Acta Biomaterialia 10 (2014) 1177–1186
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Design of a composite biomaterial system for tissue engineering applications B. Jiang a,b, B. Akar a,b, T.M. Waller a, J.C. Larson a,b, A.A. Appel a,b, E.M. Brey a,b,⇑ a b
Department of Biomedical Engineering, Illinois Institute of Technology, Chicago, IL, USA Research Service, Edward Hines, Jr. V.A. Hospital, Hines, IL, USA
a r t i c l e
i n f o
Article history: Received 29 July 2013 Received in revised form 10 November 2013 Accepted 29 November 2013 Available online 7 December 2013 Keywords: Neovascularization Tissue engineering Fibroblast growth factor Platelet-derived growth factor Poly(lactic-co-glycolic acid)
a b s t r a c t Biomaterials that regulate vascularized tissue formation have the potential to contribute to new methods of tissue replacement and reconstruction. The goal of this study was to develop a porous, degradable tissue engineering scaffold that could deliver multiple growth factors and regulate vessel assembly within the porous structure of the material. Porous hydrogels of poly(ethylene glycol)-co-(L-lactic acid) (PEG– PLLA) were prepared via salt leaching. The degradation time of the hydrogels could be controlled between 1 and 7 weeks, based on hydrogel composition. Fibrin was incorporated into the interconnected pores of the hydrogels to promote neovascularization and as a reservoir for rapid (30 days) growth factor delivery. Fibroblast growth factor-1 (FGF-1) and platelet-derived growth factor-BB (PDGF-BB) were delivered from the system owing to their roles in the promotion of angiogenesis and vascular stabilization, respectively. Hydrogels tested in vivo with a subcutaneous implantation model were selected based on the results from in vitro degradation and growth factor release kinetics. Dual growth factor delivery promoted significantly more tissue ingrowth in the scaffold compared with blank or single growth factor delivery. The sequential delivery of FGF-1 following PDGF-BB promoted more persistent and mature blood vessels. In conclusion, a biomaterials system was developed to provide structural support for tissue regeneration, as well as delivery of growth factors that stimulate neovascularization within the structure prior to complete degradation. Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Biomaterials are often used as scaffolds to provide mechanical and structural support in tissue engineering and organ regeneration applications [1,2]. Depending on the type of tissue to be engineered (e.g. bone, cartilage or soft tissue), biomaterials are fabricated into tissue engineering scaffolds to meet specific tissue requirements in regard to mechanical properties and degradation rates. Generally speaking, a three-dimensional scaffold with controlled degradation rate, high porosity and appropriate pore size is probably required for generating a proper engineered tissue for use in reconstruction [3]. In plastic and reconstructive surgery, coordinating the processes of tissue regeneration and biomaterial degradation is essential for controlling the geometry of the resultant tissue to match preimplantation design [4]. Tissue regeneration involves the regulation of a series of signaling molecules, such as cytokines, growth factors and hormones. Regeneration within a biomaterial may require control over the ⇑ Corresponding author at: 3255 South Dearborn Street, WH 314, Chicago, IL 60616-3793, USA. Tel.: +1 312 567 5098; fax: +1 312 567 5707. E-mail address: [email protected] (E.M. Brey).
