Article pubs.acs.org/IECR
Microspheres of Gelatin and Poly(ethylene glycol) Coated with Ethyl Cellulose for Controlled Release of Metronidazole Keerti V. Phadke,† Lata S. Manjeshwar,*,† and Tejraj M. Aminabhavi‡ †
Department of Chemistry, Karnatak University, Dharwad, 580 003, Karnataka, India SET’s College of Pharmacy, Dharwad, 580 002, Karnataka, India
‡
ABSTRACT: A novel type of metronidazole (MN) loaded ethyl cellulose (EC) coated blend microspheres of gelatin (GE) with poly(ethylene glycol) (PEG) were prepared by the water-in-oil (w/o) emulsion method and cross-linked with glutaraldehyde (GA) to achieve an encapsulation efficiency of 68%. The microspheres were characterized by Fourier transform infrared spectroscopy (FTIR) to understand the chemical interactions between the drug and the polymer matrix. Differential scanning calorimetry (DSC) was used to investigate the molecular level drug dispersion into the polymer matrix as well as surface coating onto the microspheres, while scanning electron microscopy (SEM) was employed to investigate the surface morphology. The diffusion coefficient, Dv, and solvent front velocity, u, based on volume expansion of the microspheres were calculated from the dynamic swelling data of the matrixes to obtain insights into the transport phenomenon that showed a non-Fickian trend. In vitro release of MN was dependent on the extent of hydrophobic coating.
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CR formulations13,14 in addition to acting as a drug release modifier.15−18 Gelatin (GE) is a natural protein, obtained by thermal degradation of collagen from the animal skin, bones, and rarely from fish scales. It mainly contains residues of three amino acids, viz., glycine, proline, and 4-hydroxyproline in its structure.19 The extended left-handed proline helix conformation incorporated with 300−4000 amino acids results in the formation of a stronger gel.20 Being a biocompatible, biodegradable, and edible polymer, GE is an ideal material for developing CR formulations for extending the release time of short-acting drugs. It is insoluble at ambient temperature but dissolves rather rapidly at body temperature in aqueous media, making it somewhat difficult to prepare long-lasting drug delivery systems.21 This problem may possibly be solved by coating GE microspheres with a hydrophobic polymer such as ethyl cellulose (EC).22 Ethyl cellulose is chosen as the costeffective polymer in the present study. Metronidazole (MN) used in this study is the preferred drug for treatment of intestinal amoebiasis and is well absorbed orally. It has a plasma elimination half-life of 6−7 h. The administration of MN in the conventional type of tablet dosage forms provides a minimum amount of MN for local action in the colon, resulting in the relief of amoebiasis, but with the unwanted systemic effects.23 Timed release MN-loaded formulations are, therefore, a better option to overcome the limitations of conventional dosage forms. Application of a hydrophobic coat onto the hydrophilic core microspheres would help to delay the release of MN in the small intestine while targeting the colon.24 Numerous matrix-type and polymer coated formulations have been widely explored for the CR of
INTRODUCTION Over the past decades, a variety of controlled release (CR) systems have been investigated in order to decrease dosing frequency, enhance patient compliance, reduce the plasma concentration fluctuation in drug levels, and facilitate more uniform release of drug over an extended period of time. Such formulations that are quite convenient to deliver the drug in a more controlled and predetermined manner than the conventional type of dosage formulations are needed to maintain the therapeutically required plasma concentration of the drug during its systemic circulation for a prolonged time.1,2 Such advantages are successfully met by using the biocompatible and biodegradable polymers, in which the active pharmaceutical ingredient can be either physically dispersed or covalently bonded to the polymer backbone, such that the release of drug occurs by diffusion mechanism through the polymer wall.3 Of all the methods of developing CR systems, microspheres have emerged as the most attractive formulation stategies,4 as these provide many advantages over the CR tablets in view of their distribution following a predictable and controllable release kinetics.5 However, uniform distribution of drug in the gastrointestinal track (GIT) would result in a uniform drug absorption, thereby reducing the patient-to-patient variability during its administration through the oral route.6−8 Poly(ethylene glycol), a nontoxic and water-soluble polymer, is known to resist recognition by the immune system, exhibits a rapid clearance from the body system, and is an FDA approved biopolymer in a wide range of biomedical applications. The hydrogels prepared from PEG are excellent biomaterials9 and its blend with other polymers may act as a pore former, to facilitate the creation of interconnected pore networks for increasing the cumulative release of the drug. However, increased pore creation might reduce the hindrance of pore networks, resulting in an increased burst effect with an elevated drug release.10−12 The microspheres prepared from PEG are often used in combination with other polymers as the effective © 2014 American Chemical Society
Received: Revised: Accepted: Published: 6575
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MN,25 but no attempts have been made to formulate microspheres of the blend of GE with PEG coated with EC to achieve the CR of MN in order to overcome the burst release effect. The objective of the present work is to prepare EC coated blend microspheres of GE and PEG for the CR of MN. The drug-loaded microspheres were characterized by the physicochemical techniques, and their in vitro release profiles were investigated to understand the non-Fickian transport mechanism in terms of diffusion and release parameters.
