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Basal level insulin delivery: In vitro release, stability, biocompatibility, and in vivo absorption from thermosensitive triblock copolymers
Basal Level Insulin Delivery: In Vitro Release, Stability, Biocompatibility, and In Vivo Absorption from Thermosensitive Triblock Copolymers KHALED AL-TAHAMI, MAYURA OAK, RHISHIKESH MANDKE, JAGDISH SINGH Department of Pharmaceutical Sciences, College of Pharmacy, Nursing, and Allied Sciences, North Dakota State University, Fargo, North Dakota 58105 Received 10 January 2011; revised 6 April 2011; accepted 9 June 2011 Published online 28 June 2011 in Wiley Online Library (wileyonlinelibrary.com). DOI 10.1002/jps.22685 ABSTRACT: The major goal of this study was to develop the biodegradable and biocompatible thermosensitive polylactic acid–polyethylene glycol–polylactic acid triblock copolymer-based delivery systems for controlled release of basal level insulin for a longer duration after single subcutaneous injection. Insulin was dispersed into aqueous copolymer solutions to prepare the delivery system. The in vitro release profile of insulin from delivery systems was studied at 37◦ C in phosphate-buffered saline. Stability of released insulin was investigated using circular dichroism, differential scanning calorimetry, and matrix-assisted laser desorption/ionization–time-offlight mass spectrometry. A 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide assay and skin histology were used to determine the in vitro and in vivo biocompatibility of the delivery systems, respectively. Streptozotocin-induced diabetic rat model was used to study the in vivo absorption and bioactivity of insulin. In vitro release studies indicated that the delivery systems released insulin over 3 months in structurally stable form. The delivery systems were biocompatible in vitro and in vivo. In vivo absorption and bioactivity studies demonstrated elevated insulin level and corresponding decreased blood glucose level in diabetic rats. Thus, the delivery systems released insulin at a controlled rate in vitro in conformationally and chemically stable form and in vivo in biologically active form up to 3 months. © 2011 Wiley-Liss, Inc. and the American Pharmacists Association J Pharm Sci 100:4790–4803, 2011 Keywords: thermosensitive; triblock copolymer; insulin; in vitro release; conformational and chemical stability; in vivo absorption; Calorimetry (DSC); Circular Dichroism; Controlled release; Protein Delivery; Biocompatibility
INTRODUCTION Polymeric delivery systems have been studied extensively for the controlled release of insulin. Insulinloaded albumin microbead implants could release insulin in diabetic rats for up to 3 weeks.1 However, the microbeads were not injectable and had to be implanted surgically. Biodegradable, injectable microparticles consisting of poly(D,L-lactide) (PLA) or poly(D,L-lactide-co-glycolide) (PLGA) have been widely studied for controlled delivery of insulin.2–4 All formulations exhibited slow release of insulin preceded by a high initial burst. Alginate microspheres Correspondence to: Jagdish Singh (Telephone: +701-231-7943; Fax: +701-231-8333; E-mail: [email protected]) Khaled Al-Tahami’s present address is Department of Pharmaceutics, College of Medical Sciences, University of Science and Technology, Sana’a 2580, Yemen. Journal of Pharmaceutical Sciences, Vol. 100, 4790–4803 (2011) © 2011 Wiley-Liss, Inc. and the American Pharmacists Association
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could control the release of insulin up to 24 h only.5 Alginate–dextran nanospheres showed rapid release of insulin and glucose, lowering activity up to 8 h.6 Biphasic release profile characterized by a high initial release was observed from microspheres prepared from chitosan.7,8 In addition to the above problems, microspheres suffer several disadvantages such as complicated manufacturing procedure, stability of insulin during preparation, low-drug-loading capacity, and use of organic solvents for the preparation of the microspheres also raise environmental concerns. A number of thermosensitive polymeric delivery systems have been studied for controlled release of hormones and proteins such as growth hormone, testosterone, levonorgestrel, salmon calcitonin, lysozyme, and so on.9–14 PLGA–polyethylene glycol (PEG)–PLGA-based triblock copolymers have shown better control of release of hydrophobic drugs than hydrophilic drugs. Delivery systems containing
JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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PLGA–PEG–PLGA thermosensitive triblock copolymer showed a controlled release of various proteins including insulin for 2 weeks.15,16 In case of formulations containing insulin, the release profile exhibited a high burst release, most probably due to the high content of the hydrophilic glycolic acid blocks.16 Thus, there is a need to develop triblock copolymers that can control the release of insulin for longer duration. Presence of methyl groups imparts more hydrophobic character to lactic acid in comparison with glycolic acid and hence, PLA is more hydrophobic than PLGA, thereby hydrating to a lesser extent and undergoing slower hydrolysis.17,18 Therefore, triblock copolymers were synthesized by increasing the block length of PLA and PEG in the PLA–PEG–PLA triblock copolymers. We found that the PLA and PEG block lengths and copolymer concentration could alter the in vitro release of an exemplar protein, lysozyme. Increasing PLA chain length lowered the burst release and extended the overall release period, whereas increasing PEG chain length led to the opposite effects. Increasing the polymer concentration in delivery system led to further reduction of burst release and prolongation of the release period.19 Delivery systems composed of copolymer A (1496–1500–1496) and B (1584–1500–1584) exhibited a zero-order release kinetics of incorporated lysozyme, which is desired for controlled delivery systems. Therefore, copolymers A and B were investigated for the controlled delivery of insulin. One of the most encountered challenges when developing delivery systems for peptides and proteins is the conservation of proteins’ physical and chemical stability. Proteins undergo rapid denaturation and lose their conformational (noncovalent) stability and bioactivity in the environment other than their physiological ones.20 A number of physicochemical conditions, including hydrogen bonding and increased hydrophobic interactions, restricted N-terminus mobility as well as greater molecular packing, leading to increase stability of proteins.21,22 The investigation of protein unfolding had led to the conclusion that a loss of the native conformation of a protein has a drastic impact on its oxidation,23 deamidation,24 and aggregation.25 Thus, conformational and chemical stability of the released insulin from the delivery systems is warranted. Basal insulin is secreted continuously at a constant rate of 0.5–1 U/h, leading to serum concentrations of 5–15 :U/mL.26,27 Even though this level of insulin is low, it plays an important role in preventing or postponing the long-term complications of diabetes. Recently, we demonstrated that the zinc addition to insulin (1:5 insulin hexamer–zinc ion) during the preparation of the delivery system enhanced insulin stability during release.28 In the present study, we DOI 10.1002/jps
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investigated the thermosensitive polymeric delivery system for controlled insulin delivery in vitro and in vivo in diabetic rats to meet the basal insulin requirements. The overall structural stability, which comprises conformational and chemical stability, of the in vitro-released insulin was evaluated. Finally, in vitro and in vivo biocompatibility of thermosensitive systems were also studied.
EXPERIMENTAL Materials D,L-Lactide
was obtained from Toyo Kasei Kogyo Company Ltd. (Tokyo, Japan). Polyethylene glycol (PEG 1500) and stannous ethylhexanoate were purchased from Sigma Chemical Company (St. Louis, Missouri). Human recombinant insulin and microBCA protein assay kit were obtained from Celliance Corporation (Norcross, Georgia) and Pierce Biotechnology Inc. (Rockford, Illinois), respectively. Human insulin enzyme-linked immunosorbent assay (ELISA) kit was obtained from Linco Research Inc. (St. Charles, Missouri). Synthesis and Characterization of Triblock Copolymers
The PLA–PEG–PLA copolymers were synthesized using a method previously described by Singh et al.15 PEG block length of 1500 and D, L-lactide were used for synthesis. Under nitrogen, PEG was dried at 100◦ C in a three-necked flask for 30 min, followed by addition of D,L-lactide. After melting of the reactants, 0.03% (w/w) of stannous 2-ethylhexanoate was added, and the reaction was further carried out at 120◦ C for 12 h. Ice-cold water (5◦ C–8◦ C) was used to dissolve the crude polymer, and it was then reheated for purification. This step was repeated to improve the purity of the copolymer by removing any unreacted monomers. Finally, freeze-drying was performed to remove residual water. The synthesized copolymers were characterized by 1 H NMR spectrometry and Gel Permeation Chromatography (GPC) in order to determine their structure, molecular weight, molecular weight distribution, and purity. Preparation of Thermosensitive Delivery Systems of Insulin and In Vitro Release Aqueous solutions of the copolymer A (molecular weight 1496–1500–1496 Da) or copolymer B (molecular weight 1584–1500–1584 Da) were prepared by dissolving the copolymers in deionized water to obtain final concentrations of 30% and 40% (w/w). Insulin was dispersed in aqueous copolymer solution and the mixture was homogenized for 30 s at room temperature. Unless otherwise stated, zinc acetate was added to insulin delivery systems at a molar ratio of 1:5 insulin JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Table 1. Composition of Insulin Delivery Systems Based on Copolymer A and B Thermosensitive Triblock Copolymers and In Vitro Release Kinetics
Formulation
Copolymer
Copolymer Concentration (%, w/w)
Insulin Hexamer: Metal Ion
Insulin (%, w/v)
Zero Order (r2 )
First Order (r2 )
Higuchi (r2 )
A A A B B B
30 40 40 30 40 40
1:5 1:5 – 1:5 1:5 1:5
0.5 0.5 0.5 0.5 0.5 2.5
0.97 0.99 0.95 0.97 0.99 0.95
0.82 0.86 0.79 0.79 0.82 0.77
0.99 0.97 0.99 0.99 0.95 0.97
1 2 3 4 5 6
hexamer to zinc ion. The composition of the delivery systems is presented in Table 1. A delivery system (1 mL) was injected into a 20 mL polypropylene tube and incubated in a reciprocating water bath maintained at 37◦ C. After formation of the gel, 15 mL of prewarmed phosphate-buffered saline (PBS), pH 7.4 was added to the test tube as release medium. For the entire study duration, the test tube was incubated in a reciprocating water bath maintained at 37◦ C and 35 rpm. At different time points, 5 mL samples were withdrawn from the dissolution media. Fresh PBS was used to replace the removed volume. After centrifugation at 4229g for 30 min, samples were diluted with PBS. The released insulin was quantified using microBCA protein assay (Pierce Biotechnology Inc.). Stability Studies
Differential Scanning Calorimetry An ultrasensitive differential scanning calorimeter (VP-DSC, MicroCal, Northampton, Massachusetts) was used to study the conformational stability of the in vitro-released insulin. The sample preparation method described previously by Al-Tahami et al.28 has been used. Samples were centrifuged, filtered, and degassed under vacuum. Experiments were carried out using a scan rate of 1◦ C/min over a temperature range of 25◦ C–105◦ C. The thermodynamic parameters determined were the calorimetric enthalpy (H) and R midpoint transition temperature (Tm ). The Origin software (MicroCal) was used for data analysis.
