RESEARCH ARTICLE – Pharmaceutics, Drug Delivery and Pharmaceutical Technology
Multi-arm PEG/Silica Hydrogel for Sustained Ocular Drug Delivery CHANGHAI LU, PAYAM ZAHEDI, ADAM FORMAN, CHRISTINE ALLEN Leslie Dan Faculty of Pharmacy, University of Toronto, Toronto, Ontario, Canada Received 2 June 2013; revised 20 September 2013; accepted 18 October 2013 Published online 27 November 2013 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jps.23777 ABSTRACT: In the present study, a series of sustained drug delivery multiarm poly(ethylene glycol) (PEG)/silica hydrogels were prepared and characterized. The hydrogels were formed by hydrolysis and condensation of poly(4-arm PEG silicate) using the sol-gel method. The relationships between water content in the PEG/silica hydrogel and stability as well as rheological properties were evaluated. Scanning electron microscopy analysis of the PEG/silica hydrogels revealed water content-dependent changes in microstructure. An increase in water content resulted in larger pores within the hydrogel, longer gelation time and higher viscosity. The PEG/silica hydrogels were loaded with dexamethasone (DMS) or dexamethasone sodium phosphate (DMSP), drugs that are hydrophobic and hydrophilic in nature, respectively. Evaluation of in vitro release revealed a zero-order release profile for DMS over the first 6 days, suggesting that degradation of the silica hydrogel matrix was the primary mechanism of drug release. It was also found that the drug-release profile could be tailored by varying the water content used during hydrogel preparation. In contrast, more than 90% of DMSP was released within 1 h, suggesting that DMSP release was only controlled by diffusion. Overall, results from this study indicate that PEG/silica hydrogels may be promising drug-eluting C 2013 Wiley Periodicals, Inc. and the American Pharmacists depot materials for the sustained delivery of hydrophobic, ophthalmic drugs. Association J Pharm Sci 103:216–226, 2014 Keywords: drug delivery systems; injectables; hydrogels; ophthalmic drug delivery; silica
INTRODUCTION Drug delivery systems that provide sustained release have the potential to revolutionize treatment in a number of indications, including ophthalmology.1–6 More than 90% of ocular drug formulations are in the form of eye drops, which are applied in volumes ranging from 50 to 100 :L to the precorneal area, with more than 75% of that volume being lost within the first 2–6 min.7 Indeed, following administration, it has been found that only 1%–5% of the drug reaches the desired site of action, making conventional eye drop-based formulations highly inefficient.8 In contrast, it has been demonstrated that significant drug levels can be achieved in tissues of the posterior segment of the eye using sustained drug delivery vehicles.9 In comparison to eye drop-based formulations, drug-eluting sustained delivery systems such as particulate systems10 , punctal plugs,11 contact lenses,3,12 and implants13,14 composed of biodegradable or nonbiodegradable polymers, injected and or implanted into the subconjunctival tissue and vitreous body, have the ability to enhance drug bioavailability and provide longer drug residence times.15 Therefore, these systems may be a favorable alternative to the conventional formulations commonly used to treat both anterior and posterior ailments of the eye. Despite significant research into the development of new depot technologies, only a handful of sustained-release systems have been approved for clinical use.15–17 Successful design of these systems is challenging given the number of criteria that must be met. Specifically, sustained-release depot systems must be biocompatible and biodegradable, ideally with degradation of the system occurring over the timescale of drug release.18–20 The system must also carry sufficient drug load
Correspondence to: Christine Allen (Telephone: +416-946-8594; Fax: +416978-8511; E-mail: [email protected]) Journal of Pharmaceutical Sciences, Vol. 103, 216–226 (2014)
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to deliver a therapeutic dose, and minimal initial burst release is often favored.21,22 In addition, for the majority of applications, injectable depot systems are less invasive and preferred over implantable systems that require invasive surgery for implantation.23 To date, a wide range of materials have been pursued for the preparation of depot delivery systems, with the most common being polyesters and polysaccharides.24–27 Silica xerogels have also been pursued for depot delivery applications owing to their biocompatibility, biodegradability, and their ability to provide sustained drug release.28–33 These hydrogels are prepared by a sol-gel process using alkoxysilicates, such as tetraethoxysilane (TEOS) and tetramethoxysilane. This process involves hydrolysis of silicate and the subsequent condensation of the silanol, either with itself or other polyols. For example, a number of groups have synthesized silica hydrogel networks using TEOS and linear poly(ethylene glycol) (PEG).34,35 Sugars, such as maltose and dextran, as well as glycerol, have also been combined with TEOS to produce hydrogel materials.36–39 Silica hydrogels are known to degrade in vivo to form monosilicic acid [Si(OH)4 ], which can be absorbed by the body and eliminated via the kidneys.40–43 The degradation of silica hydrogels under physiological conditions can be controlled by altering the waterto-alkoxide ratio, amount of solvent used during preparation, catalyst concentrations, and time of aging and drying.44,45 Thus, silica-based biodegradable matrices are ideally suited for use as depot drug delivery systems. Linear and multiarm PEG-based hydrogels have also been investigated as matrices for sustained drug delivery because of PEG’s versatility of functionalization and excellent biocompatibility.46–51 To the best of our knowledge, a combined PEG/silica hydrogel has yet to be evaluated as a depot drug delivery system for ophthalmic applications. In the current study, injectable PEG/silica hydrogels were prepared and their physicochemical and rheological properties were characterized. The ophthalmic drugs dexamethasone (DMS) and
RESEARCH ARTICLE – Pharmaceutics, Drug Delivery and Pharmaceutical Technology
Table 1. Material Components and Ratios Used in the Synthesis of the PEG/Silica Hydrogels Sample PEGS (g) H2 O (mL) TRISa (mL) Water Content (%)b
Rc
R1.5 R2.5 R3.5
1.5 2.5 3.5
0.1 0.1 0.1
0 0.1 0.2
0.15 0.15 0.15
60 71.4 77.7
a
pH value is 8.6. Water contents are calculated values based on the mass ratio of water and TRIS buffer to total the mass of water, TRIS, and PEGS in each hydrogel. c R values are generated from the mass ratio of water and TRIS buffer to the mass of PEGS. b
dexamethasone sodium phosphate (DMSP) were incorporated into the PEG/silica hydrogels, their in vitro release examined, and mechanisms of drug release elucidated. Overall, results from this study indicate that PEG/silica hydrogels may be promising drug-eluting depot materials for the sustained delivery of hydrophobic ocular drugs.
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prepared as described above using a solution containing DMSP (80 mg/mL) instead of distilled water. For preparation of DMS-loaded hydrogels, because of the limited aqueous solubility of this drug, DMS (4 mg) and PEGS (0.1 g) were first dissolved in methanol (2 mL) and then the solvent was evaporated under vacuum. Distilled water and TRIS buffer were then added to the obtained DMS/PEGS mixture as described above and the solutions were injected into rod-shaped molds. Microstructure of PEG/Silica Hydrogels Cross-sectional images of the hydrogels were obtained using a Hitachi S-3400 scanning electron microscope (Tokyo, Japan) at an acceleration voltage of 15 keV and probe current of 25 mA. Hydrogel samples were dehydrated by freeze-drying, plunged into liquid nitrogen for 5 min and broken into small pieces. Prior to scanning electron microscopy (SEM) analysis, each sample was gold sputter coated (SC7640 Sputter).
MATERIALS AND METHODS
Fourier Transform Infrared Spectroscopy Analysis
Materials
Fourier transform infrared (FTIR) spectra of freeze-dried 4arm PEG, PEGS, PEG/silica hydrogels, and drug-loaded hydrogels were obtained (Spectrum One FTIR; Perkin-Elmer, Woodbridge, ON, Canada). The spectra were recorded between 4000 and 650 cm−1 and analyzed using Spectrum V5.0.1 software (PerkinElmer). All spectra were an average of 20 scans at a resolution of 2 cm−1 and repeated in triplicate.
