Controlled drug delivery through a novel PEG hydrogel encapsulated silica aerogel system Seda Giray, Tug˘ba Bal, Ayse M. Kartal, Seda Kızılel, Can Erkey Department of Chemical and Biological Engineering, Koc University, 34450 Sariyer, Istanbul, Turkey Received 30 May 2011; revised 2 December 2011; accepted 8 December 2011 Published online 28 February 2012 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.34056 Abstract: A novel composite material consisting of a silica aerogel core coated by a poly(ethylene) glycol (PEG) hydrogel was developed. The potential of this novel composite as a drug delivery system was tested with ketoprofen as a model drug due to its solubility in supercritical carbon dioxide. The results indicated that both drug loading capacity and drug release proﬁles could be tuned by changing hydrophobicity of aerogels, and that drug loading capacity increased with decreased hydrophobicity, while slower release rates were achieved with increased hydrophobicity. Furthermore, higher concentration of
PEG diacrylate in the prepolymer solution of the hydrogel coating delayed the release of the drug which can be attributed to the lower permeability at higher PEG diacrylate concentrations. The novel composite developed in this study can be easily implemented to achieve the controlled delivery of various drugs C 2012 Wiley Periodicals, and/or proteins for speciﬁc applications. V Inc. J Biomed Mater Res Part A: 100A: 1307–1315, 2012.
Key Words: PEG hydrogel, aerogel, supercritical CO2, ketoprofen
How to cite this article: Giray S, Bal T, Kartal AM, Kızılel S, Erkey C. 2012. Controlled drug delivery through a novel PEG hydrogel encapsulated silica aerogel system. J Biomed Mater Res Part A 2012:100A:1307–1315.
There is a growing need for efﬁcient and effective drug delivery vehicle systems to enhance release proﬁle, drug loading capacity, bioavailability, and selectivity. Composite materials are currently of interest, because they address the limitations of an individual material, while at the same time they enable researchers to beneﬁt from advantages of various materials. Recent advances in polymers and materials have improved design possibilities for nanostructured composites tremendously.1,2 Porous silica aerogels are sol–gel derived nanostructured materials with high surface areas, high pore volumes, and low densities.3 They are produced by supercritical drying of gels obtained via hydrolysis and condensation reactions of a silicon alkoxide precursor such as tetraethylorthosilicate (TEOS) in a solvent. Due to their tunable pore size and volume, nontoxic, and biocompatible character, aerogels have been receiving increased attention as suitable hosts for many applications including enzyme immobilization, biosensors, waste treatment, and drug release.4–8 Among these, drug-release studies demonstrated that high drug loadings can be achieved with aerogels and the drugs adsorbed on hydrophilic silica aerogels dissolve faster than the crystalline drugs. It was also proposed that hydrophilic aerogels can be used as carrier materials for oral delivery of drugs where immediate release is desirable. Also, due to their bio-
compatibility, their applications in pulmonary drug delivery and feasibility to apply as oral drug delivery devices have been investigated. As a result, they are targeted as potential delivery vehicles for various drugs including diclofenac,9 griseofulvin, miconazole, and ketoprofen.8,10–12 Among these, ketoprofen, 2-(3-benzoylphenyl)-propionic acid, is a hydrophobic, nonsteroidal anti-inﬂammatory drug (NSAID) that inhibits prostaglandin synthase function (Fig. 1). This type of NSAID is usually applied to remove the symptoms relating to rheumatoid artritis,12,13 osteoarthritis,14 and ankylosing spondylitis as well as widely used in pain relief and dysmenorrhea.15,16 For these diseases, there is a great effort to control sustained release of ketoprofen from various carriers such as poly(vinyl alcohol) nanoﬁbers,17 dendrimers,16 bioadhesive gels,15 and microparticles.18 Although dosage concentration ranges from 25 to 200 mg per tablet in clinics, in these applications, amount of the drug applied ranges from 4 to 285 mg.19–25 Due to its short shelf life, this drug requires frequent dosage when administered orally, which results in increased level of adverse effects such as gastrointestinal side effects (irritation, bleeding) and renal side effects. When it is applied transdermally, it faces the natural barrier, skin, which limits the penetration of the drug.16,17 In addition to aerogels, hydrogels are commonly used in tissue engineering and drug delivery as a matrix either by
Correspondence to: S. Kizilel; e-mail: [email protected]
or C. Erkey; e-mail: ce[email protected]
Contract grant sponsor: College of Engineering at Koc University in Turkey, TUBITAK; contract grant number: 107M326 Contract grant sponsor: Marie Curie FP7-IRG; contract grant number: 239471
C 2012 WILEY PERIODICALS, INC. V
FIGURE 1. Structure of ketoprofen.