Circular Dichroism Jasco J-815 circular dichroism (CD) spectrophotometer (Jasco, Tokyo, Japan) was used to evaluate the secondary structure of insulin. We have earlier described the details about the sample preparation method.28 Spectra of buffer and insulin were scanned from 200 to 300 nm. The spectra of the buffer were used as baseline and were subsequently subtracted from insulin sample spectra. MicroBCA assay (Pierce Biotechnology Inc.) was used to determine the concentration of insulin in order to calculate the molar ellipticity (θ). JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
Matrix-Assisted Laser Desorption/ Ionization–Time-of-Flight Mass Spectrometry A Bruker matrix-assisted laser desorption/ ionization–time-of-flight (MALDI–TOF) II mass spectrometer (Bruker Daltonics Inc., Billercia, Massachusetts) equipped with smartbeam in combination with a solid-state laser (200 Hz) was used to study the chemical stability of insulin. The details of the method can be found in an earlier publication.28 R FlexAnalysis software (Bruker Daltonics Inc., Billercia, MA) was used to analyze the data.
In Vitro and In Vivo Biocompatibility An MTT assay was used to determine in vitro biocompatibility of in situ gel depot in human embryonic kidney (HEK293) cells. The detail methods for in vitro and in vivo biocompatibility are described in the previous publication.28 Briefly, the polymers were extracted in PBS maintained at 37◦ C and 70◦ C for 10 days. As degradation rate of polymers is greater at higher temperature, the extract prepared at 70◦ C simulates the long-term effects of the delivery system. The cytotoxicity of the polymeric extracts, growth medium without polymer (negative control), and 2% dimethyl sulfoxide in growth medium (positive control) was tested in HEK293 cells. The in vivo biocompatibility of the delivery systems was determined by studying inflammatory changes in the rat skin. The formulation was administered to the rats subcutaneously into the ventral side of the neck. Skin sections from injection sites were surgically removed after euthanizing the rats at day 1, 30, and 90 time points. The skin samples were prepared for histological analysis and studied for the signs of inflammation, necrosis, or any changes in the morphology.
In Vivo Absorption and Bioactivity Studies In vivo absorption and bioactivity of insulin was tested by measuring blood glucose reduction in diabetic male Sprague–Dawley (SD) rats (200–224 g body weight). The protocol for animal studies was approved by Institutional Animal Care and Use Committee at North Dakota State University, Fargo, North Dakota. The animal experiments were carried out according to the guidelines and humane animal care standards DOI 10.1002/jps
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described in the National Institutes of Health Guide for the Care and Use of Experimental Animals. A single intraperitoneal dose of streptozotocin (STZ; 60 mg/ kg) in citrate buffer (pH 4.5) was administered to inR (Bayer Corporation, duce diabetes. Glucometer Elite Elkhart, Indiana) was used to estimate blood glucose level. R , Ovation PharPentobarbital sodium (Nembutal maceuticals Inc., Deerfield, IL) at a dose of 30 mg/kg body weight was injected intraperitoneally to anesthetize the rats. The temperature-sensitive delivery system was injected subcutaneously at doses of 90 insulin unit (U)/kg body weight at the neck region with the help of 25-gauge needle. An aqueous solution of insulin in PBS (2 U/kg body weight) was also studied. The control was polymeric solution without insulin. Blood was withdrawn at regular intervals from tail vein. Insulin serum concentration and blood glucose level were determined by ELISA and glucose oxidase method, respectively. The details of the method have been described in an earlier publication.28 Enzyme immunoassay (human insulin ELISA kit; Linco Research Inc.) was performed for insulin according to the manufacturer’s instructions.29 Briefly, the plates, samples, and insulin standards (2–200 :U/mL) were prepared as instructed. Insulin standards and samples were added to wells followed by detection antibody solution. Plates were covered and incubated with mixing at room temperature. Solutions were discarded and plates were washed thoroughly with horseradish peroxidase (HRP) wash buffer. Enzyme solution was then added to each well followed by incubation at room temperature. Plates were washed again with HRP buffer after decanting the enzyme solutions. Substrate solution (3,3 ,5,5 -tertramethylbenzidine) was added to wells, and plates were incubated until blue color was developed. The reaction was stopped using ELISA stop solution (0.3 M HCl). The blue color turned yellow after acidification and absorbance was read at 450 nm immediately. The dose-absorbance curve of insulin standards was fitted to a sigmoidal logistic equation and used to determine the concentration of insulin in samples. After completion of the experiments, rats were euthanized by intraperitoneal injection of pentobarbital (150 mg/kg body weight).
Data Analysis Analysis of variance (ANOVA) and Student’s t-test were used for computing the results. A p value of 0.05 was taken as significant. The serum insulin concentration versus time profile was used to determine the maximum serum concentration (Cmax ), time to Cmax (Tmax ), and area under the curve (AUC).30 DOI 10.1002/jps
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Figure 1. Effect of zinc addition on in vitro release of insulin from 40% (w/w) copolymer A (1496–1500–1496) (n = 4). Key: (䊏) zinc acetate and () no zinc salt added. [Key: ∗ , significant compared with 1:5 insulin hexamer to zinc ion at p < 0.05]
RESULTS Synthesis and Characterization of Copolymers Proton nuclear magnetic resonance (NMR) was used to determine the structure and molecular weight of the copolymers. The chemical shift (*) signals originating from different protons confirmed the structure of the synthesized copolymer. The spectrum was similar as reported in the literature by Jeong et al.31 The synthesized copolymers A and B showed narrow molecular weight distribution (polydispersity index 1.1). The molecular weights were determined as described by Jeong et al.31 using signal integration method in proton NMR, and was found to be 1496–1500–1496 and 1584–1500–1584 for copolymer A and B, respectively.
In Vitro Release of Insulin Table 1 lists the composition and release kinetics of insulin from delivery systems made of copolymer A and B. The effect of metal ion addition and polymer concentration on insulin release from copolymer A-based thermosensitive delivery systems are shown in Figures 1 and 2, respectively. Figure 1 demonstrates that the addition of zinc lowered the burst release following zero-order release kinetics (r2 ∼0.99) in comparison with formulation without zinc. The release of insulin was controlled for 2 months. Increasing the polymer concentration from 30% to 40% (w/w) reduced the burst release and improved release kinetics (Fig. 2 and Table 1). Similar results were observed when copolymer B (1584–1500–1584) was used (Fig. 3 and Table 1). Increasing chain length of PLA in copolymer B (1584–1500–1584) compared with copolymer A (1496–1500–1496) extended the release period. The insulin was released from delivery JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 2. Effect of polymer concentration on in vitro release of insulin from triblock copolymer A (1496–1500–1496) (n = 4). Key: copolymer concentration of (䊏) 30% and (•) 40% (w/w). [Key: ∗ , significant compared with 40% (w/w) copolymer A at p < 0.05]
systems containing 40% (w/w) copolymer B over 3 months with best fit for zero-order release kinetics (r2 ∼0.99). Increasing the insulin loading dose from 5 (0.5%, w/v) to 25 mg (2.5%, w/v) led to a small increase in burst release (Fig. 4). Stability Studies Differential scanning calorimetry studies showed that the insulin released after 1 month had lower H and Tm compared with fresh insulin (Fig. 5). Fresh and in vitro-released insulin showed Tm values of 82.49 ± 0.92 and 83.02 ± 0.13, and H values of 14.3 ± 0.7 and 10.1 ± 0.5 Cal/mol × 103 , respectively. CD spectra for fresh and insulin released in vitro after 1 and 2 months is represented in Figure 6. The magnitude of minima at 209 and 222 nm was decreased in the released insulin samples as com-
Figure 3. Effect of polymer concentration on in vitro release of insulin from triblock copolymer B (1584–1500–1584) (n = 4). Key: copolymer concentration of (䊏) 30% and (•) 40% (w/w). [Key: ∗ , significant compared with 40% (w/w) copolymer B at p < 0.05] JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
Figure 4. Effect of insulin loading on the in vitro release of insulin from triblock copolymer B (1584–1500–1584) (n = 4). Key: Insulin amount incorporated (䊏) 5 mg and (䉬) 25 mg. [Key: ∗ , significant compared with 5 mg insulin at p < 0.05]
pared with fresh insulin. However, the presence of the two minima confirms that the secondary structure of insulin was relatively conserved during delivery system preparation and release. The structural integrity of released insulin from thermosensitive delivery systems was evaluated using MALDI–TOF mass spectroscopy. Figure 7 shows the MALDI spectra of fresh insulin, 1 and 2 months released insulin from the delivery system containing 40% (w/w) copolymer B. MALDI–TOF (M+H)+ signal at 5808.6 Da indicated the conservation of primary structure of insulin in the released samples. No major degradation products were detected. There was no evidence of formation of acylation products (MW+ 72) or high-molecular weight transformation products. However, when the released insulin spectra were enlarged, a signal corresponding to a molecular weight of 5792.9 Da was observed (Figure 7b and 7c). As mentioned previously, this indicates a loss of a water molecule from the insulin, suggesting the formation of cyclic imide product that usually follows deamidation reactions.28 The insulin released after 2 months (Fig. 7c) showed a higher peak for the 5792 Da signal, implying that further deamidation took place. The 5792 Da signal relative intensity percentage (compared with 5808 Da signal, which corresponds to the native insulin) was equal to 13.7% and 37.8% in released insulin samples withdrawn after 1 and 2 months, respectively. This increase in deamidation products could be caused by the prolonged contact of insulin with the microenvironment inside the depot or the release medium. Two additional signals were observed in 2 months released samples corresponding to molecular weights of 5217.1 and 5070.2 Da (Fig. 7c). These two signals appear to be products of peptide-bond hydrolysis, which led to the formation of lower molecular weight peptides. DOI 10.1002/jps
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Figure 5. DSC thermograms of released insulin after 1 month from copolymer B thermosensitive system (2.5%, w/v insulin): Fresh insulin (solid line), and released insulin (dotted line), released insulin (without zinc) (dashed line).