Tetraethoxysilane (≥99.9%) and Trizma base, (TRIS, ≥99.9%) were purchased from Sigma–Aldrich (Oakville, ON, Canada). 4-arm PEG-OH (MW = 2000 Da) was purchased from Creative PEGWorks (Winston-Salem, NC). DMS (98%) was purchased from Alfa Aesar (Georgetwon, ON, Canada). DMSP was purchased from Medisca (Richmond, BC, Canada). All other reagents were of analytical grade. R
1
H NMR Characterization
1
H NMR spectra were recorded in Dimethyl sulfoxided6 (DMSO-d6 ) (for TEOS and poly(ethylene glycol)-silicate, PEGS] or D2 O (for monitoring PEG/silica hydrogel formation) using a Varian Gemini 400 MHz spectrometer and referenced to residual protium in the solvent. Synthesis of 4-Arm PEGS Tetraethoxysilane (0.89 mL, 4.0 mmol) was dissolved in methanol (2 mL) and HCl (0.6 N) was added (0.1 mL) to reduce the pH of the solution to 1.5. Following the addition of 4-arm PEG-OH (2 g, 1.0 mmol) to the above mixture, the solution was stirred at 25◦ C for 0.5 h and then heated up to 60◦ C for an additional 40 h prior to evaporation of the methanol and ethanol generated from the reaction. The PEGS obtained was a colorless and viscous liquid at room temperature. 1 H NMR (400 MHz, DMSO-d6 , * in ppm): 1.16 (br, SiOCH2 CH3 ), 3.48 (m, O–CH2 CH2 –O), 3.79 (br, SiOCH2 CH3 ). Preparation of PEG/Silica Hydrogels The PEG/silica hydrogels were prepared by mixing PEGS (0.1 g) with distilled water in mass ratios of 1:0, 1:1, and 1:2 as summarized in Table 1. Solutions were vortexed for several minutes to ensure complete dissolution. TRIS buffer (0.15 mL, 0.05 mM, pH 8.6) was then added to each solution and the mixtures were vortexed and then immediately injected into rod-shaped molds (20–25 mm in length and 2 mm in diameter). In situ gelation of the hydrogels occurred at room temperature within several minutes. DMSP-loaded hydrogels were DOI 10.1002/jps.23777
Rheological Studies The rheological properties of the PEG/silica hydrogels with varying water content were characterized using a stresscontrolled rheometer with a 4-cm cone and a 2◦ angle plate geometry attachment at 37◦ C (AR-2000; TA Instruments). The rheometer was calibrated and rotational mapping was performed according to instrument specifications. Time dependent changes in viscosity were measured using an oscillatory time sweep test, where a frequency of 1.0 rad/s was applied. Hydrogel samples with varying water content were prepared and vortexed with TRIS buffer (pH 8.6). Then 100 :L of each sample was injected onto the rheometer plate for analysis. The strain was set to 10% and the resulting shear stress was monitored over time.
In Vitro Degradation The rod-shaped PEG/silica hydrogel samples (50mg) were incubated in 6 mL of phosphate-buffered saline (PBS, pH 7.4) at 37◦ C. At specific time intervals, a 6-mL aliquot was removed and replaced with fresh buffer. The degradation byproduct, Si(OH)4 , of the hydrogels was measured using an established colorimetric assay as described elsewhere.52 Briefly, ammonium molybdate (10%) was added to each aliquot to form silic-monobdenic acid. Then, concentrated hydrogen chloride (36%) was added. Following this, the reducing agent and oxalic acid (10%) were added to each sample to obtain a silicmolybdenic blue complex. Subsequently, the UV absorbance of samples at 815 nm was measured using a Varian Cary 50 Series spectrometer (Agilent Technologies Canada Inc., Mississauga, ON, Canada) . The percent degradation of the PEG/silica Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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hydrogel was calculated using the following equation:
to measure statistical significance between pairs of results. A p value of less than 0.05 was considered to be significant.
Percent Degradation(%) =
Total weight of Si(OH)4 from PEG/silica hydrogel × 100% Initial weight of TEOS in PEG/silica hydrogel
In Vitro Release of DMS and DMSP from PEG/Silica Hydrogels Hydrogel samples loaded with either DMS or DMSP were placed in dialysis bags, suspended in glass vials containing 15 mL of PBS (pH 7.4) and incubated at 37◦ C with constant shaking. At specific time intervals, 14 mL of the PBS was removed and replaced with fresh buffer. The concentration of drug in PBS was measured by HPLC analysis. In brief, an Agilent Series 1200 HPLC (Agilent Technologies, Canada) equipped with a Waters 4.6 ×250 mm2 column (XTerra MS C18; 5 :m particle size), Waters 2487 dual absorbance detector (Waters), and ChemStation software (Agilent Technologies) was used for analysis. Both DMS and DMSP were detected at 254 nm. For analysis of the DMS samples, the mobile phase consisted of 50% acetonitrile and 50% water with a flow rate of 1 mL/min (retention time = 4.2 min). For DMSP samples, the mobile phase consisted of 50% water and 50% methanol with a flow rate of 0.5 mL/min (retention time = 10 min). Cumulative amounts of DMS and DMSP released from the hydrogels were determined using calibration curves with linear ranges between 0.206 and 206 :g/mL for DMS and 0.150 and 200 :g/mL for DMSP. R
Statistical Analysis All results were obtained from data groups of n ≥ 3 and are expressed as mean ± standard error. A two-sample t-test was used
RESULTS AND DISCUSSION Synthesis of the PEG/Silica Hydrogels The PEG/silica hydrogels were prepared as illustrated in Figure 1a, following a method previously reported by Krupa et al.52 with slight modifications. PEGS was prepared from TEOS and 4-arm PEG (2 kDa) under acidic conditions in methanol. It has been reported that the direct replacement of ethoxy groups on TEOS is possible when reacted with lower molecular weight polyols.39 In the presence of acid, one of the ethoxy groups on TEOS is replaced by the hydroxyl group of the 4arm PEG to generate PEGS via transesterification and the release of ethanol as the byproduct. Upon solvent removal, PEGS was obtained as a transparent solution with enhanced solubility in water compared with TEOS. Successful preparation of PEGS was confirmed by 1 H NMR, which showed new peaks at 1.16 and 3.79 ppm (Fig. 2b) corresponding to the triethoxysilyl groups. These two peaks appear broader in PEGS than in TEOS (Fig. 2a) mainly because of the restricted mobility of the ethoxy groups in the PEG/silica matrix. To form PEG/silica hydrogels via the sol-gel process, TRIS buffer (pH 8.6) and set amounts of double distilled water were mixed with PEGS and then injected into rod-shaped molds. The composition of each PEG/silica hydrogel prepared is listed in Table 1. In the presence of water, PEGS undergoes hydrolysis to generate free silanols, which condense to form the PEGS/silica hydrogel. The hydrolysis and polycondensation of PEGS were carried out under weakly basic conditions (pH 8.6)
Figure 1. An illustration of the formation of the PEG/silica hydrogel and drug release. (a) Synthetic scheme for the PEG/silica hydrogel. (b) Illustration of the proposed mechanisms of release for hydrophilic and hydrophobic drugs from the PEG/silica hydrogels. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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RESEARCH ARTICLE – Pharmaceutics, Drug Delivery and Pharmaceutical Technology
Figure 2. groups.
1H
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NMR spectra of (a) TEOS, (b) PEGS showing triethoxysilyl groups, and (c) PEG/silica hydrogel after hydrolysis of triethoxysilyl
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Figure 3. Fourier transform infrared spectra of 4-arm PEG, PEGS, and PEG/silica hydrogel.
as it had been shown that basic conditions increase the rate of polycondensation and result in a more rapid gelation of silica hydrogels.53 To monitor the formation of PEG/silica hydrogel, the 1 H NMR spectrum of PEGS was recorded in D2 O (Fig. 2c) and compared with the spectrum obtained in deuterated DMSO (Fig. 2b). When D2 O is employed as the solvent, we observe
the disappearance of peaks corresponding to the triethoxysilyl groups (Fig. 2c) of PEGS, which indicates the hydrolysis of PEGS to form free silanols and the release of ethanol as a byproduct. Condensation of the free silanols occurs concurrently with the hydrolysis of the triethoxysilyl groups resulting in the formation of the PEG/silica hydrogel. Upon complete gelation, the rod-shaped PEG/silica hydrogels of neutral pH and varying water content (R1.5 , R2.5 , and R3.5 ) were obtained. Fourier transform infrared analysis of the PEG/silica hydrogel and its individual components are shown in Figure 3. The peak at 2870 cm−1 corresponds to the C–H stretching vibrations from PEG. The peaks at 1091 and 844 cm−1 are asymmetric and symmetric C–O–C (PEG) and Si–O–Si (PEGS and PEG/silica hydrogel) stretching vibrations, respectively. As well, the small peak at 1249 cm−1 can be attributed to asymmetric stretching of the same functional groups. The peaks at 807 and 946 cm−1 correspond to Si–OH vibrations. The broad peak from 3100 to 3700 cm−1 is because of the stretching vibrations of hydrogen bonds formed between Si–OH and the absorbed water. These peaks further confirm the formation of the hydrogels when PEGS is brought into contact with aqueous medium. Microstructure Scanning electron microscopy was used to study the fractural surface of the PEG/silica hydrogels prepared with varying water content. The R1.5 hydrogel sample showed a smooth microstructure (Figs. 4a and 4b). For the R2.5 hydrogel, the
Figure 4. Scanning electron microscopy images of PEG/silica hydrogels. (a) R1.5 , (c) R2.5 , and (e) R3.5 samples at low magnification and (b) R1.5 , (d) R2.5 , and (f) R3.5 samples at high magnification. Scale bars are 50 and 10 :m for the low- and high-magnification images, respectively. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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increase in water content in the PEGS solution resulted in a fractural surface that was rougher in appearance (Figs. 4c and 4d). When the mass ratio of water to PEGS (i.e., R value) increased to 3.5, the SEM image revealed a porous hydrogel containing interconnected macropores with a diameter of around 10 :m (Figs. 4e and 4f). Although the discontinuity in morphology was found to become more prevalent with an increase in the amount of water in the hydrogel, any distinct signs of domain structure arising from the presence of largely 4-arm PEG or silicate-enriched phases were not observed.
