pH-Sensitive Polymer Blends used as Coating Materials to Control

Biomacromolecules 2005, 6, 2074-2083

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pH-Sensitive Polymer Blends used as Coating Materials to Control Drug Release from Spherical Beads: Importance of the Type of Core Florence Lecomte,† Juergen Siepmann,*,† Mathias Walther,‡ Ross J. MacRae,‡ and Roland Bodmeier† College of Pharmacy, Freie Universitaet Berlin, Kelchstr. 31, 12169 Berlin, Germany and R&D, Pfizer Ltd, Sittingbourne Research Centre, Sittingbourne, Kent ME9 8AG, United Kingdom Received January 31, 2005; Revised Manuscript Received March 24, 2005

The aim of this study was to coat theophylline-loaded spherical beads with pH-sensitive polymer blends to control the resulting drug release kinetics. Various mixtures of ethylcellulose (water-insoluble) and Eudragit L (methacrylic-acid-ethyl-acrylate-copolymer; water-insoluble/water-soluble below/above pH 5.5) were used as coating materials. Two types of theophylline cores were studied: pure drug matrixes and theophyllinelayered sugar cores. Importantly, the type of core significantly affected the resulting drug release patterns. Interestingly, not only the slope, but also the shape of the release curves was altered, indicating changes in the underlying mass transport mechanisms, despite of the identical composition of the polymeric coatings. The observed differences could be explained based on the physicochemical properties of the film coatings and the swelling behavior of the beads upon exposure to the release media. Using this knowledge the development/optimization of this type of drug delivery system can be facilitated and the safety of the pharmacotherapies be improved. Introduction The drug concentration at the site of action in the human body is of central importance for the success of a pharmacotherapy. Too high drug concentrations can lead to serious side effects, whereas too low drug levels lead to the failure of the medical treatment. Using controlled drug delivery systems, the rate at which the drug appears at the target site can be adjusted and, thus, the effects of the pharmacotherapy can be optimized.1-5 Furthermore, as drug release can be controlled over prolonged periods of time, the frequency of drug administration can be reduced (e.g., once daily instead of three times per day). Thus, the quality of life and the compliance of the patient can be improved. Often, polymeric materials are used to control the release rate of the drug out of the pharmaceutical dosage form. Generally, the drug is either directly embedded within a macromolecular network (matrix system) or a drug depot is surrounded by a polymeric membrane (film coating; reservoir system). In the latter case, the diffusion of the drug through the macromolecular shell can control the resulting release kinetics. To adjust desired drug release rates, the coating thickness and type of polymer can be varied or different types and amounts of plasticizers can be added. However, these * To whom correspondence should be addressed. Present address: College of Pharmacy, Universite´ de Lille, 3, rue du Professeur Laguesse, 59006 Lille, France. Phone: +33-3-20964708. Fax: +33-3-20964942. E-mail: [email protected]. † Freie Universitaet Berlin. ‡ Pfizer Ltd.

variations are often restricted, because adequate film coating properties (e.g., mechanical stability and nontoxicity) must be provided. To overcome these limitations, polymer blends can be used as coating materials. By simply varying the polymer blend ratio broad ranges of film properties and, thus, drug release kinetics can be obtained.6,7 Blends of water-insoluble and pH-sensitive polymers (water-insoluble at low pH, water-soluble at high pH) are of particular interest, because the release rate of the drug can be triggered by the pH of the environment.8,9 In the stomach (at low pH), both polymers are insoluble, whereas in the intestine (at high pH), the pH-sensitive component might leach out of the polymeric film, resulting in dynamic changes in the physicochemical properties of the coatings (e.g., increased permeability for the drug). These changes can for instance be used to provide about constant release rates of basic drugs with strongly pH-dependent solubility (being freely water-soluble at low and poorly water-soluble at high pH). The idea is to be able to compensate the decrease in drug solubility along the gastro intestinal tract by a simultaneous increase in drug permeability of the film coating.10 However, the development and optimization of this type of advanced drug delivery systems is difficult, because many factors can be involved in the control of drug release and need to be considered.8,11,12 In addition to the properties of the polymeric film coatings (and changes thereof upon exposure to the release media), also the type of core can play a major role.13 For example, Wesseling and Bodmeier14 showed significantly different drug release patterns from