temporal delivery of some of these molecules in order to re-create this process. For example, neovascularization occurs via a complex temporal response of cells to a number of growth factors. Angiogenic factors, such as those from the vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF) families, are involved in the initiation of angiogenesis, while vessel maturation factors, including platelet-derived growth factor (PDGF)-BB and angiopoietin-1, are involved in the later stage of vessel stabilization [5]. Release of factors with distinct kinetics is likely to improve vessel assembly, with initial administration of angiogenic factor(s) followed by vessel maturation factor(s). The sequential delivery of FGF-2 (also known as bFGF) [6] or VEGF [7] followed by PDGF-BB has been shown to increase vessel density and stability over single growth factor delivery or simultaneous dual delivery when the drug delivery system is implanted subcutaneously [8]. Previously, it was found that fibrin–poly(ethylene glycol) (PEG) composite hydrogels stimulated vascularized tissue regeneration after implantation [9]. In this system, the PEG-based porous hydrogels provided a biologically inert material with a controlled three-dimensional (3-D) structure, while fibrin served as a bioactive material that stimulated vessel invasion into the pores. However, the hydrogels exhibit little degradation under physiological conditions. In the study presented
1742-7061/$ - see front matter Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.actbio.2013.11.029
1178
B. Jiang et al. / Acta Biomaterialia 10 (2014) 1177–1186
here, monomer units of poly(L-lactic acid) (PLLA) were introduced into the PEG structure to allow controlled degradation via hydrolysis to decouple degradation from tissue invasion [10]. Two growth factors (FGF-1 and PDGF-BB) involved in neovascularization were incorporated into the hydrogels in order to influence the density, persistence and maturity of the vessels formed. Previous studies evaluating the effect of sequential growth factor delivery on vascularization were focused mostly on the response in tissues surrounding the implants or following complete biomaterial degradation [6–8]. However, biomaterial scaffolds are often designed to precisely the structure and volumetric features of a given tissue defect. Engineering tissues of a specific 3-D shape requires that vascularized tissue invasion occurs simultaneously with implant degradation [4]. In fact, vascularized tissue must grow into the scaffold prior to degradation in order to maintain the preformed structure of the material. In the present study, vascular structure within the scaffold structure prior to degradation was evaluated. 2. Materials and methods 2.1. Materials PLGA 85:15 (Mw 50,000–75,000, acid terminated), poly(vinyl alcohol) (PVA, Mw 13,000–23,000), bovine serum albumin (BSA), sucrose (>99.5%), magnesium hydroxide (95%), PEG (Mw 3350), 3,6-dimethyl-1,4-dioxane-2,5-dione, Tin(II) 2-ethylhexanoate (95%), fibrinogen from human plasma (Fg, 50–70% protein), thrombin from human plasma (Tb), 2-hydroxy-2-methylpropiophenone (Irgacure 1173), acryloyl chloride (98%) and thriethylamine (TEA, 99.5%) were obtained from Sigma-Aldrich (St. Louis, MO). Recombinant rat PDGF-BB with and without carrier was obtained from R&D Systems (Minneapolis, MN). Recombinant rat FGF-1 was obtained from PeproTech (Rocky Hill, NJ). Iodine 125 was obtained from Perkin Elmer Inc. (Waltham, MA). Dichloromethane (DCM, 99.9%), sodium chloride (99.5%), magnesium sulfate anhydrous (97%), diethyl ether (anhydrous), ethyl alcohol (95%), PEG (Mw 8000), Shandon xylene substitute and phosphate buffered saline (PBS) solution were obtained from Fisher Scientific (Hampton, NH). Endothelial cell growth medium (EGM) and endothelial cell basal medium (EBM) were obtained from Lonza Inc. (Allendale, NJ). 2.2. Hydrogel synthesis 2.2.1. PEG–PLLA–DA and PEG–DA synthesis The procedure for synthesis of PEG–PLLA–DA has been described in detail elsewhere [10]. Briefly, 20 g of PEG (Mw 3350) was mixed with 4.24 g of 3,6-dimethyl-1,4-dioxane-2,5-dione (L-lactide) in the presence of 80 ll of stannous octoate. The mixture was allowed to react at 140 °C for 4 h under anhydrous conditions. After purification and lyophilization, the resulting PEG–PLLA product was reacted with acryloyl chloride (4 molar ratio) and TEA (2 molar ratio) to acrylate the polymer. The PEG–PLLA–DA was purified, lyophilized and stored at 20 °C in the dark prior to use. PEG–DA was synthesized by acrylating PEG (Mw 3350), using the same method. 2.2.2. Porous hydrogel fabrication A salt leaching technique was adapted for generation of porous PEG–PLLA–DA hydrogels [11]. Briefly, PEG–PLLA–DA alone or in combination with PEG–DA was dissolved in ethanol, with 2-hydroxy-2-methylpropiophenone (0.5% v/v) added as a photo-initiator. Sodium chloride was ground with a pestle and mortar and sieved to select crystals of a defined size range. Three hundred milligrams