Rheological Properties To understand the gelation kinetics associated with the sol-gel process for the formation of PEG/silica hydrogels, time sweep measurements were carried out on all hydrogel samples listed in Table 1 at 37◦ C. Specifically, the storage modulus (G ) and loss modulus (G ) were measured as a function of time. Figure 5 shows G and G as a function of time for the PEG/silica hydrogel samples with different water content (i.e., R = 1.5, 2.5, or 3.5). During the sol-gel process, the alkoxide groups of PEGS are hydrolyzed quickly under alkaline conditions and are replaced by hydroxyl groups. As gelation proceeds, the hydrogel matrix forms through condensation of the silanol groups on each PEGS causing G to increase above G . The crossover point between G and G is said to indicate the gelation time (tgel ).54 The tgel value was found to be 0.58 min for the hydrogel sample with the lowest water content (R1.5 ). A 10-fold increase in the tgel value was observed for the hydrogel with the highest water content (5.2 min for R3.5 ). This is attributed to the associated decrease in the concentration of the PEGS solution that occurs with an increase in water content that reduces the rate of condensation. In addition, a “bump” in G was observed (Fig. 5a) for the PEG/silica hydrogel sample with the lowest water content (R1.5 ). This may be attributed to the “syneresis property” of the PEG/silica hydrogel that occurs because of reorganization of bonds within the hydrogel network.55 Similar phenomena have been reported for other hydrogel systems.56–58 A plot of the change in viscosity versus time further demonstrates that the dilution that occurs because of the increase in water content in the PEG/silica hydrogel samples reduces the rate of gelation. As shown in Figure 6, the viscosity increases more slowly in the samples with higher water content.
In Vitro Degradation The degradation of PEG/silica hydrogels in PBS (pH 7.4) was monitored using a molybdenum-based colorimetric assay that detects the formation of the degradation product, Si(OH)4 , as a silic-molybdenic blue complex. As shown in Figure 7, the PEG/silica hydrogel with the lowest water content (R1.5 ) degraded at a faster rate than the hydrogel with the highest water content (R3.5 ). For all samples, degradation occurred at a linear rate for the first 6 days. Following the 13-day incubation period, total degradation was observed for PEG/silica hydrogel samples with R1.5 , whereas the degree of degradation was 95% for R2.5 samples, and 73% for R3.5 samples. SEM images were also obtained for the R2.5 PEG/silica hydrogels to qualitatively examine the degree of degradation (Fig. 8). Following a 3-day period, SEM revealed many cracks on the surface of DOI 10.1002/jps.23777
Figure 5. The gelation kinetics of PEG/silica hydrogels with varying water content. (a) R1.5 , (b) R2.5 , and (c) R3.5 . Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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In Vitro Release of DMS and DMSP
Figure 6. Effect of water content on the viscosity of the PEG/silica hydrogels.
Figure 7. The degradation of the PEG/silica hydrogels formed with varying water content.
the hydrogel (Fig. 8a) when compared with the nondegraded sample (Fig. 4c). As shown in Figure 8b, 9 days after incubation in PBS, degradation of the PEG/silica hydrogel leads to the disappearance of a significant portion of the hydrogel matrix. These results clearly demonstrate that the water content has a significant influence on the stability and degradation properties of the PEG/silica hydrogel. A similar trend was observed by Viitala et al.59 wherein the rate of degradation of silica gels was found to decrease with an increase in water content.
Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
The in vitro release of DMS and DMSP from various PEG/silica hydrogel samples (R1.5 and R2.5 ) was assessed in PBS to determine the mechanism and kinetics of hydrophobic and hydrophilic drug release (Fig. 1b). The total loading level of DMS and DMSP in each hydrogel sample was 0.85 and 5.85 mg, respectively, and the weight ratio of drug to hydrogel was 1:42 for the DMS/hydrogel samples and 1:9 for the DMSP/hydrogel samples. As shown in Figure 9a, for the lower water content sample (R1.5 ), over 60% of DMSP was released within 10 min and nearly 100% of the drug was released in 1 h. On the contrary, for the higher water content sample (R2.5 ), 84% of DMSP was released within 1 h, with 61% of the drug being released in the first 20 min. In comparison, only 4.7% and 2.7% of the hydrophobic drug, DMS, was released in the first hour from the PEG/silica hydrogels with R1.5 and R2.5 , respectively (Fig. 10). For the R1.5 PEG/silica hydrogel, DMS was released at a constant rate of 34% per day for the first 2 days of incubation. The remaining DMS was released over the following 4 days. However, for the higher water content PEG/silica hydrogel (R2.5 ), DMS was released at a constant rate of 13.8% per day for the first 6 days. From day 6 to 10 the release leveled off with a total of 12% still remaining in the hydrogel after 10 days of incubation. Silica-based hydrogels obtained under weakly alkaline conditions have been shown to be porous, which has proven to be useful for drug loading and release.60 During the hydrogel formation, DMSP likely associates with PEG and is dissolved in the solution that is contained in the micropores within the hydrogel matrix (Fig. 1b). When buffer enters the hydrogel, the large interconnecting pores in the PEG/silica hydrogel allow for rapid drug release.35 To determine whether the observed differences in the release of the hydrophilic drug DMSP and the hydrophobic drug DMS was a result of the sample preparation method, DMSP was also loaded into the PEG/silica hydrogel using the solvent cast method that was employed for loading DMS. As shown in Figure 9b, a change in the method employed for drug loading did not result in a significant difference in the in vitro drug release profile. This confirms that the differential release of the two drugs is because of the distinct nature and physicochemical properties of each drug rather than the method employed for drug loading. Further analysis of the release kinetics demonstrates a linear dependence of −ln(1−Mt /M0 ) on time for DMSP released from the PEG/silica hydrogel under sink conditions (Fig. 11). Mt is the cumulative amount of DMSP released at time t, and M0 is the initial amount of DMSP in the PEG/silica hydrogel samples. The linear relationship observed for −ln(1−Mt /M0 ) versus time for DMSP release from the hydrogels with r2 values of 0.98 (R1.5 ) and 0.95 (R2.5 ) indicates that the release profile follows first-order kinetics. As shown in Figure 10, DMS is released in a sustained manner over 2 or 6 days for the R1.5 or R2.5 hydrogels, respectively. In comparing Figure 7 and Figure 10, it is clear that the release profiles can be related to the degradation of the hydrogel (Fig. 1b). In the early stages of the sol-gel process, especially at the higher pH value, discrete colloidal silica capsule particles are formed55 and as the polycondensation reaction proceeds, the small sized DMS-loaded PEG/silica hydrogel particles grow
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Figure 8. Scanning electron microscopy images of the degradation of PEG/silica hydrogels with R2.5 in PBS buffer at 37◦ C following (a) 3 and (b) 9 days incubation periods.
and aggregate, and the heterogeneous structure of the silica hydrogel matrix is formed. A constant rate of release of DMS was observed over 2 and 6 days from PEG/silica R1.5 (r2 = 0.98) and PEG/silica R2.5 (r2 = 0.99) hydrogels, respectively. The burst release within the first 10 min can be attributed to dissolution of the fraction of DMS that was not well entrapped within the hydrogel matrix. The zero-order release profile observed may be largely attributed to the linear degradation of the PEG/silica hydrogel matrix as shown in Figure 7. Therefore, overall this hydrogel system is capable of achieving sustained release of hydrophobic drugs such as DMS, the rate of which can be tailored by varying the water content used during preparation.