10.1021/bm0500704 CCC: $30.25 © 2005 American Chemical Society Published on Web 04/22/2005

pH-Sensitive Polymer Blends

polymer-coated pure drug matrixes and drug-layered sugar cores, despite the identical composition of the macromolecular shells. The mechanisms controlling drug release from polymercoated beads are complex and not yet fully understood.15,16 Several processes can be involved, such as water-imbibition, drug dissolution and drug diffusion through the intact film coatings and/or water-filled cracks. For example, Ozturk et al.17 showed that phenylpropanolamine hydrochloride release from ethylcellulose-coated beads was controlled by a combination of drug diffusion and osmotic effects. The importance of crack formation due to the creation of significant hydrostatic pressure within different types of polymer-coated drug delivery systems has recently been pointed out in a comprehensive review by Stubbe et al.18 So-called “pulsatile drug release” is achieved when the intact polymer coating is (almost) impermeable for the drug, resulting in a “lag phase” with negligible drug release, followed by rapid and complete drug release upon shell bursting. This type of release patterns is of great interest for certain drug therapies and dosing regimes (e.g., substitution of hormones with circadian fluctuations and simulation of night dosing). As early as 1975, this type of advanced drug delivery systems has been described.19 A particularly interesting system has recently been proposed: “degradation controlled exploding microcapsules”.20,21 The idea is to coat a biodegradable gel (containing the drug) with a polymeric membrane being permeable only for water, not for the drug and gel degradation products. Upon exposure to biological fluids, water enters the system and the macromolecules of the gel are degraded into osmotically active small molecules. As soon as the hydrostatic pressure developed within the microcapsules is sufficient to rupture the polymer coating, the drug is rapidly and completely released. In the present study, different types of drug-loaded cores (approximately 0.8 mm in diameter) were coated with blends of ethylcellulose and Eudragit L (a methacrylic-acid-ethylacrylate copolymer). To facilitate oral administration, these beads can be either filled into hard gelatin capsules or compressed into tablets. Upon contact with the contents of the gastrointestinal tract, the gelatin capsules rapidly dissolve/ tablets disintegrate, releasing the beads. Once in contact with biological fluids, the macromolecular shells of the beads interact with their environment and control the subsequent release of the drug into the human body. To guarantee a secure medical treatment, it is important to know which processes are involved in the control of drug release. As pointed out above, these phenomena are complex and yet not fully understood. This is especially true for film coatings which do not consist of one single polymer but of pHsensitive polymer blends. The major aims of the present study were (i) to coat theophylline-layered sugar cores and pure theophylline matrixes with pH-sensitive polymer blends consisting of ethylcellulose (water-insoluble) and Eudragit L (waterinsoluble below and water-soluble above pH 5.5), (ii) to study the effects of the type of core and polymer blend ratio on the resulting drug release profiles, and (iii) to get further insight into the underlying mass transport mechanisms.

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Experimental Section Chemicals. Theophylline (Boehringer Ingelheim, Ingelheim, Germany) (melting point: 270-274 °C, solubility: 15.4 and 12.0 mg/mL in 0.1 M HCl and phosphate buffer pH 7.4 at 37 °C, respectively22), spherical theophylline matrixes (710-850 µm, 94% drug content; Klinge Pharma GmbH, Munich, Germany), spherical sugar cores (Suglets sugar spheres NF; 710-850 µm; NP Pharm S. A., Bazainville, France) (melting point of sucrose: 160-186 °C, solubility: 2 g/mL at 20 °C), Aquacoat ECD (aqueous dispersion of ethylcellulose, EC; FMC c/o Interorgana, Cologne, Germany), Eudragit L30D-55 (aqueous dispersion of a methacrylic acid-ethyl acrylate copolymer 1:1; Ro¨hm, Darmstadt, Germany), triethyl citrate (TEC; Morflex, Greensboro, NC), hydroxypropyl methylcellulose (HPMC, Methocel E5; Colorcon, Orpington, U.K.), and poly(ethylene glycol) (PEG 4 000; BASF, Ludwigshafen, Germany) were used as received. Preparation and Characterization of Thin, Polymeric Films. Aqueous polymer dispersions were plasticized overnight with triethyl citrate (25% w/w; based on the total polymer mass) and adjusted to 15% (w/w) polymer content with purified water. Films were prepared by spraying pure EC or Eudragit L dispersions or blends thereof (75:25, 50: 50, 25:75 w/w) onto Teflon plates and subsequent curing at 60 °C for 24 °h (spray gun: Pilot Mini, Walther Spritz- und Lackiersysteme GmbH, Wuppertal, Germany; the Teflon plates were heated to 37 °C with an IR lamp during spraying). The thickness of the films (approximately 40 µm) was measured using a thickness gauge (Minitest 600; Erichsen, Hemer, Germany). The films were cut into pieces of 5 × 5 cm, which were placed into 500 mL plastic containers filled with 400 mL of preheated phosphate buffer pH 7.4 (USP XXV), followed by horizontal shaking for 8 h (37 °C, 75 rpm, GFL 3033; Gesellschaft fu¨r Labortechnik, Burgwedel, Germany). At predetermined time intervals, samples were withdrawn (n ) 3), accurately weighed [wet mass(t)], and dried to constant mass at 60 °C [dry mass(t)]. The water content (%) and dry film mass (%) at time t were calculated as follows: water content (%)(t) )