the core matix.28 Microspheres (100 mg) were dispersed in 25 mL of ethyl acetate solution containing various amounts of EC (viz., 1:1 and 1:2; core:coat ratio) and 0.02% w/v of Span-80 and stirred for 10 min at 100 rpm. Subsequently, 50 mL of nhexane (nonsolvent) was poured at a rate of 1 mL/min, and the polymer solution was stirred for 60 min until the coating was complete. The coated microspheres were filtered, washed with an excess of n-hexane, and dried at the ambient temperature. Encapsulation Efficiency. Aqueous drug solutions were prepared by taking 0.2 mg/mL of MN, and the pH was adjusted by adding a required amount29 of 0.1 N HCl. Subsequently, the solution was equilibrated for 1 h and centrifuged to remove any polymer debris. The clear supernatant was analyzed by UV spectrophotometer (Secomam, model Anthelie, France) at a λmax value of 278 nm30 to calculate the % encapsulation efficiency (% EE) using
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EXPERIMENTAL SECTION Materials. MN was purchased from Loba Chemicals, Mumbai, India. PEG (MW ≈ 6000), GE, EC, analytical-grade GA solution 25% (V/V), ethyl acetate, n-heptane, n-hexane, and light liquid paraffin oil were all purchased from s.d. fine chemicals, Mumbai, India. Span-80 was purchased from Loba Chemicals, Mumbai, India, and the water used was doubledistilled and deionized. Other chemicals were used without further purification. Preparation of GE−PEG Core Microspheres. Blend microspheres of GE and PEG were prepared by the water-in-oil (w/o) emulsion cross-linking method.26 Here, 10 mL of 15% (w/v) polymer solution was first prepared by taking varying amounts of GE and PEG in double distilled deionized water. Different quantities of MN were dissolved in the above polymer blend solution to obtain a homogeneous solution by stirring, which was then emulsified by slowly adding into 50 mL of the mixture (4:1 v/v) of light liquid paraffin and n-heptane containing 1% (w/w) Span-80 under constant stirring at 400 rpm using a Eurostar high-speed stirrer (IKA labortechik, Staufen, Germany). After 10 min, different amounts (viz., 1.5, 3.0, and 4.5 mL) of GA containing 0.5 mL of 0.1 N HCl were added for cross-linking the matrix, and the matrix was stirred for another 2 h. The hardened microspheres were then filtered and washed with n-hexane to remove light liquid paraffin along with the residual Span-80. Microspheres were again washed with 0.1 M glycine solution and water simultaneously to remove the unreacted GA. Brady’s test for this solution was found to be negative, indicating the absence27 of unreacted GA and, thus, the safety of the developed formulations. Solid microspheres obtained were vacuum-dried at 40 °C for 24 h and stored in a desiccator before use. All formulations were prepared as per details given in Table 1. Coating of Blend Microspheres with EC. The coacervation phase separation method was used to prepare the coated microspheres of drug-loaded blend microspheres as
⎛ Weight of drug in microspheres ⎞ % Drug loading = ⎜ ⎟ × 100 Weight of microspheres ⎝ ⎠ (1)
% Encapsulation efficiency ⎛ Actual drug loading ⎞ =⎜ ⎟ × 100 ⎝ Theoretical drug loading ⎠
Fourier Transform Infrared Spectral (FTIR) Studies. FTIR spectra of GE, PEG, EC, uncoated placebo microspheres, pristine MN, MN-loaded uncoated microspheres, and MNloaded coated microspheres were all taken using a Nicolet (Model Impact 410, Milwaukee, WI, USA) instrument in KBr pellets over a scanning range of 4000−500 cm−1 to confirm the formation of cross-linked blend microspheres. Differential Scanning Calorimetric (DSC) Studies. DSC (Rheometric Scientific, Surrey, U.K.) was employed to detect any possible polymorphic changes in drug as well as its chemical interactions with polymers and excipient.5 Samples of pristine MN, powdered uncoated placebo (F0) microspheres, drug-loaded uncoated (F2) microspheres, and drug-loaded coated (F8) microspheres were weighed directly in pierced aluminum pans (∼2 mg) and scanned between 25 and 400 °C at a heating rate of 10 °C min−1 under a static nitrogen gas pressure of 20 psi and at a gas flow rate of 20 mL/min. Particle Size Measurements. Particle size measurements were carried out using the calibrated ocular microscope under regular polarized