CONCLUSIONS
Figure 9. In vitro drug release profiles of (a) DMSP from the PEG/silica hydrogels (R = 1.5 or 2.5) and (b) in vitro release profiles of DMSP from the PEG/silica hydrogels (R = 1.5) prepared by the in situ and solvent cast methods.
DOI 10.1002/jps.23777
To develop and implement effective ocular depot drug delivery systems, a number of criteria must be met, including biocompatibility, biodegradability, optimal drug release profile, and injectability. In the present study, PEG/silica hydrogels were prepared via a sol-gel process and investigated as delivery systems for hydrophilic (i.e., DMSP) and hydrophobic (i.e., DMS) ophthalmic drugs. These hydrogels fulfill all of the aforementioned criteria for an effective ocular depot delivery system: (1) the materials used to prepare the hydrogels are biocompatible; (2) the depot systems are biodegradable; (3) a sustained drug release profile is achieved with minimal initial burst; and (4) the hydrogels are injectable. The results of this study demonstrate that the properties of the PEG/silica hydrogels such as tgel and rate of degradation depend heavily on initial water content. Evaluation of in vitro release revealed different release profiles for the hydrophilic and hydrophobic drugs, with DMSP being released more quickly than DMS. Release of DMSP was shown to follow first-order kinetics, which was postulated to result from diffusion of DMSP through micron-sized pores within the PEG/silica hydrogel. The release of DMS followed zero-order kinetics, depending on the water content in the hydrogel. The zero-order release of this drug was attributed to the linear degradation rate of the hydrogel. In general, this sol-gel derived rapid in situ forming, and degradable PEG/silica hydrogel has been shown to be promising for the sustained release of hydrophobic drugs. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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Figure 10. In vitro drug release profiles of DMS from the PEG/silica hydrogels (R = 1.5 or 2.5).
Figure 11. The plots of −ln(1-Mt /M0 ) versus time for DMSP release from PEG/silica hydrogels (R = 1.5 or 2.5). Mt is the cumulative amount of DMSP released at time t and M0 is the initial amount of DMSP in the PEG/silica hydrogel samples.
ACKNOWLEDGMENTS This project is supported by an NSERC Operating grant to C. Allen and the 20/20 NSERC Ophthalmic Materials Network.
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nal vein occlusion thirty-six-month results. Ophthalmology 119(1):132– 137. 3. Lu CH, Yoganathan RB, Kociolek M, Allen C. 2013. Hydrogel containing silica shell cross-linked micelles for ocular drug delivery. J Pharm Sci 102(2):627–637. 4. Alvarez-Lorenzo C, Hiratani H, Gomez-Amoza JL, Martinez-Pacheco R, Souto C, Concheiro A. 2002. Soft contact lenses capable of sustained delivery of timolol. J Pharm Sci 91(10):2182–2192. 5. Kapoor Y, Thomas JC, Tan G, John VT, Chauhan A. 2009. Surfactant-laden soft contact lenses for extended delivery of ophthalmic drugs. Biomaterials 30(5):867–878. DOI 10.1002/jps.23777
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6. Tieppo A, White CJ, Paine AC, Voyles ML, McBride MK, Byrne ME. 2012. Sustained in vivo release from imprinted therapeutic contact lenses. Journal Control Release 157(3):391–397. 7. Gulsen D, Chauhan A. 2004. Ophthalmic drug delivery through contact lenses. Invest Ophthalmol Vis Sci 45(7):2342–2347. 8. Kapoor Y, Chauhan A. 2008. Drug and surfactant transport in cyclosporine A and Brij 98 laden p-HEMA hydrogels. J Colloid Interface Sci 322(2):624–633. 9. Schultz CL, Morck DW. 2010. Contact lenses as a drug delivery device for epidermal growth factor in the treatment of ocular wounds. Clin Exp Optom 93(2):61–65. 10. Wong V, Wood L. 2006. Conveniently implantable sustained release drug compositions. US Patent WO2006/039336 A2. 11. Ariturk N, Oge I, Erkan D, Sullu Y, Sahn M. 1996. The effects of nasolacrimal canal blockage on topical medications for glaucoma. Acta Ophthalmol Scan 74(4):411–413. 12. Peng CC, Kim J, Chauhan A. 2010. Extended delivery of hydrophilic drugs from silicone-hydrogel contact lenses containing vitamin E diffusion barriers. Biomaterials 31(14):4032–4047. 13. Kim H, Csaky KG, Gilger BC, Dunn JP, Lee SS, Tremblay M, de Monasterio F, Tansey G, Yuan P, Bungay PM, Lutz RJ, Robinson MR. 2005. Preclinical evaluation of a novel episcleral cyclosporine implant for ocular graft-versus-host disease. Invest Ophthalmol Vis Sci 46(2):655–662. 14. Jaffe GJ, McCallum RM, Branchaud B, Skalak C, Butuner Z, Ashton P. 2005. Long-term follow-up results of a pilot trial of a fluocinolone acetonide implant to treat posterior uveitis. Ophthalmology 112(7):1192–1198. 15. Kuno N, Fujii S. 2011. Recent advances in ocular drug delivery systems. Polymers 3(1):193–221. 16. Molokhia SA, Thomas SC, Garff KJ, Mandell KJ, Wirostko BM. 2013. Anterior eye segment drug delivery systems: Current treatments and future challenges. J Ocul Pharmacol Ther 29(2):92–105. 17. Kempe S, Mader K. 2012. In situ forming implants—An attractive formulation principle for parenteral depot formulations. J Control Release 161(2):668–679. 18. Einmahl S, Behar-Cohen F, D’Hermies F, Rudaz S, Tabatabay C, Renard G, Gurny R. 2001. A new poly(ortho ester)-based drug delivery system as an adjunct treatment in filtering surgery. Invest Ophthalmol Vis Sci 42(3):695–700. 19. Qiao MX, Chen DW, Ma XC, Liu YJ. 2005. Injectable biodegradable temperature-responsive PLGA-PEG-PLGA copolymers: Synthesis and effect of copolymer composition on the drug release from the copolymerbased hydrogels. Int J Pharm 294(1–2):103–112. 20. Jansen J, Koopmans SA, Los LI, van der Worp RJ, Podt JG, Hooymans JMM, Feijen J, Grijpma DW. 2011. Intraocular degradation behavior of crosslinked and linear poly(trimethylene carbonate) and poly(D,L-lactic acid). Biomaterials 32(22):4994–5002. 21. Natarajan JV, Chattopadhyay S, Ang M, Darwitan A, Foo S, Zhen M, Koo M, Wong TT, Venkatraman SS. 2011. Sustained release of an anti-glaucoma drug: Demonstration of efficacy of a liposomal formulation in the rabbit eye. PLoS One 6(9). 22. Packhaeuser CB, Schnieders J, Oster CG, Kissel T. 2004. In situ forming parenteral drug delivery systems: An overview. Eu J Pharm Biopharm 58(2):445–455. 23. Hatefi A, Amsden B. 2002. Biodegradable injectable in situ forming drug delivery systems. J Control Release 80(1–3):9–28. 24. Lee DS, Shim WS, Kim JH, Kim K, Kim YS, Park RW, Kim IS, Kwon IC. 2007. pH- and temperature-sensitive, injectable, biodegradable block copolymer hydrogels as carriers for paclitaxel. Int J Pharm 331(1):11–18. 25. Kurisawa M, Chung JE, Yang YY, Gao SJ, Uyama H. 2005. Injectable biodegradable hydrogels composed of hyaluronic acid-tyramine conjugates for drug delivery and tissue engineering. Chem Commun (34):4312–4314. 26. Lee DS, Minh KN. 2010. Injectable biodegradable hydrogels. Macromol Biosci 10(6):563–579.