of salt crystals were added per 200 ll of precursor solution and polymerized under UV light (k = 365 nm) for 5 min. The hydrogels were then incubated in 50 ml of deionized (DI) water overnight to leach out the salt crystals and ethanol. 2.2.3. PLGA microspheres The preparation of PLGA microspheres for protein delivery was performed using a water in oil in water emulsion as described previously [12]. Briefly, PLGA 85:15 (250 mg ml 1) was dissolved in DCM (oil phase) with Mg(OH)2 (7.5 mg ml 1) added as an anti-acid agent. The oil phase mixture (1 ml) was emulsified with PBS (0.2 ml) containing protein (5 lg ml 1 of either PDGF-BB or FGF1) and protein stabilizers (sucrose, BSA and PEG–8000) to form a water in oil emulsion. The mixture was then vortex mixed with 2% PVA (10 ml) to form a double emulsion. The resulting microspheres were harvested by centrifugation, lyophilized and kept at 80 °C prior to use. When incorporating into the hydrogels, the microspheres were suspended in ethanol to avoid aggregates formation in precursor. The suspension was then used to dissolve PEG–PLLA–DA and PEG–DA polymers and photo-initiator to form a viscous precursor solution containing the microspheres. Three hundred microliters of salt crystals of selected size was then added to the mixture (200 ll) with vortex mixing, before immediately placing the solution under UV light for photo-polymerization. 2.2.4. Fibrin loading Fibrin was loaded in the hydrogel pores as described in Jiang et al. [9]. Briefly, thrombin (100 U ml 1) was added to the hydrogel precursor prior to photo-polymerization. After polymerization and salt leaching, water on the hydrogel surface was mostly removed by blotting with sterile gauze. The hydrogels were air-dried for 1 h to allow evaporation of water from the pores. Three hundred microliters of Fg solution was dropped onto the porous hydrogel surface. For growth factor delivery, proteins ± heparin (5 U ml 1) were mixed with the fibrinogen solution prior to addition to the pores. The hydrogels were then incubated at room temperature for 30 min to allow fibrin gelation. A schematic representation of the hydrogels is shown in Fig. 1. 2.3. Characterization 2.3.1. Biomaterial variables A number of parameters were varied during hydrogel preparation, including pore size, polymer concentration, polymer ratio, PLGA microsphere concentration and fibrin concentration (Table 1). Pore size was controlled by varying the size range of the salt crystals used. The concentrations of PEG–PLLA–DA and PEG–DA in the precursor were varied, either resulting in a change in the overall polymer concentration of PEG–PLLA–DA or to investigate different polymer ratios (PEG–PLLA–DA:PEG–DA) at a constant overall polymer concentration at 250 mg ml 1. Varying amounts of PLGA microspheres were suspended in the hydrogel precursor prior to mixing with salt crystals, and the Fg concentration added to the pores was also varied. 2.3.2. Imaging In order to image the structure of the hydrogels, PLGA microspheres were prepared with 0.1% PKH-26 (hydrophobic red fluorescent dye) in the oil phase, and fibrinogen solution was mixed with 0.1 mg ml 1 of Fg conjugated with Alexa Fluor 647. PEG– PLLA–DA-based hydrogels exhibit auto-fluorescence that can be used to image a 3-D structure [10]. The hydrogels were sectioned in half from the center and imaged by confocal microscopy (Carl Zeiss AG, Germany) with fluorescence used to image PLGA microspheres (543 nm excitation, 560 nm long-pass filter, purple), fibrin
B. Jiang et al. / Acta Biomaterialia 10 (2014) 1177–1186
1179
Fig. 1. Schematic representation of composite hydrogels containing fibrin and PLGA microspheres for sequential growth factor delivery.