light. Particle sizes were measured by considering a minimum of 100 particles using a dry sample adopter. The calculated volume mean diameter data are recorded in Table 2. Scanning Electron Microscopic (SEM) Studies. SEM images of uncoated and coated microspheres were obtained using (JEOL model JSM-8404, Japan) scanning electron microscopy. Samples were prepared by placing a group of drug-loaded microspheres on a carbon tape and residing on a stub, which was coated with a thin layer of gold. Equilibrium Swelling Studies. The extent of matrix swelling and attainment of equilibrium swelling have an effect on drug release characteristics. In order to measure equilibrium swelling, the microspheres were swollen in water for 24 h and removed from the test bottles. The excess surface-adhered water droplets were removed by blotting with the tissue paper wraps and weighed (±0.01 mg) on an electronic microbalance (Mettler, AT 120, Switzerland). Later, the microspheres were
Table 1. Formulation Parameters and Compositions of Respective Microspheres formulation codes
GEL (% w/w)
PEG (% w/w)
MN (% w/w)
GA (mL)
F1 F2 F3 F4 F5 F6 F7 F8 F9 F0
100 90 80 90 90 90 90 90 90 90
0 10 20 10 10 10 10 10 10 10
10 10 10 5 10 10 10 10 20 0
3 3 3 3 1.5 4.5 3 3 3 3
(2)
coating (core:coat)
1:1 1:2
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mL of dissolution medium previously maintained at 37 °C. Then, 5 mL of the sample aliquot was withdrawn at different time intervals and filtered through a 0.45 mm filter (Sartorius, Goettingen, Germany) to estimate the concentration of the drug (MN) using the spectrophotometer at a fixed λmax value of 278 nm. The dissolution medium was then replaced with 5 mL of fresh dissolution solution. From such triplicate measurements for each sample, the average values were calculated and used in data analysis and graphical presentation.
Table 2. Results of % Encapsulation Efficiency (% EE) from eq 2, Volume Mean Particle Size, % Equilibrium Water Uptake from eq 3, and the Values of n along with Correlation Coefficients r, from eq 8 for Various Formulations formulation code F1 F2 F3 F4 F5 F6 F7 F8 F9 F0
% EE 62 61 53 66 57 68 61 61 58
volume mean particle size (μm)
% equilibrium water uptake
42 38 35 37 49 31 58 69 44 34
109 272 307 223 357 227 168 145 344 219
n
r
0.35 0.29 0.23 0.29 0.30 0.17 0.18 0.13 0.23
0.991 0.993 0.984 0.982 0.987 0.991 0.985 0.989 0.995
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RESULTS AND DISCUSSION An appropriate combination of core matrix and coating material (EC) is required to derive an optimized level of the best combination of formulation.32 PEG was used as a plasticizer33 to prepare the microspheres particularly in combination with hydrophobic polymer, such that the devised blend matrix will have increased hydrophilicity and after coating with a hydrophobic polymer such as EC would inhibit the burst release of the drug. Earlier,34 PEG was used in combination with other natural hydrophilic polymers like chitosan or its derivatives to increase mucoadhesivity as well as cumulative release of the drug. However, no attempts have been made to prepare the EC coated blend microspheres of GE with PEG using the water-in-oil emulsion method in the earlier literature. Further coating of these microspheres by EC is a unique formulation to study the release of MN. Earlier in the literature, Kong et al. developed the blend microspheres of GE with PEG that were chemically cross-linked, but this study did not involve drug release aspects except swelling aspects.35 In another study by Morita et al.,36 GE microparticles were prepared by colyophilization with PEG and the effect of the encapsulation efficiency of these particles into PLGA/PLA microspheres was investigated with no coating. However, our study is unique in that coating was achieved to reduce the burst release of the drug. The microspheres loaded with MN were prepared by the emulsion cross-linking method using GA as a cross-linker and coated with EC, by the coacervation phase separation method in order to avoid the possible burst release effect at the early stage of the drug release. The schematics of the preparation of the formulation is given in Figure 1. Encapsulation Efficiency. The % encapsulation efficiency (% EE) of MN, as listed in Table 2, ranges between 53 and 68%, which shows the dependence on formulation parameters such as % blend composition, extent of drug loading, and concentration of cross-linker. With increasing concentration of PEG in the blend matrix, the % EE decreased, since PEG might have facilitated the diffusion of drug particles into the outer aqueous phase, as it is known to form pores in the matrix. This