themselves or as a part of a composite material for drugs, peptides, and proteins due to their three dimensional, hydrophilic, and tissue-like properties.26–33 Speciﬁcally, poly(ethylene) glycol (PEG)-based hydrogels have received signiﬁcant attention, because of their nontoxic, nonimmunogenic, and hydrophilic character. In previous studies, the kinetics of PEG hydrogel formation and diffusion of various drugs and/or proteins through various PEG-based hydrogel networks were investigated.15,34–38 For example, it has been observed that the release kinetics of proteins and drugs could be tuned by changing PEG chain length or concentration of functional PEG monomer.39 Also, the size or weight of the guest material within the hydrogel is an important parameter which affects hydrogel permeability. Another effective way to control the release of drugs or proteins from hydrogels is to make them responsive to external stimuli such as pH, temperature, and ionic strength.40 Under a proper stimulus, these responsive hydrogels can switch from a collapsed state to a swollen state, which then allows for the release of drugs/proteins encapsulated in PEG hydrogel network. It is a big challenge to control the release of drugs or proteins through a network to provide sustained and sequential release with a single material that can satisfy the intended application. For example, some materials may require targeted drug to possess a certain degree of hydrophobicity, while some other materials may have limited drug loading capacity. Therefore, there is an increasing trend in the use of composite materials as drug delivery devices.41 The use of multiple materials has the advantage that makes it possible to cap one material as an outer layer on top of the inner material which may act as a reservoir.42– 44 Thus, the capability to construct composite material with multiple layers of varied composition in a single device increases the range of applications and effectiveness of these systems. Aerogel–hydrogel composite structures, as will be presented in this study, may provide more effective drug/protein release when a single material cannot meet the requirements for the application. Furthermore, the proﬁles may be adjusted by varying the composition and/or concentration of the precursor in the PEG prepolymer solution, or by changing the hydrophobicity of the core aerogel. It may also be possible to design PEG hydrogel layer to be degradable in response to an external stimulus, such as pH34 and thus lead to faster release of the drug from the aerogel core. Here, we report a novel drug delivery vehicle synthesized by encapsulation of hydrophobic or hydrophilic aero-
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gels within PEG hydrogel via surface-initiated photopolymerization. This system can be applied for controlled delivery of drugs including ketoprofen as well as similar drugs as a part of an implant or oral tablet with a prolonged site-directed delivery and minimized side effects compared to the difﬁculties involved in its oral or transdermal administration.45 In the ﬁrst part of the article, details of the developed procedures to synthesize these novel composite materials are provided. Techniques developed to control the pore structure and hydrophobicity of the core aerogel material as well as the properties of the hydrophilic PEG hydrogel layer are also presented. In the second part, results on a drug delivery application of the synthesized composites are given. Speciﬁcally, the effects of hydrophobicity of aerogel and PEG concentration of the hydrogel coating of hydrophilic aerogels on the release rates of ketoprofen were examined. Ketoprofen (3-benzoyl-R-methylbenzeneacetic acid) was chosen as a model drug due to reasons explained previously. Furthermore, its relatively high solubility in supercritical carbon dioxide (scCO2) enabled loading of ketoprofen into aerogels by supercritical deposition. EXPERIMENTAL