Biocompatibility of Polymeric Delivery Systems
In Vitro Biocompatibility Figures 8 and 9 show the results of MTT cell viability assay for thermosensitive delivery system extracts prepared from copolymer A and copolymer B, respectively. No significant difference (p > 0.05) in the cell viability was observed between polymers extract groups and PBS group for all the dilutions and exposure time
Figure 6. CD spectra of released insulin from copolymer B thermosensitive system (2.5%, w/v insulin): Fresh insulin (solid line), 1 month released insulin (dashed line), and 2 months released insulin (dotted line). DOI 10.1002/jps
points. Interestingly, polymer extract showed greater cell survival compared with PBS group at certain dilutions (copolymer B; 1:8 dilution; 72 h exposure).
In Vivo Biocompatibility The inflammatory response of rat skin to temperature-sensitive delivery systems was evaluated at different time points after subcutaneous administration. Figures 10a and 11a show light microscopy images of skin sample taken from control rats where no formulation was injected. Figures 10b– 10d represent light microscopy images of skin tissues from where temperature-sensitive-based delivery systems (copolymer A) were administered. Similarly, Figures 11b–11d show light microscopy images of skin tissue injected with copolymer B containing delivery system. Both polymeric systems demonstrated inflammatory reactions and tissue healing process after injection. Following administration, inflammatory cells (mainly neutrophiles and lymphocytes) infiltrate to injection site. After a month, number of macrophages dropped, which are components of chronic tissue inflammation in response to injury.8 Skin samples from 3 months time point were comparable to control tissue. The connective and muscular tissues of the injection sites appeared normal and comparable to control skin tissue. We did not observe any muscle damage or necrotic damage throughout JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 7. MALDI–TOF mass spectroscopy of (a) fresh insulin (control), (b) 1 month and (c) 2 months from thermosensitive delivery system (40%, w/w copolymer B).
the studies, denoting the biocompatible nature of the delivery systems.
In Vivo Absorption and Bioactivity Figure 12 shows the serum insulin concentration and blood glucose level following subcutaneous administration of insulin (2 U/kg body weight) dissolved in PBS (pH 7.4) in diabetic SD rats. Serum insulin level increased rapidly reaching mean peak concentration (Cmax ) of 67.84 :U/mL at 2 h after administration and declined afterward to reach below 2 :U/mL (detection limit) after 12 h. Blood glucose levels decreased following insulin absorption, showing an acute and relatively short hypoglycemic effect. Blood glucose levels were restored to preadministration levels within 8 h. In vivo absorption and biological activity of insulin are presented in Figure 13. The control group composed of copolymer without insulin. Treatment groups were injected with insulin (90 U/kg body weight) dissolved in 40% (w/w) of either copolymer A or copolymer B. Zinc was added to both treatment groups as zinc acetate (1:5 ratio). Treatment groups showed continuous insulin release from thermosensitive delivery systems over a period of 3 months. The more hydrophobic copolymer B showed a lower release rate JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
for longer duration compared with copolymer A. Blood glucose levels in rats receiving either of the delivery systems were significantly lower (p < 0.05) than the diabetic rats in control group. The bioavailability of insulin in rats was calculated in terms of AUC values after correcting for dose (Table 2). When compared with subcutaneous insulin solution, copolymer B-based delivery system enhanced insulin bioavailability by approximately 11-fold.
DISCUSSION In vitro release studies showed that increasing the chain length of PLA in copolymer resulted in lowering the burst release of insulin. Because of the more hydrophobicity of lactic acid, the copolymer absorbed less water and degraded slowly.17,32 Similar results have been reported when PLGA–PEG–PLGA thermosensitive delivery systems with varying LA/GA ratios were used for the delivery of 5-fluorouracil.33 Furthermore, it was shown that increasing LA/GA ratio decreased burst release and extended the release period of bovine serum albumin from PLGA microspheres.34,35 DOI 10.1002/jps
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Figure 8. In vitro biocompatibility of thermosensitive copolymer A (1496–1500–1496) extracts (a–c: Extracts prepared at 37◦ C and d–f: Extracts prepared at 70◦ C) (n = 8). [PBS, phosphatebuffered saline; PE,polymer extract; DMSO, dimethyl Sulfoxide (2%)].
The release kinetics for delivery systems containing 30% (w/w) copolymer A showed best fit for Higuchi model (r2 = 0.97), indicating a diffusion dominant initial release phase. The release showed a best fit for zero-order release kinetics (r2 = 0.99) as the concentration of copolymer increased to 40% (w/w). This may be attributed to the reduced diffusion due to the less porous depot at higher copolymer concentration. We speculate that the addition of metal ions reduced insulin association in PEG domain by lowering insulin solubility. Thus, less insulin was available for the initial burst release and most of the incorporated DOI 10.1002/jps
insulin was associated with PLA and was released during the degradation phase. Thermosensitive delivery systems containing 40% (w/w) of copolymer A and B, to which zinc ions were added at the molar ratio 1:5 insulin hexamer to zinc ion, showed the best fit for zero-order (r2 = 0.99) release kinetics of insulin over periods of 2 and 3 months, respectively. An important aspect in the use of controlled-release systems as proteins delivery devices is the ability of loading different doses of proteins. This is mainly to count for intraindividual and interindividual drug requirement variations. Increasing insulin loading in JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 9. In vitro biocompatibility of thermosensitive copolymer B (1584–1500–1584) extracts (a–c: Extracts prepared at 37◦ C and d–f: Extracts prepared at 70◦ C) (n = 8). [PBS, phosphatebuffered saline; PE, polymer extract; DMSO, dimethyl sulfoxide (2%)].
delivery systems containing copolymer B from 5 to 25 mg led to an increase in insulin burst release and best fit for release kinetics was shifted from zero to pseudo-zero-order model. Yet, the release profile exhibited a high-zero-order correlation (r2 = 0.95). The DSC results confirmed that the addition of divalent metal ion stabilizes insulin, indicated by the higher denaturation enthalpy. Zinc has a strong inhibitory effect on fibril formation/growth due to the JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
neutralization of the charges at the core of the insulin hexameric assembly, leading to the stabilization of these hexamers. This leads to an increase in the hydrophobic interactions within insulin hexamer, thus, increasing its stability. Numerous studies have shown that insulin denaturation is highly dependent on insulin being in the monomeric form in solutions.36 It has also been reported that complexation with zinc was able to stabilize insulin during release from DOI 10.1002/jps
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Figure 10. Light micrographs of skin taken from injection site (a) Control, (b) after 1 day, (c) 1 month, and (d) 3 months after subcutaneous administration of thermosensitive delivery system (copolymer A). Acute inflammatory cells infiltration (N, neutrophiles and L, lymphocytes) between muscle fibers and hemorrhage in injection site is apparent. [Magnification: 20×].
Figure 11. Light micrographs of skin taken from injection site (a) Control, (b) after 1 day, (c) 1 month, and (d) 3 months after subcutaneous administration of thermosensitive delivery system (copolymer B). Acute inflammatory cells infiltration (N, neutrophiles and L, lymphocytes) between muscle fibers and hemorrhage in injection site is apparent. [Magnification: 20×].