Table 1 Parameters varied in hydrogel synthesis and characterization. Variable
Values
Pore size (lm) Polymer concentration (mg ml PEG–PLLA–DA:PEG–DA ratio (wt.%:wt.%) Microsphere concentration (mg ml 1) Fibrinogen concentration (mg ml 1)
1
)
Non-porous,
Contents lists available at ScienceDirect
Acta Biomaterialia journal homepage: www.elsevier.com/locate/actabiomat
Design of a composite biomaterial system for tissue engineering applications B. Jiang a,b, B. Akar a,b, T.M. Waller a, J.C. Larson a,b, A.A. Appel a,b, E.M. Brey a,b,⇑ a b
Department of Biomedical Engineering, Illinois Institute of Technology, Chicago, IL, USA Research Service, Edward Hines, Jr. V.A. Hospital, Hines, IL, USA
a r t i c l e
i n f o
Article history: Received 29 July 2013 Received in revised form 10 November 2013 Accepted 29 November 2013 Available online 7 December 2013 Keywords: Neovascularization Tissue engineering Fibroblast growth factor Platelet-derived growth factor Poly(lactic-co-glycolic acid)
a b s t r a c t Biomaterials that regulate vascularized tissue formation have the potential to contribute to new methods of tissue replacement and reconstruction. The goal of this study was to develop a porous, degradable tissue engineering scaffold that could deliver multiple growth factors and regulate vessel assembly within the porous structure of the material. Porous hydrogels of poly(ethylene glycol)-co-(L-lactic acid) (PEG– PLLA) were prepared via salt leaching. The degradation time of the hydrogels could be controlled between 1 and 7 weeks, based on hydrogel composition. Fibrin was incorporated into the interconnected pores of the hydrogels to promote neovascularization and as a reservoir for rapid (30 days) growth factor delivery. Fibroblast growth factor-1 (FGF-1) and platelet-derived growth factor-BB (PDGF-BB) were delivered from the system owing to their roles in the promotion of angiogenesis and vascular stabilization, respectively. Hydrogels tested in vivo with a subcutaneous implantation model were selected based on the results from in vitro degradation and growth factor release kinetics. Dual growth factor delivery promoted significantly more tissue ingrowth in the scaffold compared with blank or single growth factor delivery. The sequential delivery of FGF-1 following PDGF-BB promoted more persistent and mature blood vessels. In conclusion, a biomaterials system was developed to provide structural support for tissue regeneration, as well as delivery of growth factors that stimulate neovascularization within the structure prior to complete degradation. Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Biomaterials are often used as scaffolds to provide mechanical and structural support in tissue engineering and organ regeneration applications [1,2]. Depending on the type of tissue to be engineered (e.g. bone, cartilage or soft tissue), biomaterials are fabricated into tissue engineering scaffolds to meet specific tissue requirements in regard to mechanical properties and degradation rates. Generally speaking, a three-dimensional scaffold with controlled degradation rate, high porosity and appropriate pore size is probably required for generating a proper engineered tissue for use in reconstruction [3]. In plastic and reconstructive surgery, coordinating the processes of tissue regeneration and biomaterial degradation is essential for controlling the geometry of the resultant tissue to match preimplantation design [4]. Tissue regeneration involves the regulation of a series of signaling molecules, such as cytokines, growth factors and hormones. Regeneration within a biomaterial may require control over the ⇑ Corresponding author at: 3255 South Dearborn Street, WH 314, Chicago, IL 60616-3793, USA. Tel.: +1 312 567 5098; fax: +1 312 567 5707. E-mail address: [email protected] (E.M. Brey).