Materials For the synthesis of silica aerogels, tetraethylorthosilicate (TEOS) (98.0%) and NH4OH (2.0M in ethanol) were purchased from Aldrich, HCl was purchased from Riedel-de Haen (37%). Ethanol was obtained from Merck (99.9%). For the surface modiﬁcation of aerogel, hexamethydisilizane (HMDS) was obtained from Alfa Aesar (98%). Model drug, ketoprofen (MW ¼ 254.30 g/mol) was obtained from Aldrich. For the hydrogel formation, eosin Y (98%), 1-vinyl 2-pyrrolidinone (NVP) (99þ%), poly(ethylene glycol) diacrylate (PEGDA) (MW ¼ 575 Da) were obtained from Aldrich. Triethanolamine (TEA) (>99.5%) was obtained from Fluka. Carbon dioxide (99.998 %) was purchased from Messer Aligaz (Istanbul, Turkey). The chemicals were used as received. Synthesis of silica aerogel Disk-shaped silica aerogels were synthesized by a two-step procedure using TEOS as precursor, HCl as hydrolysis catalyst, and NH4OH as condensation catalyst. A solution of TEOS (50 wt % in ethanol) was prepared. Subsequently, water and acid catalyst were added to start hydrolysis under continuous stirring. After about 30 min, the base catalyst was added to accelerate the condensation reaction and the sol was taken into cylindrical molds with a diameter of 1 cm for complete gelation [Fig. 2(a)]. The overall molar ratio of TEOS:water:HCl:NH4OH were kept constant at 1:4:2.44 103:2 102, respectively. After gelation occurred, the alcogels were taken out of the mold and placed in an aging solution (50 vol % ethanol and water), and left in an oven at 323.2 K for 20 h. The aim of aging step was to prevent shrinkage during supercritical drying by improving the mechanical strength of the alcogel. Next, the alcogels were kept for three more days in pure ethanol in order to remove water and all impurities [Fig. 2(b)]. After
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was prepared by placing a known amount of HMDS into the vessel within a glass vial along with glass beads which kept the glass vial intact and the vessel was charged with CO2. This vessel was also kept at 10.34 MPa at ambient temperature for 48 h until all HMDS was dissolved in CO2. The surface modiﬁcation of the hydrophilic aerogels was achieved by injection of CO2–HMDS mixture from the mixing vessel to the reaction vessel. This was realized by pressurizing the mixing vessel up to 20.68 MPa by charging with CO2 from the syringe pump. The CO2–HMDS mixture at 20.68 MPa and at room temperature was injected into the main vessel by opening the valve between reaction vessel and mixing vessel. This procedure was repeated until the pressure in the reaction vessel reached 20.68 MPa. The vessel was isolated at 20.68 MPa and at 333.2 K for the reaction of silica aerogel surface with HMDS for 30 min. After reaction, extraction of excess HMDS and other reaction byproducts was carried out using scCO2 at 10.34 MPa and at reaction temperature. Then, the vessel was depressurized at the same temperature and modiﬁed samples were obtained after the vessel was cooled. The surface modiﬁcation reactions were carried out at different ratios of HMDS/aerogel in scCO2 ranging from 0 to 4.2.
