Biocompatible electrospun polymer blends for biomedical applications Hrishikesh Ramesh Munj,1 M. Tyler Nelson,2 Prathamesh Sadanand Karandikar,1 John Joseph Lannutti,3 David Lane Tomasko1 1
Department of Chemical and Biomolecular Engineering, Ohio State University, Columbus, Ohio 43210 Department of Biomedical Engineering, Ohio State University, Columbus, Ohio 43210 3 Materials Science and Engineering, Ohio State University, Columbus, Ohio 43210 2
Received 26 July 2013; revised 26 January 2014; accepted 18 February 2014 Published online 00 Month 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.b.33132 Abstract: Blends of natural and synthetic polymers have received considerable attention as biomaterials due to the potential to optimize both mechanical and bioactive properties. Electrospinning of biocompatible polymers is an efficient method producing biomimetic topographies suited to various applications. In the ultimate application, electrospun scaffolds must also incorporate drug/protein delivery for effective cell growth and tissue repair. This study explored the suitability of a ternary Polymethylmethacrylate-Polycaprolactone-gelatin blend in the preparation of electrospun scaffolds for biomedical applications. Tuning the blend composition allows control over scaffold mechanical properties and degradation rate. Significant improvements were observed in the mechanical properties of the blend compared with the individual components. In order to study drug delivery potential, triblends were impregnated with the model compound Rhodamine-B
using sub/supercritical CO2 infusion under benign conditions. Results show significantly distinct release profiles of the impregnated dye from the triblends. Specific factors such as porosity, degradation rate, stress relaxation, dye-polymer interactions, play key roles in impregnation and release. Each polymer component of the triblends shows distinct behavior during impregnation and release process. This affects the aforementioned factors and the release profiles of the dye. Careful control over blend composition and infusion conditions creates the flexibility needed to produce biocompatible electrospun scaffolds for a variety of biomedical applications. C 2014 Wiley Periodicals, Inc. J Biomed Mater Res Part B: Appl BioV
mater 00B:000–000, 2014.
Key Words: electrospinning, polymer blends, impregnation, supercritical CO2
How to cite this article: Ramesh Munj H, Tyler Nelson M, Sadanand Karandikar P, Joseph Lannutti J, Lane Tomasko D. 2014. Biocompatible electrospun polymer blends for biomedical applications. J Biomed Mater Res Part B 2014:00B:000–000.
INTRODUCTION
Many challenges are faced in the development of tissue engineering scaffolds capable of controlled drug delivery, the desired cellular response, mechanical stability and tailored in vivo degradation.1 Electrospinning is a simple process to prepare polymeric fibrous scaffolds with high surface area and porosity required for drug delivery and tissue engineering.2 Electrospinning presents a threedimensional, porous microenvironment closely mimicking the extracellular matrix thus providing a suitable platform for the proliferation and differentiation of many different cell types.3 The efficacy of such scaffolds is based on several parameters such as cellular adhesion and proliferation in addition to structural integrity, void volume and fiber diameter.4,5 The design of the optimal scaffold continues to present a series of complex challenges due to these competing demands. Biocompatible polymers are widely used in tissue engineering and controlled drug release systems.6 Polycaprolactone (PCL) is an attractive option for tissue
engineered implants due to its mechanical properties and intermediate resistance to degradation.7 Electrospun PCLbased scaffolds have been used as tissue engineered scaffolds for in vitro development of blood vessels and nerve regeneration.8,9 Poly-methyl methacrylate (PMMA), on the other hand, is a synthetic, nondegradable biocompatible polymer historically useful in bone repair and hard tissue regeneration.10 Electrospun PMMA scaffolds have demonstrated good cellular adhesion, proliferation, and viability for in vitro and in vivo tissue engineering applications.11,12 Scaffold biofunctionality is essential to efficient cell adhesion and proliferation. A biofunctionalized scaffold has active sites recognizable by cells providing enhanced cell attachment.13 It has been shown that use of active molecules derived from natural extra cellular matrix proteins can supply these functionalized scaffolds. As an attempt to improve upon the biochemical properties of the two synthetic materials, gelatin has also been chosen for incorporation into these electrospun composite fiber scaffolds, creating improved cellular adhesion and proliferation.14,15