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Figure 12. Serum insulin concentration (a) and blood glucose level (b) following subcutaneous administration of insulin solution (n = 6).
microspheres and also prevented nonspecific adsorption to hydrophobic polymer surfaces.37 The CD spectrum of insulin had been reported to exhibit two minima at 209 and 223 nm,37 indicating a dominant "-helical structure (∼38%) with a small part of betasheets (∼17%). The two minima were greatly attenuated for preparations containing no zinc ion after 28 days of incubation ("-helix ∼18% and $-sheets ∼4%) in comparison with preparations containing 1:5 insulin hexamer to zinc ion ratio ("-helix ∼36% and $-sheets ∼14%). This attenuation denotes the aggregation process and the loss of bioactive form of insulin from the solution. The conformational stability and secondary structure of released insulin was partially reduced, which may be due to prolonged exposure of released insulin in agitated media at 37◦ C temperature.36,38–40 The correlation between protein unfolding and chemical stability is not straightforward as many facJOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
tors affect protein chemical stability. The chemical degradation, which takes place in protein domains that are protected in native state, occurs at higher rates when the protein is denatured.41 It has been shown that there is a high correlation between DSC results and insulin B3 deamidation products.42 However, the DSC data failed to predict the rate of covalent insulin dimer (CID) formation. In addition, because DSC measures thermal stability, it cannot be used to predict chemical stability at low temperatures. Insulin undergoes many chemical modification reactions depending on the surrounding environment and that can affect its biological activity. Biological activity of native insulin and its deamidation products is similar, whereas the products formed due to hydrolysis (A8–A9) and CIDs exhibit only 2% and 15% relative potency, respectively.43 The formation of CID and deamidation products of insulin has been DOI 10.1002/jps
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Figure 13. Serum insulin concentration (a) and blood glucose level (b) following subcutaneous administration of thermosensitive polymeric delivery systems in SD rats (n = 6). Key: (䊏) blank thermosensitive delivery systems (control), (•) thermosensitive delivery systems composed of 40% (w/w) copolymer A (1496–1500–1496), and () thermosensitive delivery systems composed of 40% (w/w) copolymer B (1584–1500–1584).
reported to be increased by temperature and shaking.40 MALDI results from 1 month-released samples from the delivery systems did not show peaks for acylation, CID, or hydrolysis products and the spectra were comparable to fresh insulin spectrum. A small signal denoting the formation of cyclic imide Table 2.
by the loss of a water molecule was observed. Iso–Asp insulin is formed due to the deamidation reaction, which follows cyclic imide formation.44 However; insulin from the 2 months-released sample exhibited an increase in the cyclic imide signal, which denotes an increase in deamidation reactions. Additional
In Vivo Pharmacokinetic Parameters of Delivered Insulin in Rats
Group SG T1 T2
Cmax (:U/mL)
Tmax (h)
AUC [:U/(h mL)]
Increase in AUC (Folds Compared with SC)
67.84 ± 7.38 – –
1.00 – –
5493.83 ± 1263.72 41709.90 ± 8116.30a 59924.41 ± 11279.68a
– 7.59 10.91
SG, insulin solution group; T1, thermosensitive delivery systems—copolymer A (1496–1500–1496); T2, thermosensitive delivery systems—copolymer B (1584–1500–1584). a Significantly (p < 0.05) higher than solution group.
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hydrolysis degradation products signals (5217.1 and 5070.2 Da), which correspond to B24–B25 and B25–B26 split products, respectively, were observed. As discussed previously, this could be due to the prolonged presence of insulin in the shaking release media at 37◦ C. As insulin released from the delivery system is absorbed instantly, we do not anticipate this to happen in vivo. It was extremely crucial to investigate the biocompatibility of the delivery systems used in this study. Neutralized extracts of thermosensitive delivery systems did not exhibit any inhibitory effect on cell viability. On the contrary, at some time points, cell viability was greater in polymer extract groups. It has been suggested that this could be due to the increase in lactic acid, which is a metabolite in the citric acid cycle.45 In vitro biocompatibility is a very good indication for predicting in vivo biocompatibility. In the beginning, there was sign of inflammation reaction due to injury caused by injection.9 After 3 months of injection of the delivery system, all signs of inflammatory responses disappeared and skin histology of injected site was comparable to control skin samples. The findings of in vivo biocompatibility were in agreement with the in vitro biocompatibility and International Organization for Standardization (ISO) regulations, supporting the biocompatible nature of the thermosensitive copolymers as a delivery system. The success of insulin continuous delivery at basal level depends on the delivery system to deliver insulin in vivo at controlled rate in biologically active form. In this study, diabetic model was created by injecting STZ into male SD rats intraperitoneally. STZ is a glucomimetic molecule and employs glucose transporter GLUT2 for its transport to pancreatic $ cells where the mechanism of the cell death includes the nitrosourea-induced DNA alkylation.46 There was an increase in pharmacological bioavailability of insulin from the copolymer-based delivery systems in comparison with subcutaneous insulin solution in PBS. It has been shown that insulin absorption rates from subcutaneous tissue are inversely correlated to insulin concentration.47 The slower release of insulin from the delivery system led to further absorption of insulin from the releasing site, thereby less likely to degrade by proteases present in the skin.48 The polymeric delivery system has also offered additional stabilizing effect on insulin by protecting it from enzymatic degradation at the injection site. It is apparent from the data that the pharmacological activity of insulin released was preserved in vivo. The duration of in vivo release was shorter in comparison with in vitro release due to the presence of enzymes in vivo, leading to faster polymer degradation of polymer.13–14,49,50 But, in vivo release seems to be consistent with the blood glucose profile and JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
the insulin release pattern observed in the in vitro release.
CONCLUSION The present work shows that the temperaturesensitive triblock copolymers can be used for the controlled delivery of insulin at basal level in a conformationally and chemically stable and bioactive form. Insulin release can be altered by changing polymer composition, polymer concentration, and addition of metal ions. Release pattern reliability on formulation factors is advantageous as it allows for delivery of different compounds with different requirements. The stabilizing effect offered by the polymeric implant and metal ion addition was also demonstrated.
ACKNOWLEDGMENTS We acknowledge the financial support from NIH grant #HD056053. We also acknowledge the Fulbright Program for scholarship to Khaled Al-Tahami. The assistance provided by Buddhadev Layek in the preparation of the manuscript is greatly appreciated.
REFERENCES 1. Goosen MF, Leung YF, O’Shea GM, Chou S, Sun AM. 1983. Slow release of insulin from a biodegradable matrix implanted in diabetic rats. Diabetes 32(5):478–481. 2. Kang F, Singh J. 2005. In vitro release of insulin and biocompatibility of in situ forming gel systems. Int J Pharm 304(1–2):83–90. 3. Liu J, Zhang SM, Chen PP, Cheng L, Zhou W, Tang WX, Chen ZW, Ke CM. 2007. Controlled release of insulin from PLGA nanoparticles embedded within PVA hydrogels. J Mater Sci Mater Med 18(11):2205–2210. 4. Takenaga M, Yamaguchi Y, Kitagawa A, Ogawa Y, Mizushima Y, Igarashi R. 2002. A novel insulin formulation can keep providing steady levels of insulin for much longer periods in-vivo. J Pharm Pharmacol 54:1189–1194. 5. Silva CM, Ribeiro AJ, Figueiredo IV, Gonc¸alves AR, Veiga F. 2006. Alginate microspheres prepared by internal gelation: Development and effect on insulin stability. Int J Pharm 311(1–2):1–10. 6. Reis CP, Ribeiro AJ, Houng S, Veiga F, Neufeld RJ. 2007. Nanoparticulate delivery system for insulin: Design, characterization and in vitro/in vivo bioactivity. Eur J Pharm Sci 30(5):392–397. 7. Wang L, Gu Y, Su Z, Ma G. 2006. Preparation and improvement of release behavior of chitosan microspheres containing insulin. Int J Pharm 311(1–2):187–195. 8. Ubaidulla U, Khar RK, Ahmad FJ, Sultana Y, Panda AK. 2007. Development and characterization of chitosan succinate microspheres for the improved oral bioavailability of insulin. J Pharm Sci 96(11):3010–3023. 9. Tang Y, Singh J. 2009. Biodegradable and biocompatible thermosensitive polymer based injectable implant for controlled release of protein. Int J Pharm 365(1–2):34–43. 10. Chen S, Singh J. 2008. Controlled release of growth hormone from thermosensitive triblock copolymer systems: In vitro and in vivo evaluation. Int J Pharm 352(1–2):58–65. DOI 10.1002/jps
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11. Chen S, Singh J. 2005. In vitro release of levonorgestrel from phase sensitive and thermosensitive smart polymer delivery systems. Pharm Dev Technol 10(2):319–325. 12. Chen S, Pieper R, Webster DC, Singh J. 2005. Triblock copolymers: Synthesis, characterization, and delivery of a model protein. Int J Pharm 288(2):207–218. 13. Chen S, Pederson D, Oak M, Singh J. 2010. In vivo absorption of steroidal hormones from smart polymer based delivery systems. J Pharm Sci 99(8):3381–3388. 14. Tang Y, Singh J. 2010. Thermosensitive drug delivery system of salmon calcitonin: In vitro release, in vivo absorption, bioactivity and therapeutic efficacies. Pharm Res 27(2):272– 284. 15. Singh S, Webster DC, Singh J. 2007. Thermosensitive polymers: Synthesis, characterization, and delivery of proteins. Int J Pharm 341(1-2):68–77. 16. Kim YJ, Choi S, Koh JJ, Lee M, Ko KS, Kim SW. 2001. Controlled release of insulin from injectable biodegradable triblock copolymer. Pharm Res 18(4):548–550. 17. Cohen S, Alonso MJ, Langer R. 1994. Novel approaches to controlled-release antigen delivery. Int J Technol Assess Health Care 10(1):121–130. 18. Ikada Y, Hideto T. 2000. Biodegradable polyesters for medical and ecological applications. Macromol Rapid Commun 21:117–132. 19. Al-Tahami K. 2007. In situ gel forming smart polymers for the controlled release of proteins. Fargo, North Dakota:Doctoral Dissertation, North Dakota State University. 20. Jen A, Merkle HP. 2001. Diamonds in the rough: Protein crystals from a formulation perspective. Pharm Res 18(11):1483–1488. 21. Cowan DA. 1997. Thermophilic proteins: Stability and function in aqueous and organic solvents. Comp Biochem Physiol A Physiol 118(3):429–438. 22. Vogt G, Argos P. 1997. Protein thermal stability: Hydrogen bonds or internal packing? Fold Des 2(4):S40–S46. 23. Komsa-Penkova R, Koynova R, Kostov G, Tenchov B. 2000. Discrete reduction of type I collagen thermal stability upon oxidation. Biophys Chem 83(3):185–195. 24. Pikal MJ, Rigsbee DR. 1997. The stability of insulin in crystalline and amorphous solids: Observation of greater stability for the amorphous form. Pharm Res 14(10):1379–1387. 25. Azuaga AI, Dobson CM, Mateo PL, Conejero-Lara F. 2002. Unfolding and aggregation during the thermal denaturation of streptokinase. Eur J Biochem 269(16):4121–4133. 26. Galloway JA, Chance RE. 1994. Improving insulin therapy: Achievements and challenges. Horm Metab Res 26(12):591–598. 27. Fonseca V. 2006. The role of basal insulin therapy in patients with type 2 diabetes mellitus. Insulin 1:51–60. 28. Al-Tahami K, Oak M, Singh J. 2011. Controlled delivery of basal insulin from phase-sensitive polymeric systems after subcutaneous administration: In vitro release, stability, biocompatibility, in vivo absorption, and bioactivity of insulin. J Pharm Sci 100(6):2161–2171. 29. Human Insulin ELISA. 2006. User’s Guide. St. Charles, Missouri: Linco Research, Inc. 30. Rowland M, Tozer T. 1980. Interacting drugs. In Clinical pharmacokinetics: Concepts and applications; Rowland M, Tozer TN, Eds. Malvern, Pennsylvania: Lea and Febiger, pp 256–275. 31. Jeong B, Bae YH, Kim SW. 1999. Biodegradable thermosensitive micelles of PEG–PLGA–PEG triblock copolymers. Colloids and Surfaces B: Biointerfaces 16 (1–4):185–193.
