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Thermogelling Biodegradable Copolymer Aqueous Solutions for Injectable Protein Delivery and Tissue Engineering Byeongmoon Jeong,*,†,‡ Kyeonghee M. Lee,† Anna Gutowska,*,† and Yuehuei H. An§ Pacific Northwest National Laboratory, 902 Battelle Boulevard, P.O. Box 999, K2-44, Richland, Washington 99352, Department of Chemistry, Ewha Womans University, Seoul 120-750, Korea, and Department of Orthopedic Surgery at the Medical University of South Carolina, Charleston, South Carolina 29425 Received March 11, 2002; Revised Manuscript Received April 21, 2002
This paper reports on the thermogelling, biodegradable polymer formulations based on poly(DL-lactic acidco-glycolic acid)/(poly(ethylene glycol) graft copolymers for in vivo biomedical applications using animal models. The description includes diabetic control by sustained insulin delivery and cartilage repair by chondrocyte cell delivery. With one injection of the poly(DL-lactic acid-co-glycolic acid)/(poly(ethylene glycol) graft copolymers insulin formulation, the blood glucose level could be controlled from 5 to 16 days in diabetic rats by varying the polymer composition. The cartilage defect was notably repaired using chondrocyte suspension in the thermogelling PLGA-g-PEG compared with a control. This report shows that thermogelling biodegradable PLGA/PEG graft copolymer system can be a promising platform for protein and cell-based therapy. Introduction Due to potential biomedical applications, such as drug delivery and tissue engineering, the sol-gel reversible hydrogel has attracted recent attention.1-4 Typical examples are copolymers of N-isopropylacrylamide (NIPAAm),5 ethylene oxide/NIPAAm block copolymers,6 Poloxamers,7 ethylene oxide/butylene oxide block copolymers,8 Poloxamergrafted with poly(acrylic acid)s,9 chitosan/glycerol phosphate,10 poly(ethylene glycol) (PEG)/poly(lactic-co-glycolic acid) (PLGA) triblock11 or graft copolymers,12,13 and poly(fumaric acid-co-propylene glycol)/PEG triblock copolymers.14 Compared with classical microsphere and liposome technology, aqueous sol-gel reversible hydrogel systems do not need organic solvents during the fabrication procedure. Drug or cell loading can be controlled as a suspension or solution. Syringe injection can make the sustained delivery depot of pharmaceutical agents and cells simple. During the next decade, more and more biopharmaceutical treatments, including protein/peptide drug and cell-based therapy will be available due to advances in biotechnology and proteomics research.15 However, the current dosage method of protein/peptide drugs requires daily or twice per day injections due to their fast clearance by enzymatic degradation in the body and thus suffer from consequent patient compliance problems. However, cell-based therapy research, such as artificial pancreatic cells for diabetic treatment, begins to show very promising results.16 * To whom correspondence should be addressed: Jeong, tel 82-2-32772334, fax 82-2-3277-2384, e-mail
[email protected]; Gutowska, tel 509375-4443, fax 509-375-2186, e-mail
[email protected]. † Pacific Northwest National Laboratory. ‡ Ewha Womans University. § Department of Orthopedic Surgery at the Medical University of South Carolina.
Figure 1. Structure of PEG-g-PLGA and PLGA-g-PEG copolymers.
In this paper, we are reporting in vivo feasibility of diabetic control and tissue engineering by using thermogelling PEGg-PLGA and PLGA-g-PEG aqueous solutions. Experimental Section Polymer Formulation Preparation. The PEG-g-PLGA (Mn ∼ 6000, polydispersity index (PDI) ∼ 1.5, EG/LA/GA ratio ∼ 2.98/2.35/1.00, number of grafts ∼3) and PLGA-gPEG (Mn ∼ 6000, PDI ∼ 1.7, EG/LA/GA ratio ∼ 2.71/3.28/ 1.00, number of grafts ∼3) were dissolved in a phosphate buffer at 4 °C for 12 h. The preparation and sol-to-gel transition behavior of these polymers were reported elsewhere.12,13,17 Aqueous solutions of both polymers undergo sol-to-gel transition and show abrupt changes in modulus with increasing temperature. The maximum modulus of the gel is about several hundred dyn/cm2. The in situ gelation at 37 °C was also confirmed by a subcutaneous injection of the polymer solution in rats.12,17 The polymer solution (25 wt %) was filtered through a 0.2 µm nylon filter. Insulin Depot. The insulin from a bovine pancreas was purchased from Sigma (28.3 USP Unit/mg, catalog no. I5500), and used as received. Diabetes in SD rats was
10.1021/bm025536m CCC: $22.00 © 2002 American Chemical Society Published on Web 05/23/2002
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Figure 2. In vivo efficacy in diabetic rats of 50/50 PEG-g-PLGA/ PLGA-g-PEG (top) and PLGA-g-PEG (bottom) depot systems. The insulin loading was 35.54 mg/kg for both systems. Blood glucose level of rats was investigated after injection of each formulation (0.5 mL). The zero day is the one when insulin formulations were injected. The line indicates a trend line of blood glucose level. The duration of efficacy was 5 days (FI) and 16 days (FII) by injection.