FIGURE 2. Scheme for synthesis of eosin Y functionalized hydrophobic aerogel formation and its subsequent coating within PEG hydrogel. [Color ﬁgure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
aging step, alcogels were contacted with eosin Y (2 mM in ethanol) solution. The adsorption of eosin Y on the surface of alcogel led to a reddish transparent composite of silica alcogel with eosin Y [Fig. 2(c)]. The alcogels with eosin Y were subsequently dried by scCO2 at 313 K and 10.3 MPa [Fig. 2(d)]. The resulting disk-shaped aerogels were hydrophilic and had a diameter of 1 cm and a thickness of 0.2 cm. Procedure for surface modiﬁcation of silica aerogels The experimental apparatus for the surface modiﬁcation of hydrophilic silica aerogels was explained in the previous study by Kartal and Erkey.46 A vessel with an internal volume of 54 mL and equipped with two sapphire windows for viewing the contents was used for surface modiﬁcation. Another vessel was used for mixing HMDS with CO2, and had a volume of 25 mL. For a typical experiment, a certain amount of hydrophilic silica aerogel sample (around 100 mg) was placed in the 54 mL vessel and the vessel was brought to a temperature of 333.2 K by circulating water using a circulating heater (Cole-Parmer polystat temperature controller). The vessel was then charged to 10.34 MPa with CO2 using a syringe pump (Teledyne ISCO model: 260D) and isolated at these conditions. The mixing vessel
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Supercritical deposition of the ketoprofen on silica aerogel A certain amount of ketoprofen and aerogel was placed in the 54 mL vessel described above. The vessel was heated until the temperature reached 333 K, and carbon dioxide was added until the pressure was 22 MPa. The system was kept at this condition and the contents were stirred for 48 h. After reaching adsorption equilibrium, CO2 was vented and loaded aerogels was taken out of the vessel and weighed. For all loading experiments in scCO2, the ratio of the mass of aerogel to the mass of ketoprofen placed in the vessel was kept constant at unity. Since the aerogel density affected drug loading capacity of the aerogel, the density of the aerogel was also kept constant at 0.2 g/cm3. The amount of loaded drug was calculated by taking the difference of the aerogel mass before and after loading procedure. Hydrogel coating of silica aerogels The hydrogel precursor solution was prepared with TEA (225 mM), 30% and 15% (w/v) PEG diacrylate (MW ¼ 575 Da), and NVP (37 mM). The solution was adjusted to pH 8 using 6M HCl. Precursor solutions were ﬁlter sterilized using a 0.2-lm syringe Teﬂon ﬁlter. Eosin and drug-loaded hydrophobic or hydrophilic aerogels were immersed in PEGdiacrylate prepolymer solution and photopolymerization was carried out using visible light (514 nm, ﬂux ¼ 5.2 mW/ cm2) for 3 min for each surface of the aerogels [Fig. 2(f)]. Immobilized eosin Y on the surface of the aerogel initiated the formation of PEG diacrylate hydrogel on the surface.47,48 Eosin was used as the photoinitiator since its spectral properties perfectly suit its application as an initiating system for an argon ion laser49–51 This step resulted in the formation of a cross-linked thin PEG hydrogel coating around the hydrophobic or hydrophilic aerogels.
Ketoprofen release experiments Drug release experiments were conducted at 310 K and under constant string at 100 rpm. The release medium was selected as HCl solution (0.1N) to mimic gastric ﬂuid.52 Each aerogel sample was placed in a small bag made out of ﬁlter paper and immersed into a glass vessel containing HCl solution (100 mL, 0.1N). Three samples were used per condition. The glass vessel was covered and placed on a stirrer which was kept inside an oven at 310 K. At speciﬁc time points, approximately 20 lL samples were taken and placed into 0.5 mL eppendorf tubes. The concentration of ketoprofen was determined using a spectrophotometer (Thermo Fisher Scientiﬁc Nanodrop 1000) at 260 nm for release experiments, due to the speciﬁc absorption of ketoprofen at this wavelength. In control samples, neither PEG hydrogel nor aerogels gave signiﬁcant absorbance at 260 nm, which proves that absorbance at 260 nm was the result of ketoprofen release (data not shown). Calculations for diffusion coefﬁcients The effective diffusion coefﬁcients for coated and noncoated aerogels were calculated using Fick’s law53,54: 1=2 Mi De t ﬃ2 Minf pd2
Here, Mi and Minf are the released concentrations of the drug at time i and at time inﬁnity, respectively. De is the effective diffusion coefﬁcient (cm2/min), t is time (min), and d is the half of the thickness of the aerogel (cm). This equation shows that De can be calculated from the plot of Mi/ Minf versus t1/2. In order to calculate, Mi, solute diffused at time i, a mass balanced was carried out as follows: Mi ¼ Ci V