dried in an oven heated at 60 °C for 5 h until no change in the dry mass of the samples was achieved. The % equilibrium water uptake was calculated using % Equilibrium water uptake = ⎛ Mass of swollen microspheres − Mass of dry microspheres ⎞ ⎜ ⎟ × 100 Mass of dry microspheres ⎝ ⎠
(3)
Dynamic Swelling Studies. Molecular transport of aqueous media into the formulated microspheres was studied microscopically31 by monitoring the extent of swelling. The initial diameter of the dry microsphere was recorded by placing on a plate under the optical microscope. Then, by filling the container with double distilled deionized water, the change in diameter, Dt, with time due to water ingression was monitored until microspheres attained the equilibrium swelling as measured by its equilibrium diameter, D∞. Triplicate experiments were done from which the average values were considered for data analysis and graphical presentation. In vitro Drug Release Studies. In vitro release of drug from uncoated and coated blend microspheres was investigated at pH 1.2 initially for 2 h, followed by phosphate buffer of pH 7.4 until completion of dissociation. This was studied as a function of % drug loading, polymer blend composition, extent of cross-linking, and extent of surface coating. The USP apparatus-I dissolution tester (Dissotest, LabIndia, Mumbai, India) at a stirring speed of 100 rpm was used in the dissolution studies. Weighed quantities of samples equivalent to 10 mg of drug were placed in dissolution baskets and immersed in 500
Figure 1. Schematic representation of the formation of coated blend microspheres. 6577
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and 1468 cm−1, respectively, but OH stretching vibrations are observed with a broad peak at 3435 cm−1. The reaction between amino and carbonyl groups involving the formation of Schiff base is accompanied by color formation, called the Millard reaction.37 In the spectra of uncoated placebo microspheres, a new bond at 1648 cm−1 is observed, indicating the CN stretching vibrations of the imine group of Schiff base, thus confirming the cross-linking of the amino group of GE with the aldehydic group of GA. In order to investigate the possible drug−polymer interaction and to confirm the coating onto the core matrix, FTIR spectra of pristine drug, drug-loaded uncoated microspheres, pure EC, and drug-loaded EC coated microspheres are taken (see Figure 3). For nascent MN, the intramolecular hydrogen-bonded O
effect can be observed in F1, F2, and F3 formulations. As the concentration of drug loading increased from 5% (F4) to 20% (F9), the % EE decreased from 66 to 58%, due to the hydrophilic nature of MN. While preparing drug-loaded formulations, the encapsulated MN might have diffused out of the matrix that gave the low value of EE (58%). Additionally, the concentration of cross-linking agent showed a significant influence on % EE, as seen in Table 2. Thus, by increasing the concentration of GA, the % EE increased significantly to reach a maximum value of 68% when the formulation contained 4.5 mL of GA. At a high concentration of GA, the matrix has acquired rigidity and the coated formulations (F7 and F8) did not show any effect on % EE data, due to the hydrophobic nature of EC. It is to be noted that, for amebiasis treatment, 250 mg of MN trice daily is the dosage fixed for 5−10 days in adults. In order to achieve the equivalent dosage, nearly 4 g of the prepared formulations is required to maintain the therapeutic efficacy. FTIR Spectral Studies. FTIR spectra of GE, PEG, and blend microspheres displayed in Figure 2 confirm the formation
Figure 3. FTIR spectra of (a) pristine MN, (b) drug-loaded uncoated microspheres, (c) EC, and (d) drug-loaded coated microspheres. Figure 2. FTIR spectra of (a) GE, (b) PEG, and (c) uncoated placebo rnicrospheres.