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27. Bhattarai N, Gunn J, Zhang M. 2010. Chitosan-based hydrogels for controlled, localized drug delivery. Adv Drug Deliv Rev 62(1):83– 99. 28. Quintanar-Guerrero D, Ganem-Quintanar A, Nava-Arzaluz MG, Pinon-Segundo E. 2009. Silica xerogels as pharmaceutical drug carriers. Expert Opin Drug Deliv 6(5):485–498. 29. Radin S, El-Bassyouni G, Vresilovic EJ, Schepers E, Ducheyne P. 2005. In vivo tissue response to resorbable silica xerogels as controlledrelease materials. Biomaterials 26(9):1043–1052. 30. Teoli D, Parisi L, Realdon N, Guglielmi M, Rosato A, Morpurgo M. 2006. Wet sol-gel derived silica for controlled release of proteins. J Control Release 116(3):295–303. 31. Prokopowicz M. 2010. Synthesis and in vitro characterization of freeze-dried doxorubicin-loaded silica xerogels. J Sol-Gel Sci Technol 53(3):525–533. 32. Munusamy P, Seleem MN, Alqublan H, Tyler R, Sriranganathan N, Pickrell G. 2009. Targeted drug delivery using silica xerogel systems to treat diseases due to intracellular pathogens. Mat Sci Eng 29(8):2313– 2318. 33. Kortesuo P, Ahola M, Karlsson S, Kangasniemi I, Yli-Urpo A, Kiesvaara J. 2000. Silica xerogel as an implantable carrier for controlled drug delivery—Evaluation of drug distribution and tissue effects after implantation. Biomaterials 21(2):193–198. 34. Ahola M, Kortesuo P, Kangasniemi I, Kiesvaara J, Yli-Urpo A. 2000. Silica xerogel carrier material for controlled release of toremifene citrate. Int J Pharm 195(1–2):219–227. 35. Czarnobaj K, Prokopowicz M, Łukasiak J. 2006. The effect of PEG concentration on the release rate of cisplatin from PEG-modified silica xerogels. Drug Deliv 13(5):339–344. 36. Gill I, Ballesteros A. 1998. Encapsulation of biologicals within silicate, siloxane, and hybrid sol-gel polymers: An efficient and generic approach. J Am Chem Soc 120(34):8587–8598. 37. Haruvy Y, Ryabov Y, Arkhipov V, Gutina A, Axelrod E, Feldman Y. 2002. Fast-sol-gel derived silsequioxane glasses embodying glycerol moieties: Dielectric properties and morphology. J Non-Cryst Solids 305(1–3):226–234. 38. Nair PD, Nair S, Remya NS, Remya S. 2011. A biodegradable in situ injectable hydrogel based on chitosan and oxidized hyaluronic acid for tissue engineering applications. Carbohydr Polym 85(4):838– 844. 39. Brook MA, Chen Y, Guo K, Zhang Z, Brennan JD. 2004. Sugarmodified silanes: Precursors for silica monoliths. J Mater Chem 14(9):1469–1479. 40. Barbe CJ, Finnie KS, Waller DJ, Perret FL, Krause-Heuer AM, Lin HQ, Hanna JV. 2009. Biodegradability of sol-gel silica microparticles for drug delivery. J Sol-Gel Sci Technol 49(1):12–18. 41. Finnie KS, Jacques DA, McGann MJ, Blackford MG, Barbe CJ. 2006. Encapsulation and controlled release of biomolecules from silica microparticles. J Mater Chem 16(46):4494–4498. 42. Jeong J, Cho WS, Choi M, Han BS, Cho M, Oh J, Park K, Kim SJ, Kim SH. 2007. Inflammatory mediators induced by intratracheal instillation of ultrafine amorphous silica particles. Toxicol Lett 175(1– 3):24–33. 43. Popplewell JF, King SJ, Day JP, Ackrill P, Fifield LK, Cresswell RG, Di Tada ML, Liu K. 1998. Kinetics of uptake and elimination of silicic acid by a human subject: A novel application of Si-32 and accelerator mass spectrometry. J Inorg Biochem 69(3):177–180. 44. Viitala R, Jokinen M, Tuusa S, Rosenholm JB, Jalonen H. 2005. Adjustably bioresorbable sol-gel derived SiO2 matrices for release of large biologically active molecules. J Sol-Gel Sci Technol 36:147–156. 45. Viitala R, Jokinen M, Maunu SL, Jalonen H, Rosenholm JB. 2005. Chemical characterization of bioresorbable sol-gel derived SiO2 matrices prepared at protein-compatible pH. J Non-Cryst Solids 351:3225– 3234. 46. Otsuka H, Nagasaki Y, Kataoka K. 2003. PEGylated nanoparticles for biological and pharmaceutical applications. Adv Drug Deliv Rev 55(3):403–419.
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47. Liu SQ, Rachel Ee PL, Ke CY, Hedrick JL, Yang YY. 2009. Biodegradable poly(ethylene glycol)-peptide hydrogels with welldefined structure and properties for cell delivery. Biomaterials 30(8):1453–1461. 48. Shikanov A, Smith RM, Xu M, Woodruff TK, Shea LD. 2011. Hydrogel network design using multifunctional macromers to coordinate tissue maturation in ovarian follicle culture. Biomaterials 32(10):2524– 2531. 49. Van De Manakker F, Vermonden T, El Morabit N, Van Nostrum CF, Hennink WE. 2008. Rheological behavior of self-assembling PEG$-cyclodextrin/PEG-cholesterol hydrogels. Langmuir 24(21):12559– 12567. 50. Wieland JA, Houchin-Ray TL, Shea LD. 2007. Non-viral vector delivery from PEG-hyaluronic acid hydrogels. J Control Release 120(3):233–241. 51. Yang J, Jacobsen MT, Pan H, Kopeˇcek J. 2010. Synthesis and characterization of enzymatically degradable PEG-based peptidecontaining hydrogels. Macromol Biosci 10(4):445–454. 52. Krupa I, Nedelcev T, Racko D, Lacik I. 2010. Mechanical properties of silica hydrogels prepared and aged at physiological conditions: Testing in the compression mode. J Sol-Gel Sci Technol 53(1):107– 114.
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53. Tilgner IC, Fischer P, Bohnen FM, Rehage H, Maier WF. 1995. Effect of acidic, basic and fluoride-catalyzed sol-gel transitions on the preparation of sub-nanostructured silica. Microporous Mater 5(1– 2):77–90. 54. Winter HH, Chambon F. 1986. Analysis of linear viscoelasticity of a cross-linking polymer at the gel point. J Rheol 30(2):367–382. 55. Hench LL, West JK. 1990. The sol-gel process. Chem Rev 90(1):33– 72. 56. Moura MJ, Figueiredo MM, Gil MH. 2007. Rheological study of genipin cross-linked chitosan hydrogels. Biomacromolecules 8(12):3823–3829. 57. Johnson SB, Dunstan DE, Franks GV. 2004. A novel thermallyactivated crosslinking agent for chitosan in aqueous solution: A rheological investigation. Colloid Polym Sci 282(6):602–612. 58. Zhang LM, Ma D, Tu K. 2010. Bioactive supramolecular hydrogel with controlled dual drug release characteristics. Biomacromolecules 11(9):2204–2212. 59. Viitala R, Jokinen M, Rosenholm JB. 2007. Mechanistic studies on release of large and small molecules from biodegradable SiO2 . Int J Pharm 336(2):382–390. 60. Czarnobaj K. 2008. Preparation and characterization of silica xerogels as carriers for drugs. Drug Deliv 15(8):485–492.