induced by streptozotoxin (STZ) with an intraperitoneal injection of 55 mg/kg. Rats were considered diabetic, if their blood glucose level exceeded 200 mg/dL. After 10 days of observation, the rats with blood glucose levels between 400
Figure 3. Tissue response around implant (FII).
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and 600 mg/dL were selected. Noninduced rats were used as sham control. On terminal day, selected organs, such as skin around the injection site, spleen, and kidneys, were also collected and saved in formalin for possible evaluation. The insulin was mixed with PLGA-g-PEG and/or PEG-g-PLGA copolymer aqueous solution and about 0.5 mL of each formulation was subcutaneously injected into a rat, using a 20-gauge syringe. The injected amount of insulin was fixed at 35.54 mg/kg for a rat.18 Samples were collected each designated time interval, such as 1-h, 3-h, 6 h, and 1 day, then every day for the first week, and every other day for the second, third, and fourth weeks. Two rats were used for the PEG-g-PLGA/PLGA-g-PEG (50/50 wt ratio) formulation (FI). Three rats were used for the PLGA-g-PEG formulation (FII). Blood glucose level (BGL), body weight, or water consumption were recorded. Cartilage Repair. Articular cartilage defects (diameter ∼ 4 mm, depth ∼ 5 mm) were created in the right femoral trochlea of rabbits (male NZW rabbits, 3.5 ( 0.25 kg). Chondrocyte cells were harvested from rabbit scapula and expanded in a 3-D thermosensitive N-isopropyl acryl amide/ acrylic acid copolymers culture before implantation.5 The defects were implanted with the autogenic chondrocyte cells suspended in the PLGA-g-PEG aqueous polymer solutions. Chondrocyte cells suspended in the poly(N-isopropyl acryl amide-co-acrylic acid)/hydroxyapatite collagen sponge were used as a control. Two rabbits were used for each protocol. Because autologous chondrocytes were used, each rabbit was operated on twice, once for cell harvesting and the second for joint defect transplantation. For the first surgery, under general anesthesia and sterile conditions (skin shaved and cleaned with 7.5% Povidon-Iodine and 70% isopropyl
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Figure 4. Cartilage repair: (a) overview of control sample using poly(N-isopropyl acrylamide-co-acrylic acid) (left) and PLGA-g-PEG (right); (b) histology of control sample using poly(N-isopropyl acrylamide-co-acrylic acid) (left) and PLGA-g-PEG (right).