where Ci is the concentration of the drug at time i, V is the total volume of the release solution (100 mL). Calculated De values are also normalized by using D0 which is the diffusion coefﬁcient of ketoprofen in water at 37 C (7.68 106 cm2/s).54,55 Contact angle determination Contact angle was measured by using a home-made device which included a camera connected to a computer. First of all, a 2 lL drop of water was put onto the aerogel surface using a micropipette and placed in front of the camera. After taking a picture of this image, angle between the water droplet and the surface was measured by set-squared ruler. For each surface, measurement was repeated more than three times and the average of them was reported as ﬁnal value. Nitrogen adsorption–desorption measurements Effects of eosin loading and surface modiﬁcation steps on the pore structure of the aerogel were investigated with the adsorption/desorption isotherms of nitrogen at 77 K (Micromeritics ASAP 2020 surface analyzer). Each sample was degassed at 573 K for at least 150 min before analysis. For hydrogel-coated aerogels, in order to separate hydrogel from the aerogel, they were left in the furnace at 323 K. As
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water evaporated from the hydrogel, it shrank and separated from the aerogel surface. The resulting pieces of the aerogel were used in the pore structure analysis. Statistical analysis The results of all data sets are analyzed using one-way analysis of variance (ANOVA). The results are represented as the mean value (6SD) of the triplicate samples unless otherwise stated. Differences between datasets are considered statistically signiﬁcant for p-values less than 0.05. RESULTS AND DISCUSSION
Functionalization and coating of silica aerogel within PEG hydrogel Visual examination of silica aerogels after surface modiﬁcation steps showed that the originally colorless transparent aerogels retained a red color due to the presence of eosin within the aerogel structure [Fig. 3(a,b)]. It is clearly observed from these ﬁgures that eosin molecules were homogeneously distributed throughout the aerogel. Uniform eosin distribution on the surface of aerogel is important as nonuniform distribution would compromise homogeneous coating of aerogel within PEG hydrogel. The penetration depth of the eosin molecules can also be adjusted by contact time of the ethanol solution with the aerogel. The hydrophobicity of aerogel after HMDS functionalization step was quantiﬁed by placing a water droplet, and measuring the contact angle on the surface of the aerogel [Fig. 3(c)]. For one of the conditions, it was found that the contact angle of water droplet for the eosin-modiﬁed hydrophobic aerogel was 128 [Fig. 3(c)]. Figure 3(d) shows the image of a PEG hydrogel-coated eosin functionalized hydrophobic aerogel. The thickness of the hydrogel coating was approximately 300 lm. Characterization of functionalized and hydrogel-coated silica aerogels Effects of eosin loading and surface modiﬁcation step on the pore structure of the aerogel were investigated. Pore size and pore size distribution were determined by nitrogen adsorption analysis. In Table I, BET (Brunauer–Emmett–Teller) speciﬁc surface area, Barrett–Joyner–Halenda (BJH) cumulative desorption pore volume, and average pore radius were compared for pure aerogel, eosin-loaded aerogel, hydrophobic aerogel, and hydrophobic aerogel after hydrogel coating. The results suggest that the presence of eosin on the aerogel surface caused BET speciﬁc surface area to decrease slightly with no signiﬁcant changes in the average pore diameter. Further modiﬁcation of the eosin-functionalized aerogel surface with HMDS decreased the surface area from about 820 to 529 m2/g, cumulative pore volume from 2.5 to 2.2 cm3/g, and increased the average pore radius from 6 to 6.7 nm. This could be attributed to the presence of some bottleneck type pores, which are blocked by SiA(CH3)3 groups after surface modiﬁcation reaction. However, pore volumes and surface areas are still high enough for carrier applications. Adsorption isotherms and pore size distributions of modiﬁed aerogels are compared in Figure 4(a,b). All
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FIGURE 3. Image of (a) pure aerogel, (b) eosin-doped hydrophilic aerogel, (c) water droplet on the eosin-doped hydrophobic aerogel, and (d) hydrogel-coated hydrophobic aerogel. Images are taken using a computer equipped with a Q Imaging Micropublisher 3.3RTV camera. [Color ﬁgure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
samples exhibited similar pore size distribution and typical type H1 type isotherms which indicate that the materials consist of compact agglomerates of approximately uniform spheres of silica and such a network is not disrupted by eosin loading, surface modiﬁcation, and PEG hydrogel coating. The BET speciﬁc surface area and pore size distribution measurements for noncoated and coated aerogels showed that the surface area of the hydrophobic aerogel did not change after encapsulation within PEG hydrogel. Also, isotherms and pore distributions of the two samples were nearly identical which indicated that hydrogel coating was only restricted to the external surface of the monolithic disks and water-based prepolymer solution did not penetrate through the hydrophobic pore structure of the aerogel before and during photopolymerization. These results indicate that the pore properties of aerogels during sol–gel synthesis do not change signiﬁcantly by various functionalization techniques used in this study.