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32. Jalil R, Nixon JR. 1990. Biodegradable poly(lactic acid) and poly(lactide-co-glycolide) microcapsules: Problems associated with preparative techniques and release properties. J Microencapsul 7(3):297–325. 33. Qiao M, Chen D, Ma X, Liu Y. 2005. Injectable biodegradable temperature-responsive PLGA–PEG–PLGA copolymers: Synthesis and effect of copolymer composition on the drug release from the copolymer-based hydrogels. Int J Pharm 294(1–2):103–112. 34. Sah H, Chien YW. 1993. Evaluation of a microreservoir-type biodegradable microcapsule for controlled release of proteins. Drug Dev Indl Pharm 19(11):1243–1263. 35. Sah H, Toddywala R, Chien YW. 1994. The influence of biodegradable microcapsule formulations on the controlled release of a protein. J Control Release 30(3):201–211. 36. Huus K, Havelund S, Olsen HB, van de Weert M, Frokjaer S. 2005. Thermal dissociation and unfolding of insulin. Biochemistry 44(33):11171–11177. 37. Manoharan C, Singh J 2009. Evaluation of polyanhydride microspheres for basal insulin delivery: Effect of copolymer composition and zinc salt on encapsulation, in vitro release, stability, in vivo absorption and bioactivity in diabetic rats. J Pharm Sci 98(11):4237–4250. 38. Duncan G, Jess TJ, Mohamed F, Price NC, Kelly SM, van der Walle CF. 2005. The influence of protein solubilisation, conformation and size on the burst release from poly(lactideco-glycolide) microspheres. J Control Release 110(1):34–48. 39. Sluzky V, Tamada JA, Klibanov AM, Langer R. 1991. Kinetics of insulin aggregation in aqueous solutions upon agitation in the presence of hydrophobic surfaces. Proc Natl Acad Sci USA 88(21):9377–9381. 40. Sadhale Y, Shah JC. 1999. Stabilization of insulin against agitation-induced aggregation by the GMO cubic phase gel. Int J Pharm 191(1):51–64. 41. Kossiakoff AA. 1988. Tertiary structure is a principal determinant to protein deamidation. Science 240(4849):191–194. 42. Huus K, Havelund S, Olsen HB, van de Weert M, Frokjaer S 2006. Chemical and thermal stability of insulin: Effects of zinc and ligand binding to the insulin zinc-hexamer. Pharm Res 23(11):2611–2620. 43. Brange J, Langkjaer L, Havelund S, Vølund A. 1992. Chemical stability of insulin. 1. Hydrolytic degradation during storage of pharmaceutical preparations. Pharm Res 9(6):715–726. 44. Houchin ML, Heppert K, Topp EM. 2006. Deamidation, acylation and proteolysis of a model peptide in PLGA films. J Control Release 112(1):111–119. 45. Ignatius AA, Claes LE 1996. In vitro biocompatibility of bioresorbable polymers: Poly(l, dl-lactide) and poly(l-lactideco-glycolide). Biomaterials 17(8):831–839. 46. Elsner M, Guldbakke B, Tiedge M, Munday R, Lenzen S. 2000. Relative importance of transport and alkylation for pancreatic beta-cell toxicity of streptozotocin. Diabetologia 43(12):1528–1533. 47. Søeborg T, Rasmussen CH, Mosekilde E, Colding-Jørgensen M. 2009. Absorption kinetics of insulin after subcutaneous administration. Eur J Pharm Sci 36(1):78–90. 48. Lee VH. 1988. Enzymatic barriers to peptide and protein absorption. Crit Rev Ther Drug Carrier Syst 5(2):69–97. 49. Kwon YM, Kim SW. 2004. Biodegradable triblock copolymer microspheres based on thermosensitive sol-gel transition. Pharm Res 21(2):339–343. 50. Chen S, Singh J 2005. Controlled delivery of testosterone from smart polymer solution based systems: In vitro evaluation. Int J Pharm 295(1–2):183–190.
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INTRODUCTION Polymeric delivery systems have been studied extensively for the controlled release of insulin. Insulinloaded albumin microbead implants could release insulin in diabetic rats for up to 3 weeks.1 However, the microbeads were not injectable and had to be implanted surgically. Biodegradable, injectable microparticles consisting of poly(D,L-lactide) (PLA) or poly(D,L-lactide-co-glycolide) (PLGA) have been widely studied for controlled delivery of insulin.2–4 All formulations exhibited slow release of insulin preceded by a high initial burst. Alginate microspheres Correspondence to: Jagdish Singh (Telephone: +701-231-7943; Fax: +701-231-8333; E-mail: [email protected]) Khaled Al-Tahami’s present address is Department of Pharmaceutics, College of Medical Sciences, University of Science and Technology, Sana’a 2580, Yemen. Journal of Pharmaceutical Sciences, Vol. 100, 4790–4803 (2011) © 2011 Wiley-Liss, Inc. and the American Pharmacists Association
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could control the release of insulin up to 24 h only.5 Alginate–dextran nanospheres showed rapid release of insulin and glucose, lowering activity up to 8 h.6 Biphasic release profile characterized by a high initial release was observed from microspheres prepared from chitosan.7,8 In addition to the above problems, microspheres suffer several disadvantages such as complicated manufacturing procedure, stability of insulin during preparation, low-drug-loading capacity, and use of organic solvents for the preparation of the microspheres also raise environmental concerns. A number of thermosensitive polymeric delivery systems have been studied for controlled release of hormones and proteins such as growth hormone, testosterone, levonorgestrel, salmon calcitonin, lysozyme, and so on.9–14 PLGA–polyethylene glycol (PEG)–PLGA-based triblock copolymers have shown better control of release of hydrophobic drugs than hydrophilic drugs. Delivery systems containing
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PLGA–PEG–PLGA thermosensitive triblock copolymer showed a controlled release of various proteins including insulin for 2 weeks.15,16 In case of formulations containing insulin, the release profile exhibited a high burst release, most probably due to the high content of the hydrophilic glycolic acid blocks.16 Thus, there is a need to develop triblock copolymers that can control the release of insulin for longer duration. Presence of methyl groups imparts more hydrophobic character to lactic acid in comparison with glycolic acid and hence, PLA is more hydrophobic than PLGA, thereby hydrating to a lesser extent and undergoing slower hydrolysis.17,18 Therefore, triblock copolymers were synthesized by increasing the block length of PLA and PEG in the PLA–PEG–PLA triblock copolymers. We found that the PLA and PEG block lengths and copolymer concentration could alter the in vitro release of an exemplar protein, lysozyme. Increasing PLA chain length lowered the burst release and extended the overall release period, whereas increasing PEG chain length led to the opposite effects. Increasing the polymer concentration in delivery system led to further reduction of burst release and prolongation of the release period.19 Delivery systems composed of copolymer A (1496–1500–1496) and B (1584–1500–1584) exhibited a zero-order release kinetics of incorporated lysozyme, which is desired for controlled delivery systems. Therefore, copolymers A and B were investigated for the controlled delivery of insulin. One of the most encountered challenges when developing delivery systems for peptides and proteins is the conservation of proteins’ physical and chemical stability. Proteins undergo rapid denaturation and lose their conformational (noncovalent) stability and bioactivity in the environment other than their physiological ones.20 A number of physicochemical conditions, including hydrogen bonding and increased hydrophobic interactions, restricted N-terminus mobility as well as greater molecular packing, leading to increase stability of proteins.21,22 The investigation of protein unfolding had led to the conclusion that a loss of the native conformation of a protein has a drastic impact on its oxidation,23 deamidation,24 and aggregation.25 Thus, conformational and chemical stability of the released insulin from the delivery systems is warranted. Basal insulin is secreted continuously at a constant rate of 0.5–1 U/h, leading to serum concentrations of 5–15 :U/mL.26,27 Even though this level of insulin is low, it plays an important role in preventing or postponing the long-term complications of diabetes. Recently, we demonstrated that the zinc addition to insulin (1:5 insulin hexamer–zinc ion) during the preparation of the delivery system enhanced insulin stability during release.28 In the present study, we DOI 10.1002/jps
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investigated the thermosensitive polymeric delivery system for controlled insulin delivery in vitro and in vivo in diabetic rats to meet the basal insulin requirements. The overall structural stability, which comprises conformational and chemical stability, of the in vitro-released insulin was evaluated. Finally, in vitro and in vivo biocompatibility of thermosensitive systems were also studied.