DOI 10.1002/jps.23777
Multi-arm PEG/Silica Hydrogel for Sustained Ocular Drug Delivery CHANGHAI LU, PAYAM ZAHEDI, ADAM FORMAN, CHRISTINE ALLEN Leslie Dan Faculty of Pharmacy, University of Toronto, Toronto, Ontario, Canada Received 2 June 2013; revised 20 September 2013; accepted 18 October 2013 Published online 27 November 2013 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jps.23777 ABSTRACT: In the present study, a series of sustained drug delivery multiarm poly(ethylene glycol) (PEG)/silica hydrogels were prepared and characterized. The hydrogels were formed by hydrolysis and condensation of poly(4-arm PEG silicate) using the sol-gel method. The relationships between water content in the PEG/silica hydrogel and stability as well as rheological properties were evaluated. Scanning electron microscopy analysis of the PEG/silica hydrogels revealed water content-dependent changes in microstructure. An increase in water content resulted in larger pores within the hydrogel, longer gelation time and higher viscosity. The PEG/silica hydrogels were loaded with dexamethasone (DMS) or dexamethasone sodium phosphate (DMSP), drugs that are hydrophobic and hydrophilic in nature, respectively. Evaluation of in vitro release revealed a zero-order release profile for DMS over the first 6 days, suggesting that degradation of the silica hydrogel matrix was the primary mechanism of drug release. It was also found that the drug-release profile could be tailored by varying the water content used during hydrogel preparation. In contrast, more than 90% of DMSP was released within 1 h, suggesting that DMSP release was only controlled by diffusion. Overall, results from this study indicate that PEG/silica hydrogels may be promising drug-eluting C 2013 Wiley Periodicals, Inc. and the American Pharmacists depot materials for the sustained delivery of hydrophobic, ophthalmic drugs. Association J Pharm Sci 103:216–226, 2014 Keywords: drug delivery systems; injectables; hydrogels; ophthalmic drug delivery; silica
INTRODUCTION Drug delivery systems that provide sustained release have the potential to revolutionize treatment in a number of indications, including ophthalmology.1–6 More than 90% of ocular drug formulations are in the form of eye drops, which are applied in volumes ranging from 50 to 100 :L to the precorneal area, with more than 75% of that volume being lost within the first 2–6 min.7 Indeed, following administration, it has been found that only 1%–5% of the drug reaches the desired site of action, making conventional eye drop-based formulations highly inefficient.8 In contrast, it has been demonstrated that significant drug levels can be achieved in tissues of the posterior segment of the eye using sustained drug delivery vehicles.9 In comparison to eye drop-based formulations, drug-eluting sustained delivery systems such as particulate systems10 , punctal plugs,11 contact lenses,3,12 and implants13,14 composed of biodegradable or nonbiodegradable polymers, injected and or implanted into the subconjunctival tissue and vitreous body, have the ability to enhance drug bioavailability and provide longer drug residence times.15 Therefore, these systems may be a favorable alternative to the conventional formulations commonly used to treat both anterior and posterior ailments of the eye. Despite significant research into the development of new depot technologies, only a handful of sustained-release systems have been approved for clinical use.15–17 Successful design of these systems is challenging given the number of criteria that must be met. Specifically, sustained-release depot systems must be biocompatible and biodegradable, ideally with degradation of the system occurring over the timescale of drug release.18–20 The system must also carry sufficient drug load
Correspondence to: Christine Allen (Telephone: +416-946-8594; Fax: +416978-8511; E-mail: [email protected]) Journal of Pharmaceutical Sciences, Vol. 103, 216–226 (2014)
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to deliver a therapeutic dose, and minimal initial burst release is often favored.21,22 In addition, for the majority of applications, injectable depot systems are less invasive and preferred over implantable systems that require invasive surgery for implantation.23 To date, a wide range of materials have been pursued for the preparation of depot delivery systems, with the most common being polyesters and polysaccharides.24–27 Silica xerogels have also been pursued for depot delivery applications owing to their biocompatibility, biodegradability, and their ability to provide sustained drug release.28–33 These hydrogels are prepared by a sol-gel process using alkoxysilicates, such as tetraethoxysilane (TEOS) and tetramethoxysilane. This process involves hydrolysis of silicate and the subsequent condensation of the silanol, either with itself or other polyols. For example, a number of groups have synthesized silica hydrogel networks using TEOS and linear poly(ethylene glycol) (PEG).34,35 Sugars, such as maltose and dextran, as well as glycerol, have also been combined with TEOS to produce hydrogel materials.36–39 Silica hydrogels are known to degrade in vivo to form monosilicic acid [Si(OH)4 ], which can be absorbed by the body and eliminated via the kidneys.40–43 The degradation of silica hydrogels under physiological conditions can be controlled by altering the waterto-alkoxide ratio, amount of solvent used during preparation, catalyst concentrations, and time of aging and drying.44,45 Thus, silica-based biodegradable matrices are ideally suited for use as depot drug delivery systems. Linear and multiarm PEG-based hydrogels have also been investigated as matrices for sustained drug delivery because of PEG’s versatility of functionalization and excellent biocompatibility.46–51 To the best of our knowledge, a combined PEG/silica hydrogel has yet to be evaluated as a depot drug delivery system for ophthalmic applications. In the current study, injectable PEG/silica hydrogels were prepared and their physicochemical and rheological properties were characterized. The ophthalmic drugs dexamethasone (DMS) and
RESEARCH ARTICLE – Pharmaceutics, Drug Delivery and Pharmaceutical Technology
Table 1. Material Components and Ratios Used in the Synthesis of the PEG/Silica Hydrogels Sample PEGS (g) H2 O (mL) TRISa (mL) Water Content (%)b
Rc
R1.5 R2.5 R3.5
1.5 2.5 3.5
0.1 0.1 0.1
0 0.1 0.2
0.15 0.15 0.15
60 71.4 77.7
a
pH value is 8.6. Water contents are calculated values based on the mass ratio of water and TRIS buffer to total the mass of water, TRIS, and PEGS in each hydrogel. c R values are generated from the mass ratio of water and TRIS buffer to the mass of PEGS. b
dexamethasone sodium phosphate (DMSP) were incorporated into the PEG/silica hydrogels, their in vitro release examined, and mechanisms of drug release elucidated. Overall, results from this study indicate that PEG/silica hydrogels may be promising drug-eluting depot materials for the sustained delivery of hydrophobic ocular drugs.
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prepared as described above using a solution containing DMSP (80 mg/mL) instead of distilled water. For preparation of DMS-loaded hydrogels, because of the limited aqueous solubility of this drug, DMS (4 mg) and PEGS (0.1 g) were first dissolved in methanol (2 mL) and then the solvent was evaporated under vacuum. Distilled water and TRIS buffer were then added to the obtained DMS/PEGS mixture as described above and the solutions were injected into rod-shaped molds. Microstructure of PEG/Silica Hydrogels Cross-sectional images of the hydrogels were obtained using a Hitachi S-3400 scanning electron microscope (Tokyo, Japan) at an acceleration voltage of 15 keV and probe current of 25 mA. Hydrogel samples were dehydrated by freeze-drying, plunged into liquid nitrogen for 5 min and broken into small pieces. Prior to scanning electron microscopy (SEM) analysis, each sample was gold sputter coated (SC7640 Sputter).
MATERIALS AND METHODS
Fourier Transform Infrared Spectroscopy Analysis
Materials
Fourier transform infrared (FTIR) spectra of freeze-dried 4arm PEG, PEGS, PEG/silica hydrogels, and drug-loaded hydrogels were obtained (Spectrum One FTIR; Perkin-Elmer, Woodbridge, ON, Canada). The spectra were recorded between 4000 and 650 cm−1 and analyzed using Spectrum V5.0.1 software (PerkinElmer). All spectra were an average of 20 scans at a resolution of 2 cm−1 and repeated in triplicate.
Tetraethoxysilane (≥99.9%) and Trizma base, (TRIS, ≥99.9%) were purchased from Sigma–Aldrich (Oakville, ON, Canada). 4-arm PEG-OH (MW = 2000 Da) was purchased from Creative PEGWorks (Winston-Salem, NC). DMS (98%) was purchased from Alfa Aesar (Georgetwon, ON, Canada). DMSP was purchased from Medisca (Richmond, BC, Canada). All other reagents were of analytical grade. R
1
H NMR Characterization
1
H NMR spectra were recorded in Dimethyl sulfoxided6 (DMSO-d6 ) (for TEOS and poly(ethylene glycol)-silicate, PEGS] or D2 O (for monitoring PEG/silica hydrogel formation) using a Varian Gemini 400 MHz spectrometer and referenced to residual protium in the solvent. Synthesis of 4-Arm PEGS Tetraethoxysilane (0.89 mL, 4.0 mmol) was dissolved in methanol (2 mL) and HCl (0.6 N) was added (0.1 mL) to reduce the pH of the solution to 1.5. Following the addition of 4-arm PEG-OH (2 g, 1.0 mmol) to the above mixture, the solution was stirred at 25◦ C for 0.5 h and then heated up to 60◦ C for an additional 40 h prior to evaporation of the methanol and ethanol generated from the reaction. The PEGS obtained was a colorless and viscous liquid at room temperature. 1 H NMR (400 MHz, DMSO-d6 , * in ppm): 1.16 (br, SiOCH2 CH3 ), 3.48 (m, O–CH2 CH2 –O), 3.79 (br, SiOCH2 CH3 ). Preparation of PEG/Silica Hydrogels The PEG/silica hydrogels were prepared by mixing PEGS (0.1 g) with distilled water in mass ratios of 1:0, 1:1, and 1:2 as summarized in Table 1. Solutions were vortexed for several minutes to ensure complete dissolution. TRIS buffer (0.15 mL, 0.05 mM, pH 8.6) was then added to each solution and the mixtures were vortexed and then immediately injected into rod-shaped molds (20–25 mm in length and 2 mm in diameter). In situ gelation of the hydrogels occurred at room temperature within several minutes. DMSP-loaded hydrogels were DOI 10.1002/jps.23777
Rheological Studies The rheological properties of the PEG/silica hydrogels with varying water content were characterized using a stresscontrolled rheometer with a 4-cm cone and a 2◦ angle plate geometry attachment at 37◦ C (AR-2000; TA Instruments). The rheometer was calibrated and rotational mapping was performed according to instrument specifications. Time dependent changes in viscosity were measured using an oscillatory time sweep test, where a frequency of 1.0 rad/s was applied. Hydrogel samples with varying water content were prepared and vortexed with TRIS buffer (pH 8.6). Then 100 :L of each sample was injected onto the rheometer plate for analysis. The strain was set to 10% and the resulting shear stress was monitored over time.