Correspondence to: D. Tomasko; e-mail: [email protected] Contract grant sponsor: National Science Foundation; contract grant number: EEC-0914790
C 2014 WILEY PERIODICALS, INC. V
1
Such precise compositional adjustments give control over blend mechanical and biological properties. Blends of natural polymers with synthetic polymers and these have proven to perform better in tissue engineering applications than their purely synthetic counterparts .16 Gelatin acts as a biological cue for cells and provides better biocompatibility than a synthetic polymer. It can also moderate the hydrophobic nature of some synthetic polymers to provide an added advantage.17 Supercritical fluids (SCF) have easily tunable physical properties (e.g. density, viscosity, etc.) useful in a wide range of applications.18 Carbon dioxide (CO2) is one of the most popular SCF due to its benign critical point conditions (7.38 MPa at 31.1 C). Recent studies have explored the potential of CO2 in sustainable and inexpensive processing of many polymers during foaming, sterilization, solvent extraction and infusion.19,20 CO2 has specific molecular interactions with many polymers to result in altered polymer properties that ease plasticization.21 In the presence of CO2, properties of many polymers (such as glass transition temperature Tg, etc.) are significantly lower than the pure form and can be further tuned by adjusting CO2 pressure and temperature.22 CO2 offers a green alternative way for the infusion of drugs and biomolecules into polymer substrates for tissue engineering. Various studies have already considered the plasticization and impregnation of PCL and PMMA in high pressure CO2.23,24 Although subcritical gaseous CO2 can reversibly swell PCL during impregnation, it results in fast release of the infused additive due its low solubility in dense CO2. Liquid CO2 can swell pure PCL reversibly without permanent deformation but has low diffusivity. Supercritical CO2 (SCCO2) has moderate diffusivity in polymers and adequate solubility for additives compared with that of dense gas or liquid CO2. Unfortunately, if pure PCL is subjected to SCCO2, it melts and loses mechanical integrity.25 Similarly, PMMA shows an irreversible change in the structure if processed under certain SCCO2 conditions.26,27 Preserving the initial structure of a scaffold (the porosity of foams and the fiber structure characteristic of fibrous scaffolds) is a factor essential to cell proliferation. In our recent work, we showed that if PCL is blended with gelatin the resulting blend exhibits distinct properties. When subjected to high pressure CO2, the blend can be swelled reversibly by dense, liquid and SCCO2.28 Our recent results have confirmed that pure PCL swells under high pressure CO2 whereas pure gelatin compresses due to dehydration in the presence of CO2. Simultaneous swelling of PCL and shrinking of gelatin stabilize the blend under SCCO2 without deformation. This discovery opens new avenue to explore CO2 assisted infusion into biocompatible blends containing gelatin for drug delivery and tissue engineering applications. The addition of PMMA to the existing PCL-Gelatin binary blend increases loading capacity of the blend due to amorphous nature of PMMA. Recently, several studies have investigated potential of polymeric ternary blends in the biomedical applications.29–32
2
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
In this article, we have explored the postinfusion properties of PMMA-PCL-gelatin ternary blends for biomedical applications. The major goal of this investigation is to understand the influence of blend composition and processing factors simultaneously on mechanical properties, impregnation and release from complex electrospun polymer matrices. Adjusting the blend ratio of the individual components allows control over mechanical properties and degradation rate. High pressure CO2 was used to assist in the infusion of Rhodamine B dye. Dye release from these scaffolds was studied over specific periods of time. Results indicate that PMMAPCL-gelatin ternary blends have promising properties for various biomedical applications. METHOD AND MATERIALS