wet mass(t) - dry mass(t) ×100% wet mass(t) (1)

dry film mass (%)(t) )

dry mass(t) ×100% dry mass(0)

(2)

The mechanical properties (puncture strength, % elongation, and energy at break) of the films were measured using the puncture test and a texture analyzer (TA.XT Plus; Winopal Forschungsbedarf GmbH, Ahnsbeck, Germany). Film specimens were mounted on a film holder. The puncture probe (spherical end: 5 mm diameter) was fixed on the load cell (5 kg) and driven downward with a cross-head speed of 0.1 mm/s to the center of the film holder’s hole. Load versus displacement curves were recorded until rupture of

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Figure 1. Effects of the EC:Eudragit L blend ratio (indicated in the figures) on the relative amount of theophylline released from coated beads: (a) in 0.1 M HCl, drug-layered sugar core; (b) in 0.1 M HCl, drug matrix; (c) in phosphate buffer pH 7.4, drug-layered sugar core; (d) in phosphate buffer pH 7.4, drug matrix (dotted curves indicate drug release controlled by diffusion through water-filled cracks at low pH).

the films (n ) 6) and used to determine the energy at break as follows: puncture strength )

F A

(3)

where F is the load required to puncture the film; A represents the cross-sectional area of the edge of the film located in the path elongationat break (%) )

xR2 + D2 - R R

×100%

(4)

Here, R denotes the radius of the film exposed in the cylindrical hole of the holder and D the displacement to puncture energyat break )

AUC V

(5)

where AUC is the area under the load versus displacement curve and V is the volume of the film located in the die cavity of the film holder (the energy at break is normalized to the film’s volume). Preparation and Characterization of Coated Beads. Two types of cores were studied: (i) theophylline matrixes

(94% w/w drug loading) and (ii) theophylline-layered sugar cores (10% w/w drug loading). The latter were prepared by layering a drug-binder solution (21.7% w/w theophylline, 1% w/w HPMC, 0.1% w/w PEG 4 000, 40.8% w/w ethanol, 36.4% w/w water) onto sugar cores in a fluidized bed (Kugelcoater UNILAB-05; Hu¨ttlin, Steinen, Germany). The process parameters were as follows: product temperature ) 40 ( 2 °C, spray rate ) 3-4 g/min, atomization pressure ) 0.4 bar, pressure of microclimate ) 0.2 bar, nozzle diameter ) 0.8 mm. The cores were coated with Aquacoat ECD, Eudragit L30D-55 and blends thereof (25:75, 50:50, 75:25 w/w) in a fluidized bed (Kugelcoater UNILAB-05). The respective aqueous polymer dispersions were plasticized overnight with triethyl citrate (25% w/w, based on the total polymer mass) and adjusted to 15% (w/w) polymer content with purified water prior to coating. The coating dispersions were sprayed onto mixtures of theophylline matrixes or theophyllinelayered sugar cores with placebo beads (1:4 w/w, 500 g) until a weight gain of 20% (w/w) was achieved (coating thickness: approximately 30 µm). The process parameters were as follows: product temperature ) 37 ( 2 °C, spray rate ) 3-4 g/min, atomization pressure ) 0.4 bar, pressure of microclimate ) 0.2 bar, nozzle diameter ) 0.8 mm. After