EXPERIMENTAL Materials D,L-Lactide
was obtained from Toyo Kasei Kogyo Company Ltd. (Tokyo, Japan). Polyethylene glycol (PEG 1500) and stannous ethylhexanoate were purchased from Sigma Chemical Company (St. Louis, Missouri). Human recombinant insulin and microBCA protein assay kit were obtained from Celliance Corporation (Norcross, Georgia) and Pierce Biotechnology Inc. (Rockford, Illinois), respectively. Human insulin enzyme-linked immunosorbent assay (ELISA) kit was obtained from Linco Research Inc. (St. Charles, Missouri). Synthesis and Characterization of Triblock Copolymers
The PLA–PEG–PLA copolymers were synthesized using a method previously described by Singh et al.15 PEG block length of 1500 and D, L-lactide were used for synthesis. Under nitrogen, PEG was dried at 100◦ C in a three-necked flask for 30 min, followed by addition of D,L-lactide. After melting of the reactants, 0.03% (w/w) of stannous 2-ethylhexanoate was added, and the reaction was further carried out at 120◦ C for 12 h. Ice-cold water (5◦ C–8◦ C) was used to dissolve the crude polymer, and it was then reheated for purification. This step was repeated to improve the purity of the copolymer by removing any unreacted monomers. Finally, freeze-drying was performed to remove residual water. The synthesized copolymers were characterized by 1 H NMR spectrometry and Gel Permeation Chromatography (GPC) in order to determine their structure, molecular weight, molecular weight distribution, and purity. Preparation of Thermosensitive Delivery Systems of Insulin and In Vitro Release Aqueous solutions of the copolymer A (molecular weight 1496–1500–1496 Da) or copolymer B (molecular weight 1584–1500–1584 Da) were prepared by dissolving the copolymers in deionized water to obtain final concentrations of 30% and 40% (w/w). Insulin was dispersed in aqueous copolymer solution and the mixture was homogenized for 30 s at room temperature. Unless otherwise stated, zinc acetate was added to insulin delivery systems at a molar ratio of 1:5 insulin JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Table 1. Composition of Insulin Delivery Systems Based on Copolymer A and B Thermosensitive Triblock Copolymers and In Vitro Release Kinetics
Formulation
Copolymer
Copolymer Concentration (%, w/w)
Insulin Hexamer: Metal Ion
Insulin (%, w/v)
Zero Order (r2 )
First Order (r2 )
Higuchi (r2 )
A A A B B B
30 40 40 30 40 40
1:5 1:5 – 1:5 1:5 1:5
0.5 0.5 0.5 0.5 0.5 2.5
0.97 0.99 0.95 0.97 0.99 0.95
0.82 0.86 0.79 0.79 0.82 0.77
0.99 0.97 0.99 0.99 0.95 0.97
1 2 3 4 5 6
hexamer to zinc ion. The composition of the delivery systems is presented in Table 1. A delivery system (1 mL) was injected into a 20 mL polypropylene tube and incubated in a reciprocating water bath maintained at 37◦ C. After formation of the gel, 15 mL of prewarmed phosphate-buffered saline (PBS), pH 7.4 was added to the test tube as release medium. For the entire study duration, the test tube was incubated in a reciprocating water bath maintained at 37◦ C and 35 rpm. At different time points, 5 mL samples were withdrawn from the dissolution media. Fresh PBS was used to replace the removed volume. After centrifugation at 4229g for 30 min, samples were diluted with PBS. The released insulin was quantified using microBCA protein assay (Pierce Biotechnology Inc.). Stability Studies
Differential Scanning Calorimetry An ultrasensitive differential scanning calorimeter (VP-DSC, MicroCal, Northampton, Massachusetts) was used to study the conformational stability of the in vitro-released insulin. The sample preparation method described previously by Al-Tahami et al.28 has been used. Samples were centrifuged, filtered, and degassed under vacuum. Experiments were carried out using a scan rate of 1◦ C/min over a temperature range of 25◦ C–105◦ C. The thermodynamic parameters determined were the calorimetric enthalpy (H) and R midpoint transition temperature (Tm ). The Origin software (MicroCal) was used for data analysis.
Circular Dichroism Jasco J-815 circular dichroism (CD) spectrophotometer (Jasco, Tokyo, Japan) was used to evaluate the secondary structure of insulin. We have earlier described the details about the sample preparation method.28 Spectra of buffer and insulin were scanned from 200 to 300 nm. The spectra of the buffer were used as baseline and were subsequently subtracted from insulin sample spectra. MicroBCA assay (Pierce Biotechnology Inc.) was used to determine the concentration of insulin in order to calculate the molar ellipticity (θ). JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
Matrix-Assisted Laser Desorption/ Ionization–Time-of-Flight Mass Spectrometry A Bruker matrix-assisted laser desorption/ ionization–time-of-flight (MALDI–TOF) II mass spectrometer (Bruker Daltonics Inc., Billercia, Massachusetts) equipped with smartbeam in combination with a solid-state laser (200 Hz) was used to study the chemical stability of insulin. The details of the method can be found in an earlier publication.28 R FlexAnalysis software (Bruker Daltonics Inc., Billercia, MA) was used to analyze the data.
In Vitro and In Vivo Biocompatibility An MTT assay was used to determine in vitro biocompatibility of in situ gel depot in human embryonic kidney (HEK293) cells. The detail methods for in vitro and in vivo biocompatibility are described in the previous publication.28 Briefly, the polymers were extracted in PBS maintained at 37◦ C and 70◦ C for 10 days. As degradation rate of polymers is greater at higher temperature, the extract prepared at 70◦ C simulates the long-term effects of the delivery system. The cytotoxicity of the polymeric extracts, growth medium without polymer (negative control), and 2% dimethyl sulfoxide in growth medium (positive control) was tested in HEK293 cells. The in vivo biocompatibility of the delivery systems was determined by studying inflammatory changes in the rat skin. The formulation was administered to the rats subcutaneously into the ventral side of the neck. Skin sections from injection sites were surgically removed after euthanizing the rats at day 1, 30, and 90 time points. The skin samples were prepared for histological analysis and studied for the signs of inflammation, necrosis, or any changes in the morphology.
In Vivo Absorption and Bioactivity Studies In vivo absorption and bioactivity of insulin was tested by measuring blood glucose reduction in diabetic male Sprague–Dawley (SD) rats (200–224 g body weight). The protocol for animal studies was approved by Institutional Animal Care and Use Committee at North Dakota State University, Fargo, North Dakota. The animal experiments were carried out according to the guidelines and humane animal care standards DOI 10.1002/jps
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described in the National Institutes of Health Guide for the Care and Use of Experimental Animals. A single intraperitoneal dose of streptozotocin (STZ; 60 mg/ kg) in citrate buffer (pH 4.5) was administered to inR (Bayer Corporation, duce diabetes. Glucometer Elite Elkhart, Indiana) was used to estimate blood glucose level. R , Ovation PharPentobarbital sodium (Nembutal maceuticals Inc., Deerfield, IL) at a dose of 30 mg/kg body weight was injected intraperitoneally to anesthetize the rats. The temperature-sensitive delivery system was injected subcutaneously at doses of 90 insulin unit (U)/kg body weight at the neck region with the help of 25-gauge needle. An aqueous solution of insulin in PBS (2 U/kg body weight) was also studied. The control was polymeric solution without insulin. Blood was withdrawn at regular intervals from tail vein. Insulin serum concentration and blood glucose level were determined by ELISA and glucose oxidase method, respectively. The details of the method have been described in an earlier publication.28 Enzyme immunoassay (human insulin ELISA kit; Linco Research Inc.) was performed for insulin according to the manufacturer’s instructions.29 Briefly, the plates, samples, and insulin standards (2–200 :U/mL) were prepared as instructed. Insulin standards and samples were added to wells followed by detection antibody solution. Plates were covered and incubated with mixing at room temperature. Solutions were discarded and plates were washed thoroughly with horseradish peroxidase (HRP) wash buffer. Enzyme solution was then added to each well followed by incubation at room temperature. Plates were washed again with HRP buffer after decanting the enzyme solutions. Substrate solution (3,3 ,5,5 -tertramethylbenzidine) was added to wells, and plates were incubated until blue color was developed. The reaction was stopped using ELISA stop solution (0.3 M HCl). The blue color turned yellow after acidification and absorbance was read at 450 nm immediately. The dose-absorbance curve of insulin standards was fitted to a sigmoidal logistic equation and used to determine the concentration of insulin in samples. After completion of the experiments, rats were euthanized by intraperitoneal injection of pentobarbital (150 mg/kg body weight).