H stretching vibration is observed at 3222 cm−1, while the alkene CH stretching vibration at 3097 cm−1 and alkane C H stretching vibration at 2846 cm−1 are prevalent. The sharp absorption bands observed at 1266 and 1664 cm−1 are attributed to CC and CN (imines) stretching vibrations, respectively. Symmetrical and asymmetrical stretching vibrations of NO (nitro group) are observed at 1367 and 1538 cm−1, respectively, while the bending vibrations of CH2 (methylene group) and CH3 (methyl group) are assigned for peaks at 1479 and 1429 cm−1, respectively. The peaks at 1186 and 957 cm−1 are assigned to CN stretching and CN bending vibrations, respectively.38 The FTIR spectrum of the drug-loaded but uncoated microspheres showed the bands that are observed for nascent MN, indicating the absence of
of blend microspheres. For GE, the band at 3365 cm−1 represents NH bond stretching, whereas N−H bending vibrations are observed at 1544 cm−1. The band at 1654 cm−1 is due to amide I (CO) stretching vibrations, while the CN bond stretching vibrations are observed at 1344 cm−1. Aliphatic CH asymmetric and symmetric stretching vibrations are assigned to the peaks observed at 2925 and 2855 cm−1, respectively. The bending vibrations of CH2 (methylene group) are observed at 1458 and 1398 cm−1, but the peak at 1111 cm−1 for PEG is due to the ether bond of PEG. Aliphatic CH stretching and bending vibrations are observed at 2886 6578
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polymer blend. The MN showed a sharp endothermic peak at 160 °C that corresponds to its melting point.39 The drugloaded microspheres showed the same peaks as those of placebo microspheres without significant changes, which suggests uniform and homogeneous dispersion of drug particles in the matrix with no significant interaction between the drug and the polymers. In the case of drug-loaded EC coated microspheres, all the DSC peaks observed for drug-loaded uncoated microspheres have appeared along with a small exothermic peak at 214 °C due to the loss of water molecules by the EC as a result of its hydrophobic nature.40 Particle Size. Volume mean particle sizes of the formulations range between 31 and 69 μm, depending on the parameters used in preparing the microspheres. As listed in Table 2, the concentration of cross-linking agent affected the particle sizes; i.e., for formulations with a high concentration of GA as in F6 (4.5 mL of GA), the volume mean particle size shows a decrease from 49 to 31 μm for F5 (1.5 mL of GA), indicating a reduction in particle size at a high concentration of GA. This could be because of the dense network structure of the blend, which reduces the particle size. The extent of drug loading has shown an effect on the particle size (Table 2); i.e., for formulation F4 containing 5% MN, the particle size is 37 μm, which was increased to 44 μm for F9 containing 20% MN. In preparing the microspheres, at high drug loading, the pores of the matrix, possibly created by the presence of PEG, are filled with drug particles to increase the particle size. Particle size is smaller at a high concentration of PEG in the blend, suggesting its dependence on the blend composition. Notice that the particle size of F1 (42 μm) is higher than that of F2 (38 μm), which in turn is higher than F3 (35 μm). This could be explained as being due to the formation of pores in the presence of PEG of the blend matrix, thereby showing a reduced volume mean particle size, but increased swelling at a higher concentration of PEG. Also, at a high concentration of EC as the surface coating material, the particle size also increased. The highest particle size observed in the case of F8 (69 μm) is the result of the highest concentration of EC used for coating onto the blend microspheres. Scanning Electron Microscopic (SEM) Studies. SEM images of (a) uncoated and (b) coated microspheres at 1000× magnification are displayed in Figure 5 for a group of drugloaded microspheres. In the case of uncoated microspheres, spherically shaped particle surfaces are smooth, but the surfaces of coated microspheres, even though they are smooth, show some fractured surfaces, due to the presence of EC. Some EC debris dents can be observed with the coated microspheres, because the EC might have been entrapped into the pores of the microspheres, thus hindering the diffusion of drug particles out of the matrix.41,42 Equilibrium Swelling Studies. The results of % equilibrium swelling data decrease from 357 to 227% with increasing concentration of GA from 1.5 to 4.5 mL (Table 2). At a high concentration of GA, the blend matrix is rigid due to high cross-link density. Also, at high drug loading up to 20% as in F9, the % equilibrium swelling increased to 344 due to the water-soluble nature of MN, which when it comes in contact with water diffuses out of the matrix, thereby facilitating the transport of water molecules into the blend matrix through the created free volume spaces. Another factor that influenced the % equilibrium swelling of the matrix is the composition of polymers in the blend. With increasing composition of PEG, say up to 20% (F3), the %