Ketoprofen-loaded silica aerogels There are mainly two methods for loading of aerogels with drugs or proteins. One of these is supercritical deposition which involves placing the aerogel in a solution of the drug dissolved in scCO2. After adsorption of the drug on the silica aerogel surface, CO2 can be vented out leaving behind adsorbed drug on the internal surface with varying degrees of hydrophobicity. The second method involves contacting the gels with a solution of the drug in a solvent during one of the steps in the sol–gel synthesis before supercritical drying. The ﬁrst method is more favorable for drug delivery applications, since CO2 is nontoxic and leaves no residue in the medium. The drug may be desorbed from the surface during supercritical drying step with the second method. Therefore, in this study, drug was loaded into the aerogel structure from the supercritical CO2 phase. Ketoprofen has a high enough solubility in scCO2 (15.5 105 mole fraction at 331.5 K and 22 MPa) for supercritical deposition.
TABLE I. Variation of Pore Structure of Aerogel After Surface Modiﬁcation Steps Sample Pure aerogel Eosin-loaded aerogel After surface modiﬁcation After hydrogel coating
BET surface area (m2/g) 926 820 528 529
6 6 6 6
9 8 5 5
BJH cumulative pore volume (cm3/g) 2.90 2.500 2.10 2.20
6 6 6 6
0.03 0.025 0.02 0.02
BJH average pore radius (nm) 5.90 6.00 6.40 6.70
6 6 6 6
0.06 0.06 0.07 0.07
BET surface area of pure aerogel and eosin-loaded aerogel are different from the surface modiﬁed or hydrogel-coated aerogel samples (p < 0.05). BET surface area between the surface modiﬁed and hydrogel-coated aerogel samples do not signiﬁcantly differ from each other. BJH cumulative pore volume in all groups statistically differs from each other (p < 0.05). BJH average pore radius for pure aerogel and eosin-loaded aerogel do not signiﬁcantly differ from each other, whereas signiﬁcant differences (p < 0.05) are observed between the remaining groups.
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TABLE II. Results of Contact Angle and Drug Loading of Different Hydrophobic Aerogels Ratio of HMDS/aerogel (mg/mg) in scCO2 0 1.1 1.8 2.3 3 4.2
Contact angle 0 0 66 6 87 6 128 6 128 6
1 1 1 1
Mass percentage of drug loading to the aerogel mass (% w/w) 96 6 1 – 25.00 6 0.25 13.0 6 0.1 – 7.00 6 0.07
Mass percentage of drug loading to the aerogel mass (% w/w) signiﬁcantly differ from each other for all groups (p < 0.05).
FIGURE 4. Effect of eosin loading and surface modiﬁcation on (a) nitrogen adsorption isotherms and (b) pore size distribution.