alcohol), the lower portion of the right scapular bone was approached through an upper dorsal incision. The lower rim of the scapular (about 1 × 10 × 35 mm, composed mainly fibrocartilage) was cut off and placed in a Petri dish containing phosphate-buffered saline (PBS). After surgery, the wound was closed in layers and unrestricted weight bearing was allowed. The cartilage was cleaned of soft tissues and chondrocytes were isolated using enzymatic digestion described by Grande et al.19 Further propagation of the chondrocytes was made for later transplantation in 2 weeks. For the second surgery, also under general anesthesia and sterile conditions, the distal femoral joint surface was reached through a 3.0 cm longitudinal anteromedial incision. A 4-mm diameter hole (5 mm in depth) was created in the femoral trochlea. Then the grafts with chondrocytes (4 × 106 cells per graft) were filled into the defects. After surgery, the wound was closed and unrestricted weight bearing was allowed. The animals were sacrificed 12 weeks after surgery. A histology study will be conducted for evaluating the repair of the defects. Results and Discussion The aqueous solutions of graft copolymers consisting of PEG and PLGA, PEG-PLGA and PLGA-g-PEG, are low
viscous sols at low temperature but become gels at physiological temperature (37 °C).12,17 The PEG-g-PLGA has a hydrophilic backbone while PLGA-g-PEG has a hydrophobic backbone. Figure 1 shows the structure of the polymers. The zigzag line and the straight bold line indicate hydrophilic PEG and hydrophobic PLGA, respectively. Due to the surfactant nature of these polymers, PEG-g-PLGA and PLGA-g-PEG form micelles in water. In these micelles, the hydrophilic PEG forms flexible shells while the hydrophobic PLGA forms the micelle cores. Such a core-shell structure was confirmed by 13C NMR, cryogenic temperature transmission electron microscopy (cryo-TEM), and hydrophobic dye solubilization.12-13,17 Similar to the PEG-PLGA-PEG system,11 the change in micellar conformation is thought responsible for the gelation of these graft copolymers as discussed in above references. It is interesting to see that the PEG-g-PLGA gel shows faster degradation, whereas the PLGA-g-PEG gel shows slower degradation. In vivo data showed that the PEG-g-PLGA depot disappeared within a week while the PLGA-g-PEG depot persisted for about 3 months.13 By mixing two polymers of PEG-g-PLGA and PLGA-g-PEG, the duration of depot can be controlled 1 week to 3 months. The PEG-g-PLGA has a flexible PEG backbone and short chain grafts of PLGA. It forms a mechanically
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weak gel. And, the degradation products of the inherently short chain PLGA in PEG-g-PLGA are easily washed away by body fluids around the depot compared with long chain PLGA in PLGA-g-PEG. The insulin depot formed by subcutaneous injection of graft copolymer-insulin aqueous solutions shows a different pharmacodynamic profile, depending on the polymer composition (Figure 2). The diabetic condition was induced over a period of 10 days by STZ treatment, and the blood glucose level was higher than 400 mg/dL. Day zero data corresponds to injection of the insulin formulation. Upon subcutaneous injection of the formulation, the blood glucose level dropped to less than 200 mg/dL in 1 h for all formulations. However, the insulin depots prepared from 50/50 mixture of PEG-gPLGA and PLGA-g-PEG (FI) show a hypoglycemic efficacy over 5 days, while those of PLGA-g-PEG (FII) show 16 days efficacy by single injection of the formulation. One injection of the formulation every 16 days, instead of daily injections, would improve patient compliance. FI consists of fast degrading PEG-g-PLGA and slow degrading PLGA-g-PEG. The reason of short duration of efficacy for FI is caused by faster erosion and because the gel is mechanically looser than the gel made of FII. The insulin dose of 35.54 mg/kg is based on the microsphere system that shows hypoglycemic efficacy over 10 days by a single-layer matrix system.18 A double layer matrix system decreased the initial burst of insulin and showed the longer duration of hypoglycemic efficacy. Compared with the microsphere system, our system is much easier to process, due to no use of organic solvents and a simple sterilization method by syringe filtration. Blood urea nitrogen (BUN) concentrations, measured for kidney toxicity, were within the normal range, indicating kidney toxicity by STZ was not apparent. The rat body weight steadily increased from 330-370 to 360-430 g during the 23-day treatment; the weight of diabetic