chemical interaction of the encapsulated MN with the blend matrix. Absorption bands of EC appearing at 3464 and 1316 cm−1 are due to O−H stretching and bending vibrations, respectively. Peaks at 2924 and 2871 cm−1 are due to aliphatic C−H stretching vibrations, while the bending vibrations are observed at 1450 and 1378 cm−1. A peak at 1114 cm−1 is due to the stretching vibrations of C−O−C. The drug-loaded and coated microspheres exhibit all bands that are present in drug-loaded microspheres and EC. A sharp peak at 2362 cm−1 and a broad peak at 3435 cm−1 are attributed to electrostatic interactions between EC and the uncoated microspheres, which confirms the coating of EC onto the surface of the microspheres. Differential Scanning Calorimetric (DSC) Analysis. DSC analysis was performed to study thermal transitions of the polymers during the heating process. Thermograms of the uncoated placebo microspheres, nascent drug, drug-loaded uncoated microspheres, and drug-loaded coated microspheres are displayed in Figure 4. Uncoated placebo microspheres exhibit three broad endothermic peaks, while the peak observed at 101 °C is due to endothermic transition as a result of the loss of moisture from the blend microspheres. Another broad peak observed at 291 °C is due to degradation of the blend polymer. A small endothermic peak observed at 193 °C is due to polymer chain interactions, resulting in a phase transition of the
Figure 4. DSC thermograms of (a) uncoated placebo microspheres, (b) pristine MN, (c) drug-loaded uncoated microspheres, and (d) drug-loaded coated microspheres. 6579
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Figure 6. Plot of Dt/D0 vs time t, demonstrating the effect of crosslinking density for formulations F5 (1.5 mL of GA), F2 (3 mL of GA), and F6 (4.5 mL of GA).
concentration of GA. Similar plots for microspheres loaded with different amounts of MN are displayed in Figure 7,
Figure 5. SEM images of (a) drug-loaded uncoated microspheres and (b) drug-loaded coated microspheres.
Figure 7. Plot of Dt/D0 vs time t, demonstrating the effect of % drug loading for formulations F0 (placebo microspheres), F4 (5% w/w MN), F2 (10% w/w MN), and F9 (20% w/w MN).
equilibrium swelling also increases to 307 due to the pore forming ability of PEG.43 During the preparation of microparticles, PEG branches out of the blend polymer solution to orient toward the internal and surrounding aqueous phase, thus forming a sponge-like structure. Proteins like GE have a hydrophilic nature and get attached to hydrophilic moieties like PEG, resulting in the porous structure in the microparticles. These pores would create additional free volume space in the matrix for easy water ingression, thus increasing % equilibrium uptake. However, the coated microspheres hinder the matrix swelling due to the hydrophobic nature of EC. With increasing extent of coating, the % equilibrium swelling decreased. Thus, the % equilibrium swelling for the uncoated microspheres that was originally 272 for F2 was reduced to 168 for F7, which further got reduced to 145 for F8. Such a reduction in equilibrium swelling is due to the hydrophobic nature of EC. Dynamic Swelling Studies. Dynamic swelling results are obtained by monitoring changes in diameter of the microspheres, Dt, as a function of time, t, at ambient temperature using the optical microscopy. Figure 6 shows the plot of normalized diameter, Dt/D0, vs t for formulations prepared with different extents of cross-linking agent. The normalized diameter decreases with increasing concentration of GA, due to the formation of a rigid network structure at a high
wherein the normalized diameter increases with increasing concentration of MN, because at higher drug loading more of the drug is possibly leached out of the matrix, leaving behind the free volume for the entry of water molecules. This might have enhanced the entry of excess water molecules into the matrix. This phenomenon also supports the low % EE data at high drug loading. The plot of Dt/D0 vs t for formulations containing different % composition of polymers displayed in Figure 8 suggests an increase in normalized diameter with increasing amount of PEG in the matrix, due to the increase in porosity of the overall matrix. The water transport takes place mainly through these porous channels. The increased porosity also enhances the water transport through the matrix to induce high swelling of the microspheres. However, hydrophobic coating by EC helps to significantly reduce the swelling of the coated microspheres. The plot of Dt/D0 vs t at varying amounts of EC coating is displayed in Figure 9, wherein the normalized diameter is found to decrease with increasing concentration of EC. From the swelling data, we have calculated the dimensional changes of the microspheres (i.e., volume changes) as a function of time, ΔVt, with respect to the initial volume, V0, of the microspheres. From these data, the diffusion coefficients, Dv, of water molecules were calculated using44 6580