In Vitro Degradation The rod-shaped PEG/silica hydrogel samples (50mg) were incubated in 6 mL of phosphate-buffered saline (PBS, pH 7.4) at 37◦ C. At specific time intervals, a 6-mL aliquot was removed and replaced with fresh buffer. The degradation byproduct, Si(OH)4 , of the hydrogels was measured using an established colorimetric assay as described elsewhere.52 Briefly, ammonium molybdate (10%) was added to each aliquot to form silic-monobdenic acid. Then, concentrated hydrogen chloride (36%) was added. Following this, the reducing agent and oxalic acid (10%) were added to each sample to obtain a silicmolybdenic blue complex. Subsequently, the UV absorbance of samples at 815 nm was measured using a Varian Cary 50 Series spectrometer (Agilent Technologies Canada Inc., Mississauga, ON, Canada) . The percent degradation of the PEG/silica Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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hydrogel was calculated using the following equation:
to measure statistical significance between pairs of results. A p value of less than 0.05 was considered to be significant.
Percent Degradation(%) =
Total weight of Si(OH)4 from PEG/silica hydrogel × 100% Initial weight of TEOS in PEG/silica hydrogel
In Vitro Release of DMS and DMSP from PEG/Silica Hydrogels Hydrogel samples loaded with either DMS or DMSP were placed in dialysis bags, suspended in glass vials containing 15 mL of PBS (pH 7.4) and incubated at 37◦ C with constant shaking. At specific time intervals, 14 mL of the PBS was removed and replaced with fresh buffer. The concentration of drug in PBS was measured by HPLC analysis. In brief, an Agilent Series 1200 HPLC (Agilent Technologies, Canada) equipped with a Waters 4.6 ×250 mm2 column (XTerra MS C18; 5 :m particle size), Waters 2487 dual absorbance detector (Waters), and ChemStation software (Agilent Technologies) was used for analysis. Both DMS and DMSP were detected at 254 nm. For analysis of the DMS samples, the mobile phase consisted of 50% acetonitrile and 50% water with a flow rate of 1 mL/min (retention time = 4.2 min). For DMSP samples, the mobile phase consisted of 50% water and 50% methanol with a flow rate of 0.5 mL/min (retention time = 10 min). Cumulative amounts of DMS and DMSP released from the hydrogels were determined using calibration curves with linear ranges between 0.206 and 206 :g/mL for DMS and 0.150 and 200 :g/mL for DMSP. R
Statistical Analysis All results were obtained from data groups of n ≥ 3 and are expressed as mean ± standard error. A two-sample t-test was used
RESULTS AND DISCUSSION Synthesis of the PEG/Silica Hydrogels The PEG/silica hydrogels were prepared as illustrated in Figure 1a, following a method previously reported by Krupa et al.52 with slight modifications. PEGS was prepared from TEOS and 4-arm PEG (2 kDa) under acidic conditions in methanol. It has been reported that the direct replacement of ethoxy groups on TEOS is possible when reacted with lower molecular weight polyols.39 In the presence of acid, one of the ethoxy groups on TEOS is replaced by the hydroxyl group of the 4arm PEG to generate PEGS via transesterification and the release of ethanol as the byproduct. Upon solvent removal, PEGS was obtained as a transparent solution with enhanced solubility in water compared with TEOS. Successful preparation of PEGS was confirmed by 1 H NMR, which showed new peaks at 1.16 and 3.79 ppm (Fig. 2b) corresponding to the triethoxysilyl groups. These two peaks appear broader in PEGS than in TEOS (Fig. 2a) mainly because of the restricted mobility of the ethoxy groups in the PEG/silica matrix. To form PEG/silica hydrogels via the sol-gel process, TRIS buffer (pH 8.6) and set amounts of double distilled water were mixed with PEGS and then injected into rod-shaped molds. The composition of each PEG/silica hydrogel prepared is listed in Table 1. In the presence of water, PEGS undergoes hydrolysis to generate free silanols, which condense to form the PEGS/silica hydrogel. The hydrolysis and polycondensation of PEGS were carried out under weakly basic conditions (pH 8.6)
Figure 1. An illustration of the formation of the PEG/silica hydrogel and drug release. (a) Synthetic scheme for the PEG/silica hydrogel. (b) Illustration of the proposed mechanisms of release for hydrophilic and hydrophobic drugs from the PEG/silica hydrogels. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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Figure 2. groups.
1H
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NMR spectra of (a) TEOS, (b) PEGS showing triethoxysilyl groups, and (c) PEG/silica hydrogel after hydrolysis of triethoxysilyl
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Figure 3. Fourier transform infrared spectra of 4-arm PEG, PEGS, and PEG/silica hydrogel.
as it had been shown that basic conditions increase the rate of polycondensation and result in a more rapid gelation of silica hydrogels.53 To monitor the formation of PEG/silica hydrogel, the 1 H NMR spectrum of PEGS was recorded in D2 O (Fig. 2c) and compared with the spectrum obtained in deuterated DMSO (Fig. 2b). When D2 O is employed as the solvent, we observe
the disappearance of peaks corresponding to the triethoxysilyl groups (Fig. 2c) of PEGS, which indicates the hydrolysis of PEGS to form free silanols and the release of ethanol as a byproduct. Condensation of the free silanols occurs concurrently with the hydrolysis of the triethoxysilyl groups resulting in the formation of the PEG/silica hydrogel. Upon complete gelation, the rod-shaped PEG/silica hydrogels of neutral pH and varying water content (R1.5 , R2.5 , and R3.5 ) were obtained. Fourier transform infrared analysis of the PEG/silica hydrogel and its individual components are shown in Figure 3. The peak at 2870 cm−1 corresponds to the C–H stretching vibrations from PEG. The peaks at 1091 and 844 cm−1 are asymmetric and symmetric C–O–C (PEG) and Si–O–Si (PEGS and PEG/silica hydrogel) stretching vibrations, respectively. As well, the small peak at 1249 cm−1 can be attributed to asymmetric stretching of the same functional groups. The peaks at 807 and 946 cm−1 correspond to Si–OH vibrations. The broad peak from 3100 to 3700 cm−1 is because of the stretching vibrations of hydrogen bonds formed between Si–OH and the absorbed water. These peaks further confirm the formation of the hydrogels when PEGS is brought into contact with aqueous medium. Microstructure Scanning electron microscopy was used to study the fractural surface of the PEG/silica hydrogels prepared with varying water content. The R1.5 hydrogel sample showed a smooth microstructure (Figs. 4a and 4b). For the R2.5 hydrogel, the
Figure 4. Scanning electron microscopy images of PEG/silica hydrogels. (a) R1.5 , (c) R2.5 , and (e) R3.5 samples at low magnification and (b) R1.5 , (d) R2.5 , and (f) R3.5 samples at high magnification. Scale bars are 50 and 10 :m for the low- and high-magnification images, respectively. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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increase in water content in the PEGS solution resulted in a fractural surface that was rougher in appearance (Figs. 4c and 4d). When the mass ratio of water to PEGS (i.e., R value) increased to 3.5, the SEM image revealed a porous hydrogel containing interconnected macropores with a diameter of around 10 :m (Figs. 4e and 4f). Although the discontinuity in morphology was found to become more prevalent with an increase in the amount of water in the hydrogel, any distinct signs of domain structure arising from the presence of largely 4-arm PEG or silicate-enriched phases were not observed.
Rheological Properties To understand the gelation kinetics associated with the sol-gel process for the formation of PEG/silica hydrogels, time sweep measurements were carried out on all hydrogel samples listed in Table 1 at 37◦ C. Specifically, the storage modulus (G ) and loss modulus (G ) were measured as a function of time. Figure 5 shows G and G as a function of time for the PEG/silica hydrogel samples with different water content (i.e., R = 1.5, 2.5, or 3.5). During the sol-gel process, the alkoxide groups of PEGS are hydrolyzed quickly under alkaline conditions and are replaced by hydroxyl groups. As gelation proceeds, the hydrogel matrix forms through condensation of the silanol groups on each PEGS causing G to increase above G . The crossover point between G and G is said to indicate the gelation time (tgel ).54 The tgel value was found to be 0.58 min for the hydrogel sample with the lowest water content (R1.5 ). A 10-fold increase in the tgel value was observed for the hydrogel with the highest water content (5.2 min for R3.5 ). This is attributed to the associated decrease in the concentration of the PEGS solution that occurs with an increase in water content that reduces the rate of condensation. In addition, a “bump” in G was observed (Fig. 5a) for the PEG/silica hydrogel sample with the lowest water content (R1.5 ). This may be attributed to the “syneresis property” of the PEG/silica hydrogel that occurs because of reorganization of bonds within the hydrogel network.55 Similar phenomena have been reported for other hydrogel systems.56–58 A plot of the change in viscosity versus time further demonstrates that the dilution that occurs because of the increase in water content in the PEG/silica hydrogel samples reduces the rate of gelation. As shown in Figure 6, the viscosity increases more slowly in the samples with higher water content.