Electrospinning Initially, 5% w/w polycaprolactone (PCL) (Mn 70,000– 90,000; Sigma-Aldrich, St. Louis, MO), 3% polymethyl methacrylate (PMMA) (Mn 350,000; Sigma-Aldrich, St. Louis, MO) and 6.7% w/w porcine gelatin type A (300 Bloom; SigmaAldrich, St. Louis, MO) solutions were prepared in 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) (>99% purity; Sigma-Aldrich, St. Louis, MO) by stirring at room temperature (ffi25 C) for 24 h. Weight % concentrations were chosen based on electrospinning of pure polymer to achieve required fiber morphology. Different ternary blend compositions were obtained on a volume/volume basis by adjusting the amount of each component in the blended solution. After complete mixing of the solutions, Tri-PCL (PCL 50%, PMMA 25%, Gelatin 25%), Tri-PMMA (PCL 25%, PMMA 50%, Gelatin 25%), and Tri-Gelatin (PCL 25%, PMMA 25%, Gelatin 50%) blends were then poured individually into separate 20 mL syringes, fitted with a 20-gauge blunt tip needle, and electrospun using a DC high voltage power supply (Glassman High Voltage, Inc., High Bridge, NJ) at positive 20 kV, 20 cm needle-to-collector distance, 10 mL/hr flow rate, for 45 to 60 min; 8 cm 3 8 cm randomly oriented electrospun fiber mats of Tri-PCL, Tri-PMMA and Tri-Gelatin were produced on aluminum non-stick foil. The as-spun mats were then placed in a vacuum oven (99.5% purity, 200 Proof; Sigma-Aldrich, St. Louis, MO) solution was placed uniformly onto each of the scaffolds. These scaffolds were placed in chemical hood to dry for 2 h until all the ethanol was evaporated. Scaffolds were moved using tweezers in order to avoid their adhesion to the well plate surface. In the case of adsorption-only study, scaffolds were washed three times with pure ethanol after 30 min. For CO2 infusion, dried scaffolds were then placed into a high pressure stainless steel vessel. Temperature of the vessel was controlled using heat tape and monitored using a temperature sensor (Omega CSC 32, Stamford, CT). Bonedry CO2 (99.9% purity, Praxair, Columbus, OH) was allowed to flow in the pressure vessel and pressure was controlled with the syringe pump (500 D; Teledyne ISCO, Inc., Lincoln, NE). Equilibrium conditions were maintained for 2 h. Subcritical CO2 infusion was conducted at 6.20 MPa and 25 C whereas supercritical CO2 conditions were maintained at 8.27 MPa and 37 C. In either case, pressure was released by reversing syringe pump slowly (0.5–1 mL/min) over >12 h. Scaffolds were then washed with pure ethanol three times and dried in the chemical hood. Rhodamine B release was carried out in 24-well plates for all three impregnation conditions (adsorption, subcritical CO2 and supercritical CO2) by immersing each scaffold in the 2 mL of phosphate buffer saline (PBS) and keeping it in the incubator (Wiseven, Witeg Labortechnik GmbH, Germany) at 37 C. At each time point, two aliquots were taken from each sample and stored in 96-well plates. PBS was replaced with fresh PBS in each well after every time point. All 96-well plates containing aliquots were stored in the refrigerator. A UV-Vis 96-well plate spectrometer (Spectra Max 190 Absorbance UV-VIS plate reader, Sunnyvale, CA) was used to measure the absorbance of released Rhodamine B solutions at a wavelength of 535 nm. Spectrometer readings were then converted to concentration values (lg/mL) using a polynomial calibration curve produced by making serial dilutions of Rhodamine B in PBS. Weight loss study Similar to the infusion study, 13 mm diameter discs of TriPMMA, Tri-PCL and Tri-Gelatin fibers were weighed and placed in 24-well plates. Each scaffold was immersed in 2 mL of PBS and stored in the incubator (Wiseven, Witeg Labortechnik GmbH, Germany) at 37 C. Scaffolds were then removed after specific time interval to wash with distilled water in order to get rid of salts. These scaffolds were then dried overnight in a vacuum oven to remove all water before they were weighed again to analyze weight loss. Statistical analysis Statistical analysis was performed using JMP software. Each treatment was represented by at least three independent experiments. Comparisons between multiple treatments were made by fitting model for response against different variable and then comparing means for different pairs using Tuckey HSD. A p value of PMMA > gelatin. Among the different blends, a similar trend is observed as that of pure polymers i.e., the elongation of Tri-PCL >TriPMMA > Tri-Gelatin. Figure 6(C) shows ultimate tensile strength (UTS) measurements for pure components and ternary blends. There is notable enhancement in the UTS of Tri-PMMA and Tri-Gelatin blends when evaluated against their pure states. Although Tri-PCL blend shows some improvement over pure PCL it is not as high as PMMA and gelatin. Mechanical testing shows that a remarkable change in the properties of the blends compared with the individual components. PCL is a semi-crystalline polymer with lower modulus and higher elongation.42 Being above the glass
FIGURE 6. Comparison among different mechanical properties of different compositions of PMMA-PCL-gelatin blends against PCL, PMMA, and gelatin in their pure state A) modulus, B) elongation, C) ultimate tensile strength.
maintaining the fiber structure. However, for Tri-PMMA and Tri-Gelatin, PCL is a minor component. DSC and XRD confirmed that PCL is unable to form crystals in these blends.