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Solvent front velocity, u, of the advancing aqueous boundary into microspheres was also calculated using45 u=
⎛ dv ⎞ 1 ⎜ ⎟ ⎝ dt ⎠ A
(6)
The results of equilibrium normalized diameter, D∞/D0, diffusion coefficient, Dv, and solvent front velocity, u, presented in Table 3 suggest that diffusion coefficient and solvent front velocity values are higher for formulations with a higher PEG content of the blend, a higher amount of MN, a lower extent of cross-linking, and a lower extent of surface coating. Therefore, swelling of the microspheres depends on the availability of free volume in the microspheres. With increasing concentration of GA, the extent of cross-linking increased, leading to reduction in free volume, but with increasing composition of PEG, free volume also increased, leading to an increase in liquid transport through the polymer matrix. Dynamic swelling data of all the formulations were also fitted to the empirical equation used before.37
Figure 8. Plot of Dt/D0 vs time t, demonstrating the effect of % composition of the polymer for formulations F1 (pure GE), F2 (10% w/w PEG), and F3 (20% w/w PEG).
Dt = kt n D∞
Here, k is the rate constant and n represents the mode of transport. Least-squares estimated values of n at 95% confidence limit, using dynamic swelling data fitted to eq 7, are presented in Table 3 along with the results of correlation coefficients, r. The values of n range between 0.19 and 0.39, indicating the non-Fickian mode of transport of MN through the blend matrix. In vitro Release Studies. To understand the in vitro release of MN from uncoated and coated blend microspheres, the dissolution experiments were conducted in acidic (pH 1.2) as well as alkaline media (pH 7.4). For all the formulations, the release of MN was extended up to 32 h, with a delayed burst release due to the plasticizing activity46 of PEG. As seen from the results of % cumulative release vs time plots in Figure 10, at varying amounts of GA, but at fixed blend composition and of drug, the drug release was slower for F6 that contained a higher amount of GA (4.5 mL) but was faster for F5 that has a lower concentration of GA (1.5 mL). At a higher amount of GA, the matrix was rigid, which hindered the transport of MN and, thus, lowered its release. Figure 11 displays the release profiles for blend microspheres containing different amounts of MN. Here, the release was slower for F4 that contained a lower content of MN (5% w/w) but was faster for F9 formulation containing a higher amount of MN (20% w/w). This could be the result of a higher concentration gradient created in the blend matrix, leading to a quick diffusion of MN into the media.
Figure 9. Plot of Dt/D0 vs time t, demonstrating the effect of coating for formulations F2 (uncoated), F7 (core:coat, 1:1), and F8 (core:coat, 1:2).
⎛ ΔVt ⎞ ⎛ 4(ΔV∞/V0 ⎞⎛ Dv ⎞1/2 1/2 ⎜ ⎟=⎜ ⎟⎜ ⎟ t D0 ⎝ V0 ⎠ ⎝ ⎠⎝ π ⎠
(4)
where 2 ⎛ V0D0 ⎞ Dv = ⎜(1.773 × slope) ⎟ 4ΔV∞ ⎠ ⎝
(7)
(5)
Table 3. Water Transport Results for Various Formulations at 37 °C formulation codes
equilibrium normalized diameter (D∞/D0)
n (eq 7)
r (eq 7)
Dv (×106 cm2/s)
u (×104 cm/s)
F1 F2 F3 F4 F5 F6 F7 F8 F9 F0
5.57 4.58 4.25 4.85 4.67 3.83 2.81 2.30 4.43 5.20
0.39 0.34 0.34 0.36 0.35 0.29 0.32 0.19 0.34 0.37
0.925 0.987 0.983 0.961 0.908 0.975 0.984 0.909 0.933 0.926
3.68 3.47 4.58 4.70 5.93 2.73 1.96 1.47 6.06 3.57
4.79 5.00 5.77 4.99 6.98 4.86 3.49 1.00 5.14 4.37
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release of 100% of encapsulated drug was not observed in any formulation containing MN in the earlier literature. In order to achieve the maximum release of drug, PEG was chosen while preparing the formulations. Thus, the present blend microspheres have released maximum of 98% of MN, which is mainly attributed to the presence of PEG in the blend matrix, and this is the highest observed so far in the literature. Figure 13 depicts the dependence of MN release on the amount of surface coating, wherein we have observed that, as
Figure 10. Effect of cross-linking density on the in vitro release profile for the formulations F5 (1.5 mL of GA), F2 (3 mL of GA), and F6 (4.5 mL of GA).