Data Analysis Analysis of variance (ANOVA) and Student’s t-test were used for computing the results. A p value of 0.05 was taken as significant. The serum insulin concentration versus time profile was used to determine the maximum serum concentration (Cmax ), time to Cmax (Tmax ), and area under the curve (AUC).30 DOI 10.1002/jps
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Figure 1. Effect of zinc addition on in vitro release of insulin from 40% (w/w) copolymer A (1496–1500–1496) (n = 4). Key: (䊏) zinc acetate and () no zinc salt added. [Key: ∗ , significant compared with 1:5 insulin hexamer to zinc ion at p < 0.05]
RESULTS Synthesis and Characterization of Copolymers Proton nuclear magnetic resonance (NMR) was used to determine the structure and molecular weight of the copolymers. The chemical shift (*) signals originating from different protons confirmed the structure of the synthesized copolymer. The spectrum was similar as reported in the literature by Jeong et al.31 The synthesized copolymers A and B showed narrow molecular weight distribution (polydispersity index 1.1). The molecular weights were determined as described by Jeong et al.31 using signal integration method in proton NMR, and was found to be 1496–1500–1496 and 1584–1500–1584 for copolymer A and B, respectively.
In Vitro Release of Insulin Table 1 lists the composition and release kinetics of insulin from delivery systems made of copolymer A and B. The effect of metal ion addition and polymer concentration on insulin release from copolymer A-based thermosensitive delivery systems are shown in Figures 1 and 2, respectively. Figure 1 demonstrates that the addition of zinc lowered the burst release following zero-order release kinetics (r2 ∼0.99) in comparison with formulation without zinc. The release of insulin was controlled for 2 months. Increasing the polymer concentration from 30% to 40% (w/w) reduced the burst release and improved release kinetics (Fig. 2 and Table 1). Similar results were observed when copolymer B (1584–1500–1584) was used (Fig. 3 and Table 1). Increasing chain length of PLA in copolymer B (1584–1500–1584) compared with copolymer A (1496–1500–1496) extended the release period. The insulin was released from delivery JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 2. Effect of polymer concentration on in vitro release of insulin from triblock copolymer A (1496–1500–1496) (n = 4). Key: copolymer concentration of (䊏) 30% and (•) 40% (w/w). [Key: ∗ , significant compared with 40% (w/w) copolymer A at p < 0.05]
systems containing 40% (w/w) copolymer B over 3 months with best fit for zero-order release kinetics (r2 ∼0.99). Increasing the insulin loading dose from 5 (0.5%, w/v) to 25 mg (2.5%, w/v) led to a small increase in burst release (Fig. 4). Stability Studies Differential scanning calorimetry studies showed that the insulin released after 1 month had lower H and Tm compared with fresh insulin (Fig. 5). Fresh and in vitro-released insulin showed Tm values of 82.49 ± 0.92 and 83.02 ± 0.13, and H values of 14.3 ± 0.7 and 10.1 ± 0.5 Cal/mol × 103 , respectively. CD spectra for fresh and insulin released in vitro after 1 and 2 months is represented in Figure 6. The magnitude of minima at 209 and 222 nm was decreased in the released insulin samples as com-
Figure 3. Effect of polymer concentration on in vitro release of insulin from triblock copolymer B (1584–1500–1584) (n = 4). Key: copolymer concentration of (䊏) 30% and (•) 40% (w/w). [Key: ∗ , significant compared with 40% (w/w) copolymer B at p < 0.05] JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
Figure 4. Effect of insulin loading on the in vitro release of insulin from triblock copolymer B (1584–1500–1584) (n = 4). Key: Insulin amount incorporated (䊏) 5 mg and (䉬) 25 mg. [Key: ∗ , significant compared with 5 mg insulin at p < 0.05]
pared with fresh insulin. However, the presence of the two minima confirms that the secondary structure of insulin was relatively conserved during delivery system preparation and release. The structural integrity of released insulin from thermosensitive delivery systems was evaluated using MALDI–TOF mass spectroscopy. Figure 7 shows the MALDI spectra of fresh insulin, 1 and 2 months released insulin from the delivery system containing 40% (w/w) copolymer B. MALDI–TOF (M+H)+ signal at 5808.6 Da indicated the conservation of primary structure of insulin in the released samples. No major degradation products were detected. There was no evidence of formation of acylation products (MW+ 72) or high-molecular weight transformation products. However, when the released insulin spectra were enlarged, a signal corresponding to a molecular weight of 5792.9 Da was observed (Figure 7b and 7c). As mentioned previously, this indicates a loss of a water molecule from the insulin, suggesting the formation of cyclic imide product that usually follows deamidation reactions.28 The insulin released after 2 months (Fig. 7c) showed a higher peak for the 5792 Da signal, implying that further deamidation took place. The 5792 Da signal relative intensity percentage (compared with 5808 Da signal, which corresponds to the native insulin) was equal to 13.7% and 37.8% in released insulin samples withdrawn after 1 and 2 months, respectively. This increase in deamidation products could be caused by the prolonged contact of insulin with the microenvironment inside the depot or the release medium. Two additional signals were observed in 2 months released samples corresponding to molecular weights of 5217.1 and 5070.2 Da (Fig. 7c). These two signals appear to be products of peptide-bond hydrolysis, which led to the formation of lower molecular weight peptides. DOI 10.1002/jps
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Figure 5. DSC thermograms of released insulin after 1 month from copolymer B thermosensitive system (2.5%, w/v insulin): Fresh insulin (solid line), and released insulin (dotted line), released insulin (without zinc) (dashed line).
Biocompatibility of Polymeric Delivery Systems
In Vitro Biocompatibility Figures 8 and 9 show the results of MTT cell viability assay for thermosensitive delivery system extracts prepared from copolymer A and copolymer B, respectively. No significant difference (p > 0.05) in the cell viability was observed between polymers extract groups and PBS group for all the dilutions and exposure time
Figure 6. CD spectra of released insulin from copolymer B thermosensitive system (2.5%, w/v insulin): Fresh insulin (solid line), 1 month released insulin (dashed line), and 2 months released insulin (dotted line). DOI 10.1002/jps
points. Interestingly, polymer extract showed greater cell survival compared with PBS group at certain dilutions (copolymer B; 1:8 dilution; 72 h exposure).
In Vivo Biocompatibility The inflammatory response of rat skin to temperature-sensitive delivery systems was evaluated at different time points after subcutaneous administration. Figures 10a and 11a show light microscopy images of skin sample taken from control rats where no formulation was injected. Figures 10b– 10d represent light microscopy images of skin tissues from where temperature-sensitive-based delivery systems (copolymer A) were administered. Similarly, Figures 11b–11d show light microscopy images of skin tissue injected with copolymer B containing delivery system. Both polymeric systems demonstrated inflammatory reactions and tissue healing process after injection. Following administration, inflammatory cells (mainly neutrophiles and lymphocytes) infiltrate to injection site. After a month, number of macrophages dropped, which are components of chronic tissue inflammation in response to injury.8 Skin samples from 3 months time point were comparable to control tissue. The connective and muscular tissues of the injection sites appeared normal and comparable to control skin tissue. We did not observe any muscle damage or necrotic damage throughout JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 7. MALDI–TOF mass spectroscopy of (a) fresh insulin (control), (b) 1 month and (c) 2 months from thermosensitive delivery system (40%, w/w copolymer B).
the studies, denoting the biocompatible nature of the delivery systems.
In Vivo Absorption and Bioactivity Figure 12 shows the serum insulin concentration and blood glucose level following subcutaneous administration of insulin (2 U/kg body weight) dissolved in PBS (pH 7.4) in diabetic SD rats. Serum insulin level increased rapidly reaching mean peak concentration (Cmax ) of 67.84 :U/mL at 2 h after administration and declined afterward to reach below 2 :U/mL (detection limit) after 12 h. Blood glucose levels decreased following insulin absorption, showing an acute and relatively short hypoglycemic effect. Blood glucose levels were restored to preadministration levels within 8 h. In vivo absorption and biological activity of insulin are presented in Figure 13. The control group composed of copolymer without insulin. Treatment groups were injected with insulin (90 U/kg body weight) dissolved in 40% (w/w) of either copolymer A or copolymer B. Zinc was added to both treatment groups as zinc acetate (1:5 ratio). Treatment groups showed continuous insulin release from thermosensitive delivery systems over a period of 3 months. The more hydrophobic copolymer B showed a lower release rate JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
for longer duration compared with copolymer A. Blood glucose levels in rats receiving either of the delivery systems were significantly lower (p < 0.05) than the diabetic rats in control group. The bioavailability of insulin in rats was calculated in terms of AUC values after correcting for dose (Table 2). When compared with subcutaneous insulin solution, copolymer B-based delivery system enhanced insulin bioavailability by approximately 11-fold.
DISCUSSION In vitro release studies showed that increasing the chain length of PLA in copolymer resulted in lowering the burst release of insulin. Because of the more hydrophobicity of lactic acid, the copolymer absorbed less water and degraded slowly.17,32 Similar results have been reported when PLGA–PEG–PLGA thermosensitive delivery systems with varying LA/GA ratios were used for the delivery of 5-fluorouracil.33 Furthermore, it was shown that increasing LA/GA ratio decreased burst release and extended the release period of bovine serum albumin from PLGA microspheres.34,35 DOI 10.1002/jps
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Figure 8. In vitro biocompatibility of thermosensitive copolymer A (1496–1500–1496) extracts (a–c: Extracts prepared at 37◦ C and d–f: Extracts prepared at 70◦ C) (n = 8). [PBS, phosphatebuffered saline; PE,polymer extract; DMSO, dimethyl Sulfoxide (2%)].