Figure 2. Effects of the type of core (drug-layered sugar core vs drug matrix; indicated in the figures) on the absolute amount of theophylline released per bead, coated with EC:Eudragit L blends in 0.1 M HCl (left column) and phosphate buffer pH 7.4 (right column). The polymer blend ratio is indicated in the figure. The horizontal lines indicate the respective absolute total amounts of theophylline to be released: 0.36 mg in the case of drug matrixes (thick lines), 0.03 mg in the case of drug-layered sugar cores (thin lines). Please note the discontinuous scaling of the y axis.

coating the beads were further fluidized for 15 min and subsequently cured for 24 h at 60 °C (final bead diameter: approximately 0.8 mm). Drug release from the coated beads was studied in 900 mL of 0.1 M HCl (with 0, 1 and 5% sodium chloride) and phosphate buffer pH 7.4 (USP XXVII), using the USP XXVII paddle apparatus (37 °C, 100 rpm, n ) 3). At

predetermined time intervals, 3 mL samples (not replaced with fresh medium) were withdrawn and analyzed UVspectrophotometrically (λ ) 271 nm; UV-2101 PC; Shimadzu Scientific Instruments, Columbia, MD). The swelling of the coated beads was determined by placing individual beads into glass vials filled with 0.1 M HCl (with 0, 1 and 5% sodium chloride), followed by

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Figure 3. Effects of the EC:Eudragit L blend ratio (indicated in the figures) on the swelling kinetics and dry weight loss behavior of thin polymeric films upon exposure to the different release media: (a) 0.1 M HCl, water contents; (b) phosphate buffer pH 7.4, water contents; (c) 0.1 M HCl, dry film mass; (d) phosphate buffer pH 7.4, dry film mass [results reproduced from ref 12].

horizontal shaking for 8 h (37 °C, 75 rpm; GFL 3033). At predetermined time intervals, the diameter of the beads was measured macroscopically using an optical imaging system (n ) 6). The increase in bead diameter (%) at time t was calculated as follows: increase in bead diameter (%)(t) ) diameter(t) - diameter(0) ×100% (6) diameter(0) Results and Discussion Drug Release from Coated Beads. The effects of the EC: Eudragit L blend ratio and type of core on the resulting relative drug release rate from coated beads in 0.1 M HCl and phosphate buffer pH 7.4 are shown in Figure 1. As it can be seen, theophylline release was significantly affected by the polymer blend ratio at low as well as at high pH, irrespective of the type of core. Interestingly, the release patterns were fundamentally different for drug-layered sugar cores compared to drug matrixes, despite of the identical composition of the polymeric coatings. Importantly, not only the slope, but also the shape of the release curves was affected, indicating changes in the underlying drug release mechanisms. Theophylline release was generally much

slower in the case of drug matrixes and followed zero order kinetics (at least over major parts of the observation period). In contrast, very different shapes of release curves were observed in the case of drug-layered sugar cores. For example, monotonically decreasing release rates, sigmoidal release patterns, and zero order kinetics resulted, depending on the polymer blend ratio. Generally, increasing Eudragit L contents led to increasing relative release rates, irrespective of the type of release medium and type of core. This can be attributed to the higher permeability of this polymer compared to EC and/or to its partial leaching out of the coatings at high pH.8,16 However, it has to be pointed out that only the relative amounts of drug released are illustrated in Figure 1. As the total absolute amounts of theophylline present in the coated beads are significantly different for drug-layered sugar cores and drug matrixes (0.03 vs 0.36 mg per bead), also the respective absolute release rates must be compared (Figure 2). The thick and thin horizontal lines indicate the total absolute amounts of theophylline to be released. Importantly, also the absolute release rate of theophylline was higher in the case of drug-layered sugar cores compared to drug matrixes (except for pure Eudragit L coatings), despite the higher drug amounts present within the systems. From a