temporal delivery of some of these molecules in order to re-create this process. For example, neovascularization occurs via a complex temporal response of cells to a number of growth factors. Angiogenic factors, such as those from the vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF) families, are involved in the initiation of angiogenesis, while vessel maturation factors, including platelet-derived growth factor (PDGF)-BB and angiopoietin-1, are involved in the later stage of vessel stabilization [5]. Release of factors with distinct kinetics is likely to improve vessel assembly, with initial administration of angiogenic factor(s) followed by vessel maturation factor(s). The sequential delivery of FGF-2 (also known as bFGF) [6] or VEGF [7] followed by PDGF-BB has been shown to increase vessel density and stability over single growth factor delivery or simultaneous dual delivery when the drug delivery system is implanted subcutaneously [8]. Previously, it was found that fibrin–poly(ethylene glycol) (PEG) composite hydrogels stimulated vascularized tissue regeneration after implantation [9]. In this system, the PEG-based porous hydrogels provided a biologically inert material with a controlled three-dimensional (3-D) structure, while fibrin served as a bioactive material that stimulated vessel invasion into the pores. However, the hydrogels exhibit little degradation under physiological conditions. In the study presented
1742-7061/$ - see front matter Ó 2013 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.actbio.2013.11.029
1178
B. Jiang et al. / Acta Biomaterialia 10 (2014) 1177–1186
here, monomer units of poly(L-lactic acid) (PLLA) were introduced into the PEG structure to allow controlled degradation via hydrolysis to decouple degradation from tissue invasion [10]. Two growth factors (FGF-1 and PDGF-BB) involved in neovascularization were incorporated into the hydrogels in order to influence the density, persistence and maturity of the vessels formed. Previous studies evaluating the effect of sequential growth factor delivery on vascularization were focused mostly on the response in tissues surrounding the implants or following complete biomaterial degradation [6–8]. However, biomaterial scaffolds are often designed to precisely the structure and volumetric features of a given tissue defect. Engineering tissues of a specific 3-D shape requires that vascularized tissue invasion occurs simultaneously with implant degradation [4]. In fact, vascularized tissue must grow into the scaffold prior to degradation in order to maintain the preformed structure of the material. In the present study, vascular structure within the scaffold structure prior to degradation was evaluated. 2. Materials and methods 2.1. Materials PLGA 85:15 (Mw 50,000–75,000, acid terminated), poly(vinyl alcohol) (PVA, Mw 13,000–23,000), bovine serum albumin (BSA), sucrose (>99.5%), magnesium hydroxide (95%), PEG (Mw 3350), 3,6-dimethyl-1,4-dioxane-2,5-dione, Tin(II) 2-ethylhexanoate (95%), fibrinogen from human plasma (Fg, 50–70% protein), thrombin from human plasma (Tb), 2-hydroxy-2-methylpropiophenone (Irgacure 1173), acryloyl chloride (98%) and thriethylamine (TEA, 99.5%) were obtained from Sigma-Aldrich (St. Louis, MO). Recombinant rat PDGF-BB with and without carrier was obtained from R&D Systems (Minneapolis, MN). Recombinant rat FGF-1 was obtained from PeproTech (Rocky Hill, NJ). Iodine 125 was obtained from Perkin Elmer Inc. (Waltham, MA). Dichloromethane (DCM, 99.9%), sodium chloride (99.5%), magnesium sulfate anhydrous (97%), diethyl ether (anhydrous), ethyl alcohol (95%), PEG (Mw 8000), Shandon xylene substitute and phosphate buffered saline (PBS) solution were obtained from Fisher Scientific (Hampton, NH). Endothelial cell growth medium (EGM) and endothelial cell basal medium (EBM) were obtained from Lonza Inc. (Allendale, NJ). 2.2. Hydrogel synthesis 2.2.1. PEG–PLLA–DA and PEG–DA synthesis The procedure for synthesis of PEG–PLLA–DA has been described in detail elsewhere [10]. Briefly, 20 g of PEG (Mw 3350) was mixed with 4.24 g of 3,6-dimethyl-1,4-dioxane-2,5-dione (L-lactide) in the presence of 80 ll of stannous octoate. The mixture was allowed to react at 140 °C for 4 h under anhydrous conditions. After purification and lyophilization, the resulting PEG–PLLA product was reacted with acryloyl chloride (4 molar ratio) and TEA (2 molar ratio) to acrylate the polymer. The PEG–PLLA–DA was purified, lyophilized and stored at 20 °C in the dark prior to use. PEG–DA was synthesized by acrylating PEG (Mw 3350), using the same method. 2.2.2. Porous hydrogel fabrication A salt leaching technique was adapted for generation of porous PEG–PLLA–DA hydrogels [11]. Briefly, PEG–PLLA–DA alone or in combination with PEG–DA was dissolved in ethanol, with 2-hydroxy-2-methylpropiophenone (0.5% v/v) added as a photo-initiator. Sodium chloride was ground with a pestle and mortar and sieved to select crystals of a defined size range. Three hundred milligrams