Figure 13. Effect of coating on the in vitro release profile for formulations F2 (uncoated), F7 (core:coat, 1:1), and F8 (core:coat, 1:2).
the amount of coating increased, the % cumulative release also has significantly decreased. For F8, only 5% of MN was released in the first 2 h with no burst effect, due to the hydrophobic nature of EC as the coating material, which lowered the % cumulative release of MN from the coated blend microspheres. The advantage of EC coated microparticles is that they do not allow the burst release of the drug in the stomach but facilitate the drug release in the intestine after 3 or 4 h. From this study, the F7 formulation showed the most efficient release of MN and this formulation has a core:coat ratio of 1:1. In this case, both of the requirements were satisfied; viz., the burst release was minimized and 50% drug release occurred in 12 h (for one digestive cycle). Since the polymers used in the preparation of formulations are mucoadhesive in nature, they also reside for a longer time in the intestine. In order to establish a correlation between the % cumulative drug release and drug transport, the in vitro release data have been fitted to the empirical equation:47
Figure 11. Effect of % drug loading on in vitro release profile for formulations F4 (5% w/w MN), F2 (10% w/w MN), and F9 (20% w/ w MN).
To investigate the effect of PEG in the blend matrix, the % cumulative release data are plotted as a function of time for F1, F2, and F3 formulations in Figure 12, which suggests an increase in drug release with increasing content of PEG, due to its pore forming nature. Earlier,44 the formulation prepared from GE and NaCMC released the drug ketorolac tromethamine up to 80%. In another study,45 GE and HEC blend formulations released 80% of theophylline. However, the
Mt = kt n M∞
(8)
Here, Mt/M∞ represents the fraction of drug released at time t and k and n have the same meanings as indicated in eq 7. The n values estimated for all of the formulations, using the leastsquares method at the 95% confidence level, are given in Table 2 along with the correlation coefficients, r. The n values range from 0.13 to 0.35, indicating the non-Fickian transport mode. However, the diffusion coefficients and solvent front velocities are higher for formulations containing a higher amount of PEG, a higher amount of MN, a lower extent of cross-linking, and a lower extent of surface coating. A similar trend was also observed in MN release profiles, indicating a good correlation between drug release characteristics and transport parameters.
Figure 12. Effect of % composition of the polymer on the in vitro release profile for formulations F1 (pure GE), F2 (10% w/w PEG), and F3 (20% w/w PEG). 6582
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CONCLUSION The results of this study are novel to the literature that suggest the usefulness of biocompatible and cost-effective polymer like EC when used as a coating material on the blend microspheres of GE and PEG achieved the encapsulation of MN up to 68% and drug release was highest up to 98% and extending its release time up to 32 h. FTIR confirmed the formation of blend matrix, chemical stability of MN in the blend matrix, as well as surface coating onto the microspheres. DSC confirmed the molecular level drug dispersion and absence of drug−polymer interaction. Coating of the microspheres was also confirmed by an increase in particle size of the microspheres that ranged between 31 and 69 μm. SEM images supported the coating by showing the rough surfaces for the coated microspheres. Dynamic swelling and in vitro release data indicated a good correlation between in vitro drug release rates and transport parameters, suggesting the non-Fickian transport of MN through the microspheres.
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AUTHOR INFORMATION
Corresponding Author
*E-mail:
[email protected]. Phone: +91 836 2215286. Fax: +91 836 2771275. Notes
The authors declare no competing financial interest.
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