mechanistic point of view, these absolute drug release rates are a priori to be considered when investigating the underlying drug release mechanisms. However, from a therapeutic point of view, the relative drug release rates are more important (in practice, the number of administered beads can be adjusted to provide the required total drug dose). To better understand the underlying drug release mechanisms, thin polymeric films of identical composition as the bead coatings were prepared and physicochemically characterized before and upon exposure to the different release media. Physicochemical Properties of Thin, Polymeric Films. Figure 3, parts a and b show the water uptake behavior of thin EC:Eudragit L films upon exposure to 0.1 M HCl and phosphate buffer pH 7.4, respectively. Generally, increasing Eudragit L contents led to increased water uptake rates and extents. This can be attributed to the higher hydrophilicity of this polymer compared to EC, to its higher swelling capacity and to its (partial) leaching out of the film coatings at high pH (being replaced by imbibing water). These swelling kinetics correlate very well with the observed theophylline rates from coated beads in the case of drug matrixes, at low as well as at high pH (Figures 1 and 2). With increasing water content, the permeability of the film coatings increases,16 leading to increasing relative and absolute drug release rates. At high pH, the increase in water content was more pronounced than at low pH (Figure 3) [due to the (partial) leaching of Eudragit L out of the films], resulting in higher drug permeabilities and, thus, higher relative and absolute drug release rates (Figures 1 and 2). In contrast, the water uptake kinetics of the polymeric films do not correlate with the observed theophylline release rates from coated beads in the case of drug-layered sugar cores. The monotonic increase in water content (and, thus, drug permeability) with increasing Eudragit L content cannot explain the observed significant changes in the shapes of the release curves. The effects of the EC:Eudragit L blend ratio on the dry weight loss behavior of thin polymeric films upon exposure to 0.1 M HCl and phosphate buffer pH 7.4 are shown in Figure 3, parts c and d, respectively. The observed limited dry weight loss at low pH can be attributed to the leaching of the water-soluble plasticizer triethyl citrate out of the systems, whereas the substantial dry weight loss at high pH indicates the (partial) leaching of Eudragit L out of the films. As expected, increasing initial contents of this polymer led to significantly increasing dry weight losses. These results correlate very well with the observed relative and absolute release rates of theophylline from the coated beads, irrespective of the type of core (Figures 1 and 2). The increasing loss in dry coating mass leads to increasing drug permeabilities and, thus, to increasing release rates. However, based on the swelling and dry weight loss kinetics of the polymeric films alone, the observed drug release kinetics from the investigated beads cannot be explained. To get further insight into the underlying mass transport mechanisms controlling drug release, also the mechanical properties of the film coatings (puncture strength, % elongation and energy at break) were measured (as a


Figure 4. Changes in the mechanical properties of thin EC:Eudragit L films upon exposure to 0.1 M HCl: (a) puncture strength; (b) % elongation at break; (c) energy at break. The polymer blend ratio is indicated in the figures [the results shown in Figure 4c are reproduced from ref 12].

function of the exposure time to the release media). Figure 4 shows the results obtained in 0.1 M HCl. Clearly, the mechanical resistance of the films rapidly increased at early time points (due to the plasticizing effect of imbibing water) and then leveled off (similar to the water uptake kinetics, Figure 3). The subsequent slight decrease in the mechanical resistance of the films can be explained by the leaching of the water-soluble plasticizer triethyl citrate into the release medium. Importantly, increasing Eudragit L contents led to significantly increasing energies required to break the films.

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Figure 5. Schematic presentation (not up to scale) of the underlying drug release mechanisms from the investigated theophylline-loaded beads: drug-layered sugar cores and pure drug matrixes, coated with EC and/or Eudragit L.