FIGURE 7. Different factors affecting various blend components during impregnation (ethanol and high pressure CO2) and release (release media diffusion and degradation) processes.
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
7
FIGURE 8. Comparison of release profiles of Rhodamine B dye from ternary blends after different impregnation conditions.
transition temperature under ambient conditions, low stress is required to achieve large strain in PCL. PMMA is brittle material with lower elongation and high compressive strength at room temperature.43 The mechanical properties of gelatin are closely related to water content.44 As water
8
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
content is reduced, it changes the gel behavior to that of a rigid polymer. Electrospun gelatin fibers in solid state at ambient conditions contain about 10 to 15% moisture.45 Thus, pure gelatin fibers exhibit higher moduli and lower elongation under ambient conditions. Although blends
MUNJ ET AL.
ORIGINAL RESEARCH REPORT
prepared for this study are heterogeneous blends, each component contributes to the improvement of overall mechanical properties. Gelatin increases the modulus of the blends, whereas PCL enhances their elongation and UTS. Release of a solute from polymer blends is a complex process. Each polymer in the blend has a unique capacity for solute release based on polymer-solute interactions as well as infusion conditions.46 Apart from these factors, degradation and morphology of the polymer matrix in the release medium can play an essential role in the release.47 If we consider solute-polymer interactions for ternary blends, comparing solubility parameters for different polymers with Rhodamine B can provide a basis for comparison. The solubility parameters of PCL, PMMA and gelatin are 18.25, 18.2, and 24.3 MPa5 respectively.48–50 The solubility parameter of Rhodamine B estimated using group contribution theory is about 27.4 MPa0.5.51 Thus, gelatin should have higher compatibility for Rhodamine B as compared with PCL and PMMA. The blends have mixtures of semi-crystalline PCL, amorphous PMMA and gelatin all of which show different behavior when subjected to high pressure CO2. As discussed earlier, PMMA and PCL show significant increases in volume (swelling) when subjected to high pressure CO2. However, gelatin shows compression due to dehydration in the presence of CO2. Also these polymers differ in terms of their stability in aqueous release media. PMMA is nonbiodegradable polymer which cannot be degraded hydrolytically on a relevant time scale.52 PCL, a poly-hydroxy ester, shows slow hydrolysis of ester groups in PBS which proceeds via bulk degradation.53 Uncrosslinked gelatin is extremely unstable in aqueous media and shows rapid degradation due to hydrolytic or enzymatic degradation.54 Furthermore, relaxation dynamics of polymers play an important role in the impregnation and release. PMMA and gelatin electrospun fibers in their pure state show prominent reduction in the voidage of electrospun mat when subjected to 100% ethanol. This phenomenon is explained by the relaxation of stresses built up during the extremely rapid electrospinning process. As shown in the SEM analysis [Figure 1(E,F)], both Tri-PMMA and Tri-Gelatin blends show decrease in the void volume of an electrospun mat when subjected to Rhodamine B in ethanol. For Tri-gelatin fibers, a large degree of melting was observed, whereas, Tri-PMMA fibers show no melting but reduction in the void volume and change in fiber structure. Tri-PCL scaffolds show no change in structure after pure ethanol wetting. Figure 7 shows how different factors affect each component of the ternary blend during impregnation and release. Release from the blended electrospun fiber mats is shown in Figure 8. All release curves are expressed in terms of the cumulative mass of Rhodamine B per mass of fiber to compare different blends and different impregnation conditions. Figure 8(A) shows release from Tri-Gelatin fibers infused via simple adsorption, subcritical CO2 and supercritical CO2 conditions. Subcritical CO2 infusion releases significantly more dye with time than adsorption and supercritical CO2 infused samples that are not significantly different from each other. Release curves from Tri-PCL fibers [Figure 8(B)]
FIGURE 9. Percentage weight loss of electrospun ternary blends over time in PBS.