In Vitro Degradation The degradation of PEG/silica hydrogels in PBS (pH 7.4) was monitored using a molybdenum-based colorimetric assay that detects the formation of the degradation product, Si(OH)4 , as a silic-molybdenic blue complex. As shown in Figure 7, the PEG/silica hydrogel with the lowest water content (R1.5 ) degraded at a faster rate than the hydrogel with the highest water content (R3.5 ). For all samples, degradation occurred at a linear rate for the first 6 days. Following the 13-day incubation period, total degradation was observed for PEG/silica hydrogel samples with R1.5 , whereas the degree of degradation was 95% for R2.5 samples, and 73% for R3.5 samples. SEM images were also obtained for the R2.5 PEG/silica hydrogels to qualitatively examine the degree of degradation (Fig. 8). Following a 3-day period, SEM revealed many cracks on the surface of DOI 10.1002/jps.23777
Figure 5. The gelation kinetics of PEG/silica hydrogels with varying water content. (a) R1.5 , (b) R2.5 , and (c) R3.5 . Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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In Vitro Release of DMS and DMSP
Figure 6. Effect of water content on the viscosity of the PEG/silica hydrogels.
Figure 7. The degradation of the PEG/silica hydrogels formed with varying water content.
the hydrogel (Fig. 8a) when compared with the nondegraded sample (Fig. 4c). As shown in Figure 8b, 9 days after incubation in PBS, degradation of the PEG/silica hydrogel leads to the disappearance of a significant portion of the hydrogel matrix. These results clearly demonstrate that the water content has a significant influence on the stability and degradation properties of the PEG/silica hydrogel. A similar trend was observed by Viitala et al.59 wherein the rate of degradation of silica gels was found to decrease with an increase in water content.
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The in vitro release of DMS and DMSP from various PEG/silica hydrogel samples (R1.5 and R2.5 ) was assessed in PBS to determine the mechanism and kinetics of hydrophobic and hydrophilic drug release (Fig. 1b). The total loading level of DMS and DMSP in each hydrogel sample was 0.85 and 5.85 mg, respectively, and the weight ratio of drug to hydrogel was 1:42 for the DMS/hydrogel samples and 1:9 for the DMSP/hydrogel samples. As shown in Figure 9a, for the lower water content sample (R1.5 ), over 60% of DMSP was released within 10 min and nearly 100% of the drug was released in 1 h. On the contrary, for the higher water content sample (R2.5 ), 84% of DMSP was released within 1 h, with 61% of the drug being released in the first 20 min. In comparison, only 4.7% and 2.7% of the hydrophobic drug, DMS, was released in the first hour from the PEG/silica hydrogels with R1.5 and R2.5 , respectively (Fig. 10). For the R1.5 PEG/silica hydrogel, DMS was released at a constant rate of 34% per day for the first 2 days of incubation. The remaining DMS was released over the following 4 days. However, for the higher water content PEG/silica hydrogel (R2.5 ), DMS was released at a constant rate of 13.8% per day for the first 6 days. From day 6 to 10 the release leveled off with a total of 12% still remaining in the hydrogel after 10 days of incubation. Silica-based hydrogels obtained under weakly alkaline conditions have been shown to be porous, which has proven to be useful for drug loading and release.60 During the hydrogel formation, DMSP likely associates with PEG and is dissolved in the solution that is contained in the micropores within the hydrogel matrix (Fig. 1b). When buffer enters the hydrogel, the large interconnecting pores in the PEG/silica hydrogel allow for rapid drug release.35 To determine whether the observed differences in the release of the hydrophilic drug DMSP and the hydrophobic drug DMS was a result of the sample preparation method, DMSP was also loaded into the PEG/silica hydrogel using the solvent cast method that was employed for loading DMS. As shown in Figure 9b, a change in the method employed for drug loading did not result in a significant difference in the in vitro drug release profile. This confirms that the differential release of the two drugs is because of the distinct nature and physicochemical properties of each drug rather than the method employed for drug loading. Further analysis of the release kinetics demonstrates a linear dependence of −ln(1−Mt /M0 ) on time for DMSP released from the PEG/silica hydrogel under sink conditions (Fig. 11). Mt is the cumulative amount of DMSP released at time t, and M0 is the initial amount of DMSP in the PEG/silica hydrogel samples. The linear relationship observed for −ln(1−Mt /M0 ) versus time for DMSP release from the hydrogels with r2 values of 0.98 (R1.5 ) and 0.95 (R2.5 ) indicates that the release profile follows first-order kinetics. As shown in Figure 10, DMS is released in a sustained manner over 2 or 6 days for the R1.5 or R2.5 hydrogels, respectively. In comparing Figure 7 and Figure 10, it is clear that the release profiles can be related to the degradation of the hydrogel (Fig. 1b). In the early stages of the sol-gel process, especially at the higher pH value, discrete colloidal silica capsule particles are formed55 and as the polycondensation reaction proceeds, the small sized DMS-loaded PEG/silica hydrogel particles grow
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Figure 8. Scanning electron microscopy images of the degradation of PEG/silica hydrogels with R2.5 in PBS buffer at 37◦ C following (a) 3 and (b) 9 days incubation periods.
and aggregate, and the heterogeneous structure of the silica hydrogel matrix is formed. A constant rate of release of DMS was observed over 2 and 6 days from PEG/silica R1.5 (r2 = 0.98) and PEG/silica R2.5 (r2 = 0.99) hydrogels, respectively. The burst release within the first 10 min can be attributed to dissolution of the fraction of DMS that was not well entrapped within the hydrogel matrix. The zero-order release profile observed may be largely attributed to the linear degradation of the PEG/silica hydrogel matrix as shown in Figure 7. Therefore, overall this hydrogel system is capable of achieving sustained release of hydrophobic drugs such as DMS, the rate of which can be tailored by varying the water content used during preparation.
CONCLUSIONS
Figure 9. In vitro drug release profiles of (a) DMSP from the PEG/silica hydrogels (R = 1.5 or 2.5) and (b) in vitro release profiles of DMSP from the PEG/silica hydrogels (R = 1.5) prepared by the in situ and solvent cast methods.
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To develop and implement effective ocular depot drug delivery systems, a number of criteria must be met, including biocompatibility, biodegradability, optimal drug release profile, and injectability. In the present study, PEG/silica hydrogels were prepared via a sol-gel process and investigated as delivery systems for hydrophilic (i.e., DMSP) and hydrophobic (i.e., DMS) ophthalmic drugs. These hydrogels fulfill all of the aforementioned criteria for an effective ocular depot delivery system: (1) the materials used to prepare the hydrogels are biocompatible; (2) the depot systems are biodegradable; (3) a sustained drug release profile is achieved with minimal initial burst; and (4) the hydrogels are injectable. The results of this study demonstrate that the properties of the PEG/silica hydrogels such as tgel and rate of degradation depend heavily on initial water content. Evaluation of in vitro release revealed different release profiles for the hydrophilic and hydrophobic drugs, with DMSP being released more quickly than DMS. Release of DMSP was shown to follow first-order kinetics, which was postulated to result from diffusion of DMSP through micron-sized pores within the PEG/silica hydrogel. The release of DMS followed zero-order kinetics, depending on the water content in the hydrogel. The zero-order release of this drug was attributed to the linear degradation rate of the hydrogel. In general, this sol-gel derived rapid in situ forming, and degradable PEG/silica hydrogel has been shown to be promising for the sustained release of hydrophobic drugs. Lu et al., JOURNAL OF PHARMACEUTICAL SCIENCES 103:216–226, 2014
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Figure 10. In vitro drug release profiles of DMS from the PEG/silica hydrogels (R = 1.5 or 2.5).
Figure 11. The plots of −ln(1-Mt /M0 ) versus time for DMSP release from PEG/silica hydrogels (R = 1.5 or 2.5). Mt is the cumulative amount of DMSP released at time t and M0 is the initial amount of DMSP in the PEG/silica hydrogel samples.
ACKNOWLEDGMENTS This project is supported by an NSERC Operating grant to C. Allen and the 20/20 NSERC Ophthalmic Materials Network.
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