Ketoprofen was loaded into aerogels with different degrees of hydrophobicity which were obtained by changing the ratio of HMDS/aerogel (mg/mg) in scCO2 during surface functionalization from supercritical solutions (SFFSS). Since increasing hydrophobicity is due to the replacement of surface hydroxyl groups on the silica surface by the Si(CH3)3 groups of HMDS, decreasing or increasing the amount of HMDS in scCO2 provides a control in the number of OH groups on aerogel surface. The results showed that the contact angle increased from 0 to 128 C as HMDS/aerogel (mg/ mg) ratio in scCO2 was increased from 0 to 4.2 (Table II, Fig. 5). This resulted in a decrease of the drug loading capacity from 96 to 7% (w/w) (Table II). These percentages
correspond to 40 mg and 5 mg total drug amount, respectively, for the disk-shaped aerogels used in this study. Previous studies have shown that the adsorption isotherm of ketoprofen on silica aerogel in the presence of scCO2 at constant temperature and pressure strongly depends on the hydrophobicity of silica aerogel.10 The ketoprofen loading obtained on a hydrophobic aerogel was lower than that of a hydrophilic aerogel, which was attributed to the decreased number of OH groups that provide active sites for hydrogen bonding with ketoprofen. This behavior is also observed in our studies and the results suggest that it is also possible to control drug loading by controlling the hydrophobicity of aerogel. It should be noted that a gradient in hydrophobicity was obtained through the core of the disk-shaped aerogels, when lower amounts of HMDS was used. This resulted in higher hydrophobicities at the regions closer to outer surface compared to regions closer to the center of the aerogel. This is due to the fact that the reaction between OH groups and HMDS starts at the outer surface of the aerogel and moves toward the center of the aerogel. Such a gradient may provide an additional parameter for release control. Effect of hydrophobicity on ketoprofen release behavior As shown in Figure 6, drug release rates varied with the degree of the hydrophobicity of the aerogel. As aerogels became more hydrophobic, release rate was decreased. For hydrophilic aerogels, release was completed nearly within 10 h [Fig. 6(a)]. However, for aerogel with a contact angle of 66 , release was completed in approximately 24 h. When hydrophobicity of the aerogel was increased further (up to a contact angle of 128 ), a slower release rate for ketoprofen was observed [Fig. 6(b,c)].
FIGURE 5. Image of water droplets on aerogels with different hydrophobicities. [Color ﬁgure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIGURE 7. Release behavior of ketoprofen from coated hydrogels. (a) Percent release, (b) fractional release in the region Mi/Minf < 0.6. (R2 ¼ 0.99 for hydrophilic, R2 ¼ 0.94 for coated (15%) and R2 ¼ 0.95 for coated (30%) aerogels).
FIGURE 6. Release behavior of ketoprofen from aerogels with different hydrophobicities. (h: contact angle) (a) and (b) percent release, (c) fractional release in the region Mi/Minf < 0.6 (R2 ¼ 0.98 for hydrophobic hydrogel (h ¼ 66 ), R2 ¼ 0.99 for hydrophobic (h ¼ 128 ) and R2 ¼ 0.99 for hydrophilic aerogels).
In addition, it was observed that the hydrophilic aerogel lost its disk shape during the drug release experiments and matrix erosion occurred, while all hydrophobic aerogels preserved their original shapes during the drug release experiments. The absence of matrix erosion was an indication that the release was governed by diffusion (Fig. 6).