Upon exposure to phosphate buffer pH 7.4, the films became mechanically too weak to allow accurate measurements with the experimental setup used (except for pure EC films, which showed similar results as in 0.1 M HCl, data not shown). Drug Release Mechanisms. On the basis of the experimentally measured relative and absolute drug release rates (Figures 1 and 2), water uptake and dry weight loss kinetics of thin polymeric films of identical composition as the coatings (Figure 3) and changes in the mechanical properties of the latter upon exposure to the release media (Figure 4), the drug release mechanisms illustrated in Figure 5 can be hypothesized. At low pH (in the stomach), beads containing drug-layered sugar cores take up significant amounts of water due to the high solubility of sucrose (resulting in the creation of significant osmotic pressures within the bead cores, being the driving forces for the water influx). Consequently, a considerable hydrostatic pressure is exerted onto the polymeric coatings. If the latter are mechanically weak, cracks are formed (after certain lag-times) and the drug is rapidly released through water-filled channels. This type of drug release mechanism dominates in the case of films with moderate to high EC contents (Figure 1). In contrast, if the polymeric films are mechanically strong, they can withstand the hydrostatic pressure built up within the bead cores and crack formation is suppressed. In this case, drug release is primarily controlled by diffusion through the intact film coatings. This type of drug release mechanism is dominating in film coatings containing moderate to high Eudragit L amounts (Figure 1). (Remark: Due to the significant water imbibition, the diameter of the beads can be expected to increase, resulting in an increase in surface area available for diffusion and decrease in film thickness with time.)

On the other hand, if pure drug matrixes are used, the water influx is much lower than in the case of drug-layered sugar cores, because the solubility of theophylline in 0.1 M HCl is rather low (15.4 mg/mL at 37 °C).22 Thus, no significant hydrostatic pressures are built up within the bead cores and drug release is primarily controlled by diffusion through the intact film coatings, irrespective of the polymer blend ratio (Figure 1). In phosphate buffer pH 7.4, the leaching of the pHsensitive polymer Eudragit L plays a major role at moderate to high Eudragit L contents. In the case of pure Eudragit L coatings, the rapid dissolution of this polymer results in instantaneous drug release, irrespective of the type of bead core (Figure 1, parts c and d). In the case of drug-layered sugar cores, the leaching of Eudragit L out of the films (Figure 3d) together with the significant hydrostatic pressure built up within the bead cores results in a rapid disintegration of the coatings and, thus, high drug release rates (Figure 1c). In contrast, no polymer leaching occurs from pure EC coatings and theophylline is released through water-filled cracks which are created after a certain lag-time due to the significant water influx (Figure 1c). On the other hand, in the case of pure theophylline matrixes, the hydrostatic pressures within the bead cores are much lower, and despite the (partial) Eudragit L leaching from the coatings, the latter do not disintegrate (at least at high EC contents). Thus, theophylline is released through the more or less porous film coating remnants, explaining the much lower release rates compared to coated drug-layered sugar cores. In the case of pure EC coatings, the polymeric shells remain intact and (due to their low permeability for theophylline) effectively suppress drug release during the observation period (Figure 1d). Increase in Bead Diameter. To confirm the hypothesized drug release mechanisms, changes in the beads’ diameter upon exposure to 0.1 M HCl were monitored for both types



Figure 6. Swelling behavior of theophylline-loaded, EC:Eudragit L-coated beads in 0.1 M HCl with different types of cores: (a) druglayered sugar cores; (b) drug matrixes. The polymer blend ratio is indicated in the figures.

of cores (Figure 6, parts a and b). As long as the film coating does not rupture and an excess of a freely water-soluble substance is present within the system (as in the case of druglayered sugar cores), the imbibing water results in a monotonic increase in bead diameter. However, as soon as crack formation sets in, the hydrostatic pressure built up within the cores “squeezes” parts of the bead content out into the release medium. Consequently, the bead diameter levels off and partially decreases. As it can be seen in Figure 6a, this leveling off occurs after approximately 0.5, 3, and 4 h (indicated by the black arrows) in the case of 100:0, 75: 25, and 50:50 EC:Eudragit L blends. These time points agree very well with the observed significant increases in the drug release rate shown in Figure 1a. In contrast, no leveling off of the bead diameter was observed in the case of 25:75 and 0:100 EC:Eudragit L blends (Figure 6a), agreeing well with the absence of sudden increases in the respective drug release rates (Figure 1a). On the other hand, the bead diameter remained about constant during the observation period in the case of coated drug matrixes (Figure 6b), due to the absence of freely watersoluble substances within the systems. Consequently, the film

Figure 7. Effects of the osmolality of the release medium (indicated in the figures; pH 1.2) on theophylline release from drug-layered sugar cores coated with EC:Eudragit L blends. The polymer blend ratio is indicated in the figures.