rats treated with insulin formulations was not significantly different from that of normal (nondiabetic) rats. The kidney, spleen, and pancreas were observed to be normal. Around the implant site are loose connective tissues of collagen with numerous new capillaries and arterioles and mast cells indicating a very minimal chronic inflammation in 1 month of implantation (Figure 3). To investigate the tissue engineering application using a biodegradable injectable system, a part (depth ∼5 mm, diameter ∼4 mm) of the rabbit cartilage tissue was removed. Chondrocyte cell suspension in the thermogelling PLGA-gPEG solution was injected into the defected site. The macroimages (Figure 4a) show that rabbits in the control experiment have a persistent defect, whereas the one implanted with chondrocyte/PLGA-g-PEG mixture shows almost complete healing. The histological examination (Figure 4b) correlated well with that visualized on gross inspection. The defects of both rabbits in control had failed to incorporate enough new cartilage growth, whereas the PLGA-g-PEG system appeared to fill the defect appropriately. The biodegradable system may be slowly replaced by in-growing tissue while the nonbiodegradable control system may hamper the growth of the new cartilage by forming a
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physical barrier against the new cartilage, resulting in incomplete growth of cartilage tissue. Although no attempt was made for quantitative analysis such as the O’Driscoll score for cartilage repair, the O’Driscoll score is known to be consistent with those of the histological and gross images.20 To conclude, the in vivo biomedical applications of the biodegradable thermogelling PEG-g-PLGA/PLGA-g-PEG aqueous solutions were investigated using animal models. As a protein delivery system, the efficacy could be controlled for 5-16 days by varying the composition of polymers. The cartilage defect was notably repaired using thermogelling PLGA-g-PEG compared with a control. The PLGA-g-PEG and PEG-g-PLGA systems have advantages such as (1) no organic solvent is used during fabrication, (2) no surgical procedure is needed to make depot, (3) easy sterilization is possible by syringe filtration, (4) the system is biodegradable, (5) drug or cell loading can be solution or suspension, (6) duration or profile can be controlled by change in composition. This report confirmed the PLGA/PEG graft copolymer system can be a promising platform for biopharmaceutical and cell-based therapy. Acknowledgment. This work was supported by Battelle/ Pacific Northwest National Laboratory Independent Research & Development and Department of Energy (DOE) funds. References and Notes (1) Gutowska, A.; Jeong, B.; Jasionowski, M. Anat. Rec. 2001, 263, 342349. (2) Stile, R. A.; Burghardt, W. R.; Healy, K. E. Macromolecules 1999, 32, 7370-7379. (3) Jeong, B.; Kim, S. W.; Bae, Y. H. AdV. Drug DeliVery ReV. 2002, 54, 37-51. (4) Jeong, B.; Gutowska, A. Trends Biotechnol., in press. (5) Ahn, Y. H.; Mironov, V. A.; Gutowska, A. US patent US6103528, 2001. (6) Topp, M. D. C.; Dijkstra, P. J.; Talsma, H.; Feijen, J. Macromolecules 1997, 30, 8518-8520. (7) Malstom, M.; Lindman, B. Macromolecules 1992, 25, 5446-5450. (8) Luo, Y. Z.; Nicholas, C. V.; Attwod, D.; Collett, J. H.; Price, C.; Booth, C. Colloid Polym. Sci. 1992, 270, 1094-1105. (9) Bromberg, L. J. Phys. Chem. B 1998, 102, 1956-1963. (10) Chenite, A.; Chaput, C.; Wang, D.; Combes, C.; Buschmann, M. D.; Hoemann, C. D.; Leroux J. C.; Atkinson, B. L.; Binette, F.; Selmani, A. Biomaterials 2001, 21, 2155-2161. (11) Jeong, B.; Bae, Y. H.; Kim, S. W. Macromolecules 1999, 32, 70647069. (12) Jeong, B.; Kibbey, M. R.; Birnbaum, J. C.; Won, Y. Y.; Gutowska, A. Macromolecules 2000, 33, 8317-8322. (13) Jeong, B.; Wang, L. Q.; Gutowska, A. Chem. Commun. 2001, 15161517. (14) Behravesh, E.; Shung, A. K.; Jo, S.; Mikos, A. G. Biomacromolecules 2002, 3, 153-158. (15) Chem. Eng. News 1999, Aug. 30, 29-32; 2000, Sept. 18, 49-65. (16) (a)Shapiro, J.; Lakey, J.; Ryan, E. New Engl. J. Med. 2000, 27, 230240. (b) Vernon, B.; Kim, S. W.; Bae, Y. H. J. Biomater. Sci., Polym. Ed. 1999, 10, 183-198. (17) Chung, Y. M.; Simmons, K. L.; Gutowska, A.; Jeong, B. Biomacromolecules 2002, 3, 0000-0000. (18) Yamakawa, I.; Kawahara, M.; Watanabe, S.; Miyake, Y. J. Pharm. Sci. 1990, 79 (6), 505-509. (19) Grande, D. A.; Pitman, M. I.; Peterson, L.; Menche, D.; Klein, M. J. Orthop. Res. 1989, 7 (2), 208-218. (20) O’Driscoll, S. W.; Keeley, F. W.; Saletr, R. B. J. Bone Jt. Surg. Am. Vol. 1988, 70, 595-606.
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