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Ketoprofen release from coated and noncoated hydrophilic aerogels The ketoprofen-loaded aerogels were coated with PEG hydrogel layer by applying the same method for coating the aerogels with PEG hydrogel. Figure 7 shows ketoprofen release proﬁles from noncoated aerogel, and hydrogelcoated aerogels with two different PEG diacrylate concentrations; 15% and 30% by weight. As shown in Figure 7, during the ﬁrst 51 h (3060 min), only 55% of the drug was released through 30% PEG hydrogel-coated silica aerogel. For the case of 15% (w/v) coated aerogels, 80% of the drug was released within 72 h (4320 min). These results show that coating with PEG hydrogel can retard drug release from silica aerogels and the release rate is slower with the coating prepared using 30 wt % prepolymer solution. This is because, PEG hydrogel mesh size, which depends on the crosslink density, modulates the diffusion of the drug through the hydrogel membrane, and thus affects the release rate of the drug [Fig. 7(a,b)]. Hydrophilic aerogels crumbled when they were wetted in water which led to faster release of ketoprofen.8 Coating of aerogels with PEG hydrogel retarded this process. Thus, hydrogel coating
TABLE III. Summary of the Results for Effective Diffusivities of Different Hydrophobic Aerogels and Hydrogel-Coated Hydrophilic Aerogel Samples De (cm2/min)
Aerogel sample Hydrophilic Hydrohobic, (66 ) Hydrophobic, (128 ) 30% PEG hydrogel coated 15% PEG hydrogel coated
1.38 9.82 1.84 3.16 4.03
105 106 106 107 106
6 6 6 6 6
1.20 1.02 1.72 6.13 1.08
106 106 107 108 107
acts as not only as a barrier for the diffusion of the drug molecules, but also as a barrier which retards the crumbling of the hydrophilic aerogels. The release rate of a drug from a medium depends on different parameters such as diffusion coefﬁcient of the drug in the matrix, drug particle diameter, molecular weight of the drug, and drug solubility in the released medium. The results presented in this study are consistent with the fundamental fact that diffusion coefﬁcients of proteins or drugs are higher in membranes with higher permeability compared to the values observed through membranes with lower permeability.39,56,57 It is clearly observed that increasing hydrophobicity decreased drug release rates, thus lowered both diffusion rate and effective diffusion coefﬁcients (Table III). Furthermore, coating of aerogels retarded release of ketoprofen from the composite structure, since normalized effective diffusion coefﬁcients of ketoprofen with respect to its diffusivity in water through aerogel coated with 15% (w/v) and 30% (w/v) PEG hydrogel have been measured as 0.9 and 0.07%, respectively (Fig. 8). Such lower diffusivities could be attributed to the differences in diffusion mechanism of the drug through aerogel structure. Commonly, release of small molecular weight drugs or proteins involve diffusion of these compounds through waterﬁlled networks, such as hydrogels, however, in this system the release of ketoprofen involves both its desorption from aerogel scaffold and its diffusion through water-ﬁlled aerogel pores. As hydrophobicity of aerogel increase, it takes
FIGURE 8. Effective diffusion coefﬁcient normalized with D0, diffusion coefﬁcient in water at 37 C. Asterisk (*) indicates statistical difference.
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longer for water molecules to hydrate the pores of aerogel, and hence this causes delays in the release of drug through this composite structure. As a result, low values for the normalized effective diffusivities have been observed for ketoprofen through the composite structure developed.
In summary, we have developed a novel aerogel/PEG hydrogel composite, and examined the delivery of a model drug, ketoprofen, from this composite structure. The core-shell structure can be synthesized by coating of hydrophilic and/ or hydrophobic aerogels via surface-initiated photopolymerization of PEG hydrogel precursors. By incorporating the initiator, eosin Y, into aerogel structure, it was possible to coat aerogel with PEG hydrogel membrane. By adjusting the degree of hydrophobicity of the aerogel core and permeability property of the PEG hydrogel shell, it was possible to control the release proﬁle of ketoprofen. This composite could be easily implemented as a potential drug delivery vehicle to achieve sequential release of drugs with different hydrophobicities, such that hydrophobic and hydrophilic drug could be loaded into the aerogel core and PEG hydrogel shell, respectively. Moreover, the use of pH- or temperature-responsive PEG hydrogel in the composite structure may allow for the degradation of PEG hydrogel network to achieve faster release of a drug loaded within the aerogel core. ACKNOWLEDGMENTS
Authors thank Prof. Levent Demirel for allowing them to use goniometer for contact angle measurements.
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