exhibit an obvious trend for the mass of released dye, where adsorption < subcritical CO2 < supercritical CO2. In the case of Tri-PMMA fibers [Figure 8(C)], all three infusion conditions show similar release profiles without notable change. Subcritical CO2 infused samples show a small but not prominent increase in the amount of released dye as compared with supercritical CO2 and adsorption which overlap with each other. Figure 8(D–F) show different blend compositions subjected to a specific impregnation condition. All three blend compositions show significantly different cumulative amounts released for each time point for adsorption [Figure 8(D)]. Tri-Gelatin shows the highest release followed by Tri-PMMA and finally Tri-PCL. Subcritical and supercritical conditions show Tri-PMMA and Tri-PCL are significantly different from Tri-Gelatin but not from each other [Figure 8(E,F)]. As discussed earlier, we see that release of a solute from polymer blends is a complex process and each polymer in the blend can show a unique response towards solute release rate. Among all three blends, Tri-Gelatin samples show highest amount of release of dye (per mg of fibers) for all impregnation conditions for the following reasons: (1) during impregnation large amount of Rhodamine B is retained in the TriGelatin scaffolds due to hydrophilic interactions between the dye and the gelatin; (2) although melting of scaffolds results in lower diffusion of release medium in the scaffolds, faster degradation of gelatin causes quick release of dye; (3) pure gelatin is compressed in the presence of high pressure CO2 preventing significant depth of infusion within the fiber. Nevertheless, subcritical CO2 shows smaller degree of compression for gelatin as compared with supercritical CO2 swelling. In the case of Tri-PCL fibers, a similar trend was observed as in the PCL-gelatin binary blend.28 Supercritical CO2 causes more swelling for PCL in contrast to subcritical CO2. Although gelatin is present in the blend, a compression effect is not dominant in the case of Tri-PCL fibers. All release curves of Tri-PMMA almost overlap for each of the impregnation conditions. Tri-PMMA fibers show no
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
9
TABLE I. Diffusion Coefficients (D) of Rhodamine B from Ternary Blends Infused Using Different Impregnation Conditions in a Static Release Medium (PBS).
Adsorption Subcritical CO2 Supercritical CO2
DTri-PMMA (mm2/sec)
DTri-PCL (mm2/sec)
DTri-Gelatin (mm2/sec)
4.8 3 10210 6.4 3 10210 7.7 3 10210
6.9 3 10210 9.92 3 10210 9.85 3 10210
9.3 3 10210 18.4 3 10210 20.6 3 10210
significant change in the dye released under different impregnation conditions. The most likely explanation for this abnormal behavior is restricted diffusion of release medium in the Tri-PMMA samples due to reduction in void volume. Also subcritical CO2 and supercritical CO2 show similar swelling for PMMA when subjected below 40 C.55 Figure 9 shows results from the weight loss study indicating Tri-Gelatin degrades rapidly in PBS followed by Tri-PCL and finally Tri-PMMA. If we model these scaffolds by considering a thin film, we can calculate Fickian diffusion coefficients of dye release using procedure explained by Ritger et al.56 Diffusion coefficients (Table I) indicate that all three compositions have different limiting factors in the release of dye. Tri-PMMA fibers show slowest release with lowest diffusion coefficients. This is the result of formation of fibrous membrane of Tri-PMMA scaffolds after stress relaxation. Considering Tri-PCL scaffolds can maintain the porosity during the release, higher diffusion coefficients are obtained. In the case of Tri-Gelatin fibers, rapid dissolution of gelatin releases dye quickly compared with Tri-PCL and Tri-PMMA scaffolds (Figure 9). CONCLUSION