coatings stayed intact, and drug release was primarily controlled by the (low) permeability of the polymer blends (Figure 1b). Only in the case of pure Eudragit L coatings, remarkable theophylline release was observed. Effects of the Osmolality of the Release Medium. To further confirm the hypothesized drug release mechanisms

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(by adding 1 and 5% NaCl) led to a significant decrease in the theophylline release rates. This can be explained by the decrease in osmotic pressure (being the driving force for water imbibition) and, thus, delayed onset of crack formation in the case of 100:0, 75:25, and 50:50 EC:Eudragit L blends (Figure 7). The solid lines indicate the time periods in which drug release is primarily controlled by diffusion through the intact film coatings; the dotted lines indicate the time periods, in which drug release is controlled by diffusion through water-filled cracks. Clearly, the onset of crack formation (indicated by the black arrows) is shifted to later time points with increasing NaCl amounts. The observed drug release profiles correlate very well with the experimentally measured increase in bead diameter (Figure 8). The leveling off of the latter corresponds to the onset of crack formation in all cases. Interestingly, the critical % increase in bead diameter, at which crack formation starts, is independent of the added amount of NaCl but characteristic for each type of film coating (approximately 6, 20 and 45% for 100:0, 75:25 and 50:50 EC:Eudragit L blends). This is because each film coating has a specific mechanical resistance, rupturing at a defined hydrostatic pressure. In addition, the observed critical threshold values for the % increase in bead diameter are of the same order of magnitude as the experimentally measured % elongation at break of thin free films (7, 21, and 65% elongation for 100:0, 75:25, and 50:50 EC:Eudragit L blends, respectively). This clearly indicates that the physicochemical properties of the thin free films are a good measure for the properties of the bead coatings. In the case of 50:50 EC:Eudragit L blends and 5% NaCl addition, the lag-time for crack formation was delayed for more than 8 h (Figure 7). However, looking at the increase in diameter of the respective beads in Figure 8, it becomes evident that the critical value of around 50% increase in bead diameter for this type of film coating will probably be attained after 9-10 h. In the case of 25:75 and 0:100 EC: Eudragit L blends, no leveling off of the bead diameters was observed for any of the investigated NaCl concentrations (Figure 8), explaining the absence of sudden increases in the theophylline release rate from the coated beads (Figure 7). Due to the decrease in osmotic pressure within the bead cores with increasing NaCl concentration in the release medium, the water influx slowed and, thus, the increase in the surface area of the film coatings and decrease in coating thickness slowed. As drug release from these systems is primarily controlled by diffusion through the intact polymeric films, both effects resulted in a decrease in the drug release rates (Figure 7). Figure 8. Effects of the osmolality of the release medium (indicated in the figures; pH 1.2) on the swelling behavior of theophylline-layered sugar cores coated with EC:Eudragit L blends. The polymer blend ratio is indicated in the figures.

(Figure 5), the osmolality of the release medium was varied by adding different amounts of NaCl. The resulting theophylline release patterns and changes in bead diameter upon exposure to 0.1 M HCl are illustrated in Figures 7 and 8, respectively (coated drug-layered sugar cores). Clearly, an increase in osmolality from 0.19 to 0.54 and 1.78 osm/kg

Conclusions Theophylline release from beads coated with EC and Eudragit L blends is strongly affected by the type of core (drug-layered sugar core vs drug matrix). Importantly, not only the slope, but also the shape of the release curves is altered (despite of the identical composition of the coatings), indicating changes in the underlying mass transport mechanisms. The observed phenomena can be explained based on the physicochemical properties of the film coatings and