of salt crystals were added per 200 ll of precursor solution and polymerized under UV light (k = 365 nm) for 5 min. The hydrogels were then incubated in 50 ml of deionized (DI) water overnight to leach out the salt crystals and ethanol. 2.2.3. PLGA microspheres The preparation of PLGA microspheres for protein delivery was performed using a water in oil in water emulsion as described previously [12]. Briefly, PLGA 85:15 (250 mg ml 1) was dissolved in DCM (oil phase) with Mg(OH)2 (7.5 mg ml 1) added as an anti-acid agent. The oil phase mixture (1 ml) was emulsified with PBS (0.2 ml) containing protein (5 lg ml 1 of either PDGF-BB or FGF1) and protein stabilizers (sucrose, BSA and PEG–8000) to form a water in oil emulsion. The mixture was then vortex mixed with 2% PVA (10 ml) to form a double emulsion. The resulting microspheres were harvested by centrifugation, lyophilized and kept at 80 °C prior to use. When incorporating into the hydrogels, the microspheres were suspended in ethanol to avoid aggregates formation in precursor. The suspension was then used to dissolve PEG–PLLA–DA and PEG–DA polymers and photo-initiator to form a viscous precursor solution containing the microspheres. Three hundred microliters of salt crystals of selected size was then added to the mixture (200 ll) with vortex mixing, before immediately placing the solution under UV light for photo-polymerization. 2.2.4. Fibrin loading Fibrin was loaded in the hydrogel pores as described in Jiang et al. [9]. Briefly, thrombin (100 U ml 1) was added to the hydrogel precursor prior to photo-polymerization. After polymerization and salt leaching, water on the hydrogel surface was mostly removed by blotting with sterile gauze. The hydrogels were air-dried for 1 h to allow evaporation of water from the pores. Three hundred microliters of Fg solution was dropped onto the porous hydrogel surface. For growth factor delivery, proteins ± heparin (5 U ml 1) were mixed with the fibrinogen solution prior to addition to the pores. The hydrogels were then incubated at room temperature for 30 min to allow fibrin gelation. A schematic representation of the hydrogels is shown in Fig. 1. 2.3. Characterization 2.3.1. Biomaterial variables A number of parameters were varied during hydrogel preparation, including pore size, polymer concentration, polymer ratio, PLGA microsphere concentration and fibrin concentration (Table 1). Pore size was controlled by varying the size range of the salt crystals used. The concentrations of PEG–PLLA–DA and PEG–DA in the precursor were varied, either resulting in a change in the overall polymer concentration of PEG–PLLA–DA or to investigate different polymer ratios (PEG–PLLA–DA:PEG–DA) at a constant overall polymer concentration at 250 mg ml 1. Varying amounts of PLGA microspheres were suspended in the hydrogel precursor prior to mixing with salt crystals, and the Fg concentration added to the pores was also varied. 2.3.2. Imaging In order to image the structure of the hydrogels, PLGA microspheres were prepared with 0.1% PKH-26 (hydrophobic red fluorescent dye) in the oil phase, and fibrinogen solution was mixed with 0.1 mg ml 1 of Fg conjugated with Alexa Fluor 647. PEG– PLLA–DA-based hydrogels exhibit auto-fluorescence that can be used to image a 3-D structure [10]. The hydrogels were sectioned in half from the center and imaged by confocal microscopy (Carl Zeiss AG, Germany) with fluorescence used to image PLGA microspheres (543 nm excitation, 560 nm long-pass filter, purple), fibrin
B. Jiang et al. / Acta Biomaterialia 10 (2014) 1177–1186
1179
Fig. 1. Schematic representation of composite hydrogels containing fibrin and PLGA microspheres for sequential growth factor delivery.
Table 1 Parameters varied in hydrogel synthesis and characterization. Variable
Values
Pore size (lm) Polymer concentration (mg ml PEG–PLLA–DA:PEG–DA ratio (wt.%:wt.%) Microsphere concentration (mg ml 1) Fibrinogen concentration (mg ml 1)
1
)
Non-porous,