The release kinetics for delivery systems containing 30% (w/w) copolymer A showed best fit for Higuchi model (r2 = 0.97), indicating a diffusion dominant initial release phase. The release showed a best fit for zero-order release kinetics (r2 = 0.99) as the concentration of copolymer increased to 40% (w/w). This may be attributed to the reduced diffusion due to the less porous depot at higher copolymer concentration. We speculate that the addition of metal ions reduced insulin association in PEG domain by lowering insulin solubility. Thus, less insulin was available for the initial burst release and most of the incorporated DOI 10.1002/jps
insulin was associated with PLA and was released during the degradation phase. Thermosensitive delivery systems containing 40% (w/w) of copolymer A and B, to which zinc ions were added at the molar ratio 1:5 insulin hexamer to zinc ion, showed the best fit for zero-order (r2 = 0.99) release kinetics of insulin over periods of 2 and 3 months, respectively. An important aspect in the use of controlled-release systems as proteins delivery devices is the ability of loading different doses of proteins. This is mainly to count for intraindividual and interindividual drug requirement variations. Increasing insulin loading in JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
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Figure 9. In vitro biocompatibility of thermosensitive copolymer B (1584–1500–1584) extracts (a–c: Extracts prepared at 37◦ C and d–f: Extracts prepared at 70◦ C) (n = 8). [PBS, phosphatebuffered saline; PE, polymer extract; DMSO, dimethyl sulfoxide (2%)].
delivery systems containing copolymer B from 5 to 25 mg led to an increase in insulin burst release and best fit for release kinetics was shifted from zero to pseudo-zero-order model. Yet, the release profile exhibited a high-zero-order correlation (r2 = 0.95). The DSC results confirmed that the addition of divalent metal ion stabilizes insulin, indicated by the higher denaturation enthalpy. Zinc has a strong inhibitory effect on fibril formation/growth due to the JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
neutralization of the charges at the core of the insulin hexameric assembly, leading to the stabilization of these hexamers. This leads to an increase in the hydrophobic interactions within insulin hexamer, thus, increasing its stability. Numerous studies have shown that insulin denaturation is highly dependent on insulin being in the monomeric form in solutions.36 It has also been reported that complexation with zinc was able to stabilize insulin during release from DOI 10.1002/jps
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Figure 10. Light micrographs of skin taken from injection site (a) Control, (b) after 1 day, (c) 1 month, and (d) 3 months after subcutaneous administration of thermosensitive delivery system (copolymer A). Acute inflammatory cells infiltration (N, neutrophiles and L, lymphocytes) between muscle fibers and hemorrhage in injection site is apparent. [Magnification: 20×].
Figure 11. Light micrographs of skin taken from injection site (a) Control, (b) after 1 day, (c) 1 month, and (d) 3 months after subcutaneous administration of thermosensitive delivery system (copolymer B). Acute inflammatory cells infiltration (N, neutrophiles and L, lymphocytes) between muscle fibers and hemorrhage in injection site is apparent. [Magnification: 20×].
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Figure 12. Serum insulin concentration (a) and blood glucose level (b) following subcutaneous administration of insulin solution (n = 6).
microspheres and also prevented nonspecific adsorption to hydrophobic polymer surfaces.37 The CD spectrum of insulin had been reported to exhibit two minima at 209 and 223 nm,37 indicating a dominant "-helical structure (∼38%) with a small part of betasheets (∼17%). The two minima were greatly attenuated for preparations containing no zinc ion after 28 days of incubation ("-helix ∼18% and $-sheets ∼4%) in comparison with preparations containing 1:5 insulin hexamer to zinc ion ratio ("-helix ∼36% and $-sheets ∼14%). This attenuation denotes the aggregation process and the loss of bioactive form of insulin from the solution. The conformational stability and secondary structure of released insulin was partially reduced, which may be due to prolonged exposure of released insulin in agitated media at 37◦ C temperature.36,38–40 The correlation between protein unfolding and chemical stability is not straightforward as many facJOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
tors affect protein chemical stability. The chemical degradation, which takes place in protein domains that are protected in native state, occurs at higher rates when the protein is denatured.41 It has been shown that there is a high correlation between DSC results and insulin B3 deamidation products.42 However, the DSC data failed to predict the rate of covalent insulin dimer (CID) formation. In addition, because DSC measures thermal stability, it cannot be used to predict chemical stability at low temperatures. Insulin undergoes many chemical modification reactions depending on the surrounding environment and that can affect its biological activity. Biological activity of native insulin and its deamidation products is similar, whereas the products formed due to hydrolysis (A8–A9) and CIDs exhibit only 2% and 15% relative potency, respectively.43 The formation of CID and deamidation products of insulin has been DOI 10.1002/jps
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Figure 13. Serum insulin concentration (a) and blood glucose level (b) following subcutaneous administration of thermosensitive polymeric delivery systems in SD rats (n = 6). Key: (䊏) blank thermosensitive delivery systems (control), (•) thermosensitive delivery systems composed of 40% (w/w) copolymer A (1496–1500–1496), and () thermosensitive delivery systems composed of 40% (w/w) copolymer B (1584–1500–1584).
reported to be increased by temperature and shaking.40 MALDI results from 1 month-released samples from the delivery systems did not show peaks for acylation, CID, or hydrolysis products and the spectra were comparable to fresh insulin spectrum. A small signal denoting the formation of cyclic imide Table 2.
by the loss of a water molecule was observed. Iso–Asp insulin is formed due to the deamidation reaction, which follows cyclic imide formation.44 However; insulin from the 2 months-released sample exhibited an increase in the cyclic imide signal, which denotes an increase in deamidation reactions. Additional
In Vivo Pharmacokinetic Parameters of Delivered Insulin in Rats
Group SG T1 T2
Cmax (:U/mL)
Tmax (h)
AUC [:U/(h mL)]
Increase in AUC (Folds Compared with SC)
67.84 ± 7.38 – –
1.00 – –
5493.83 ± 1263.72 41709.90 ± 8116.30a 59924.41 ± 11279.68a
– 7.59 10.91
SG, insulin solution group; T1, thermosensitive delivery systems—copolymer A (1496–1500–1496); T2, thermosensitive delivery systems—copolymer B (1584–1500–1584). a Significantly (p < 0.05) higher than solution group.
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hydrolysis degradation products signals (5217.1 and 5070.2 Da), which correspond to B24–B25 and B25–B26 split products, respectively, were observed. As discussed previously, this could be due to the prolonged presence of insulin in the shaking release media at 37◦ C. As insulin released from the delivery system is absorbed instantly, we do not anticipate this to happen in vivo. It was extremely crucial to investigate the biocompatibility of the delivery systems used in this study. Neutralized extracts of thermosensitive delivery systems did not exhibit any inhibitory effect on cell viability. On the contrary, at some time points, cell viability was greater in polymer extract groups. It has been suggested that this could be due to the increase in lactic acid, which is a metabolite in the citric acid cycle.45 In vitro biocompatibility is a very good indication for predicting in vivo biocompatibility. In the beginning, there was sign of inflammation reaction due to injury caused by injection.9 After 3 months of injection of the delivery system, all signs of inflammatory responses disappeared and skin histology of injected site was comparable to control skin samples. The findings of in vivo biocompatibility were in agreement with the in vitro biocompatibility and International Organization for Standardization (ISO) regulations, supporting the biocompatible nature of the thermosensitive copolymers as a delivery system. The success of insulin continuous delivery at basal level depends on the delivery system to deliver insulin in vivo at controlled rate in biologically active form. In this study, diabetic model was created by injecting STZ into male SD rats intraperitoneally. STZ is a glucomimetic molecule and employs glucose transporter GLUT2 for its transport to pancreatic $ cells where the mechanism of the cell death includes the nitrosourea-induced DNA alkylation.46 There was an increase in pharmacological bioavailability of insulin from the copolymer-based delivery systems in comparison with subcutaneous insulin solution in PBS. It has been shown that insulin absorption rates from subcutaneous tissue are inversely correlated to insulin concentration.47 The slower release of insulin from the delivery system led to further absorption of insulin from the releasing site, thereby less likely to degrade by proteases present in the skin.48 The polymeric delivery system has also offered additional stabilizing effect on insulin by protecting it from enzymatic degradation at the injection site. It is apparent from the data that the pharmacological activity of insulin released was preserved in vivo. The duration of in vivo release was shorter in comparison with in vitro release due to the presence of enzymes in vivo, leading to faster polymer degradation of polymer.13–14,49,50 But, in vivo release seems to be consistent with the blood glucose profile and JOURNAL OF PHARMACEUTICAL SCIENCES, VOL. 100, NO. 11, NOVEMBER 2011
the insulin release pattern observed in the in vitro release.
CONCLUSION The present work shows that the temperaturesensitive triblock copolymers can be used for the controlled delivery of insulin at basal level in a conformationally and chemically stable and bioactive form. Insulin release can be altered by changing polymer composition, polymer concentration, and addition of metal ions. Release pattern reliability on formulation factors is advantageous as it allows for delivery of different compounds with different requirements. The stabilizing effect offered by the polymeric implant and metal ion addition was also demonstrated.
ACKNOWLEDGMENTS We acknowledge the financial support from NIH grant #HD056053. We also acknowledge the Fulbright Program for scholarship to Khaled Al-Tahami. The assistance provided by Buddhadev Layek in the preparation of the manuscript is greatly appreciated.
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