the swelling behavior of the beads upon exposure to the release media. Both, the water influx into the systems as well as the mechanical resistance of the polymeric films are decisive at low pH. Due to the low solubility of theophylline, the water influx is low in the case of pure drug matrixes, and drug release is primarily controlled by diffusion through the intact coatings. In contrast, theophylline-layered sugar cores create considerable hydrostatic pressures within the beads upon exposure to 0.1 M HCl. Mechanically weak films (moderate to high EC contents) cannot withstand these pressures, resulting in crack formation and subsequent rapid drug release through water-filled channels. In contrast, mechanically strong films (high Eudragit L contents) withstand the hydrostatic pressures and drug release is controlled by diffusion through the intact coatings. The experimentally measured changes in the bead diameter confirm these drug release mechanisms. The leveling off of bead swelling agrees very well with the onset of crack formation. Furthermore, the addition of increasing amounts of NaCl to the release medium (leading to decreasing osmotic pressures within the cores and, hence, decreasing water influx rates) results in delayed onsets of leveling off of the bead diameter, crack formation and increase in release rate. At high pH, also the leaching of the pH-sensitive polymer into the release medium plays a major role, affecting the integrity and mechanical resistance of the film coatings. An interesting practical application of the obtained new knowledge on the underlying mass transport mechanisms is the possibility to help to facilitate the development and optimization of this type of controlled drug delivery system. References and Notes (1) Langer, R. Acc. Chem. Res. 1993, 26, 537-542. (2) Peppas, N. A.; Wright, S. L. Macromolecules 1996, 29, 8798-8804.

Biomacromolecules, Vol. 6, No. 4, 2005 2083 (3) Franssen, O.; Vandervennet, L.; Roders, P.; Hennink, W. E. J. Controlled Release 1999, 60, 211-221. (4) Torres-Lugo, M.; Peppas, N. A. Macromolecules 1999, 32, 66466651. (5) Stubbe, B. G.; Horkay, F.; Amsden, B.; Hennink, W. E.; De Smedt, S. C.; Demeester, J. Biomacromolecules 2003, 4, 691-695. (6) Amighi, K.; Moes, A. J. Drug DeV. Ind. Pharm. 1995, 21, 23552369. (7) Wakerly, Z.; Fell, J. T.; Attwood, D.; Parkins, D. Pharm. Res. 1996, 13, 1210-1212. (8) Lecomte, F.; Siepmann, J.; Walther, M.; MacRae, R. J.; Bodmeier, R. J. Controlled Release 2004, 99, 1-13. (9) Wu, C.; McGinity, J. W. Pharm. DeV. Technol. 2003, 8, 103110. (10) Amighi, K.; Timmermans, J.; Puigdevall, J.; Baltes, E.; Moes, A. J. Drug DeV. Ind. Pharm. 1998, 24, 509-515. (11) Lecomte, F.; Siepmann, J.; Walther, M.; MacRae, R. J.; Bodmeier, R. Pharm. Res. 2004, 21, 882-890. (12) Lecomte, F.; Siepmann, J.; Walther, M.; MacRae, R. J.; Bodmeier, R. Pharm. Res., in press. (13) Tang, L.; Schwartz, J. B.; Porter, S. C.; Schnaare, R. L.; Wigent, R. J. Pharm. DeV. Technol. 2000, 5, 383-390. (14) Wesseling, M.; Bodmeier, R. Pharm. DeV. Technol. 2001, 6, 325331. (15) Dresssman, J. B.; Palsson, B. O.; Ozturk, A. G.; Ozturk, S. S. Mechanisms of release from coated pellets. In Multiparticulates Oral Drug DeliVery; Ghebre Sellassie, I., Ed; Marcel Dekker: New York, 1997; pp 285-306. (16) Lecomte, F.; Siepmann, J.; Walther, M.; MacRae, R. J.; Bodmeier, R. J. Controlled Release 2003, 89, 457-471. (17) Ozturk, A. G.; Ozturk, S. S.; Palsson, B. O.; Weatley, T. A.; Dressman, J. B. J. Controlled Release 1990, 14, 203-213. (18) Stubbe, B. G.; De Smedt, S. C.; Demeester, J. Pharm. Res. 2004, 21, 1732-1740. (19) Baker, R. W. U.S. Patent 3,952,741, 1975. (20) Stubbe, B. G.; Braeckmans, K.; Horkay, F.; Hennink, W. E.; De Smedt, S. C.; Demeester, J. Macromolecules 2002, 35, 25012505. (21) Stubbe, B. G.; Horkay, F.; Amsden, B.; Hennink, W. E.; De Smedt, S. C.; Demeester, J. Biomacromolecules 2003, 4, 691-695. (22) Bodmeier, R.; Chen, H. J. Pharm. Sci. 1989, 78, 819-822.

BM0500704

pH-Sensitive Polymer Blends used as Coating Materials to Control

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