after 28 days of foam implantation. Each data point represents five samples (n = 5 animals) and the error bars represent the standard deviation.

3.5. In vivo bio-distribution of paclitaxel in the mouse brain This work was carried out to study the bio-distribution of paclitaxel in BALB/c mice brain tissue after implantation of paclitaxelloaded PLGA foam, F100 (100 g paclitaxel/mouse) and F200 (200 g paclitaxel/mouse) for 28 days and to investigate the effect of drug dosage on the penetration in the brain. Two control groups, namely Sham (no implants) and Placebo (blank PLGA implant) were used for LC-MSMS calibration purposes. Fig. 11 shows the bio-distribution of paclitaxel released from F100 and F200 foam implants in the brain tissue of the BALB/c mice. F100 implants showed concentrations of 1000–10000 ng/50 mg brain tissue within 2 mm of the implant, 100–1000 ng/50 mg brain tissue within 2–5 mm of the implant and 10–100 ng/50 mg brain tissue even at 10 mm away from the implant. For F200, paclitaxel concentrations were 1000–10 000 ng/50 mg brain tissue within 3 mm of the implant, 100–1000 ng/50 mg brain tissue within 3–8 mm of the implant and 10–100 ng/50 mg brain tissue thereafter. Both F100 and F200 implants were able to provide elevated levels of paclitaxel throughout the brain even after 28 days. For the lower dosage F100, although paclitaxel concentration fell sharply from 1000 to 10 000 ng/50 mg brain tissue within 4 mm of implantation to only 10–100 ng/50 mg brain tissue beyond 10 mm of implantation site, the lower concentrations are still higher than the mean peak serum concentration of paclitaxel necessary for cancer chemotherapy, which has been reported to be around 1 M or 1 ng/mg brain tissue (Sonnichsen and Relling, 1994). This suggests that the drug is capable of penetrating to a sufficient distance from the polymer implant to treat the tumor effectively. Also, paclitaxel when administered systemically in the commercial form Taxol䉸 has a short half-life of only 1.3–8.6 h (Rowinsky et al., 1990), thus rendering it unlikely to maintain high paclitaxel concentrations for prolonged periods. This is especially so with the blood brain barrier limiting the uptake of the drug. On the other hand, F100 and F200 foams are able to maintain elevated concentrations of paclitaxel intracranially for prolonged periods. In another study, Ong et al. (2009) have

reported the bio-distribution of two formulations: 5% drug-loaded PLGA 85:15 and 10% drug-loaded PLGA 50:50 foam rods each with 100 g paclitaxel in rat brain. They investigated the bio-distribution with respect to distance in the brain and time to evaluate sustainability. In comparison, we studied the dose dependent bio-distribution in the brain and its related toxicity effects on the animal. Our results clearly show that 200 g foam (F200) can be used safely in mice. Furthermore, the F200 foams provided deeper penetration of up to 8 mm in the mouse brain as compared to 3 mm by the F2 foams (10% paclitaxel) in a rat brain reported by Ong et al. (2009). The higher drug dosage in the F200 foams also renders higher cytotoxicity and sustainability than the F2 foams. Thus these foam formulations are promising drug delivery implants for post-surgical chemotherapy of malignant glioma. 4. Conclusions Paclitaxel-loaded PLGA foams were fabricated as an implant material for post-surgical chemotherapy. These foams provide an alternative to conventional methods of making implants by compression molding. By using a two-step production process of spray drying from DCM followed by supercritical gas foaming from CO2 , dissolution of paclitaxel in a porous polymer matrix was achieved. Microparticles can quickly become saturated with CO2 , so the contact time required to produce foams with pores of a given size is less than would be needed with pellets. Micro-porous foams also have short diffusion paths within the polymer and this allows for more rapid drug release. The release profiles for the micro-porous foams fabricated as 3 mm disks exhibit a relatively constant release rate up to 8 weeks in vitro. The micro-porous foams with small pores (around 50 m) and uniformly porous structure were observed to be mechanically stable and could be used as an implant in surgical experiments. Therapeutic levels of paclitaxel (mean peak serum concentration needed for cancer chemotherapy) were achievable in brain tissue even after 28 days up to 10 mm from the site of implantation. Furthermore, this technique may be applied to encapsu-

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