In this study, standard electrospinning procedure was used to prepare ternary blend scaffolds of Tri-PCL (PCL 50%, PMMA 25%, gelatin 25%), Tri-PMMA (PCL 25%, PMMA 50%, gelatin 25%), and Tri-Gelatin (PCL 25%, PMMA 25%, GELATIN 50%) blends. We have explored a polymer blend with three components having unique characteristics. PCL boosts the elongation capacity of the blend and aids in maintaining structural properties during further processing of blend. Gelatin imparts higher modulus and the ability to achieve rapid release of the infused drug. Additionally, it makes the blend bioactive and allows efficient adhesion of proteins and cells. PMMA can contribute in compressive strength of scaffold and allows longer degradation periods. Stress relaxation of PMMA fibers lead to loss of interfiber porosity. However, it maintains the fiber structure and thus has high surface area. This can be used in dermal and transdermal patches for long-term drug delivery. Significant improvement was obtained in the mechanical properties of ternary blends as compared with individual components of the blend. DSC and XRD studies show noteworthy changes in the crystalline properties of PCL in the ternary blend. CO2 assisted impregnation of drugs and biomolecules in polymers is simple, benign and “green” alternative to current impregnation processes. A dye release study was carried out from all three blend compositions under various
10
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
impregnation conditions. Stress relaxation of PMMA and gelatin in the presence of ethanol during impregnation process changes morphological properties of Tri-PMMA and TriGelatin blend but Tri-PCL remains unchanged due to high crystalline content. Three major factors were found to have a significant impact on the dye release profiles from ternary blends: (1) effect of CO2; (2) degradation rate of scaffold; (3) porosity of scaffold. The Tri-PCL blended scaffolds display long degradation rates and maintain porosity of the scaffold in the release medium. The effect of CO2 was dominant in the case of Tri-PCL blends due to melting and recrystallization on PCL in the blend after CO2 treatment. For Tri-Gelatin scaffolds, degradation rate was found to be a key factor in rapid release of the dye from scaffolds. Also ethanol treatment shows significant melting of Tri-Gelatin fibers which increases the loading of the dye in the blend. In the case of Tri-PMMA, stress relaxation in the presence of ethanol reduces the porosity of the scaffold. Although PMMA shows an increase in swelling under high pressure CO2, release of Rhodamine B is limited by porosity of the scaffold and diffusion of release medium in the scaffold. In summary, all three polymers show unique contribution in altering the properties of the ternary blends. Composition of the ternary blend was found to be a vital component in deciding several properties of these electrospun scaffolds. Adjusting the composition of PMMA-PCL-gelatin ternary blends and tuning impregnation conditions, several tissue engineered scaffolds and controlled drug delivery systems can be designed for efficient biomedical applications. REFERENCES 1. Lee K, Silva EA, Mooney DJ. Growth factor delivery-based tissue engineering: general approaches and a review of recent developments. J R Soc Interface 2011;8:153–170. 2. Nisbet DR, Forsythe JS, Shen W, Finkelstein DI, Horne MK. Review paper: A review of the cellular response on electrospun nanofibers for tissue engineering. J Biomater Appl 2009;24:7–29. 3. Jagur-Grodzinski J. Polymers for tissue engineering, medical devices, and regenerative medicine. Concise general review of recent studies. Polym Adv Technol 2006;17:395–418. 4. Gloria A, De Santis R, Ambrosio L. Polymer-based composite scaffolds for tissue engineering. J Appl Biomater Biomechanics 2010;8:57–67. 5. Place ES, George JH, Williams CK, Stevens MM. Synthetic polymer scaffolds for tissue engineering. Chem Soc Rev 2009;38: 1139–1151. 6. Ulery BD, Nair LS, Laurencin CT. Biomedical applications of biodegradable polymers. J Polym Sci Part B: Polym Phys 2011;49: 832–864. 7. Dash TK, Badireenath Konkimalla V. Poly-¾-caprolactone based formulations for drug delivery and tissue engineering: A review. J Controlled Release 2012;158:15–33. 8. Vaz CM, Van Tuijl S, Bouten CVC, Baaijens FPT. Design of scaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique. Acta Biomaterialia 2005;1:575–582. 9. Ghasemi-Mobarakeh L, Prabhakaran MP, Morshed M, NasrEsfahani MH, Ramakrishna S. Electrospun poly(e-caprolactone)/ gelatin nanofibrous scaffolds for nerve tissue engineering. Biomater 2008;29:4532–4539. 10. Mano JF, Sousa RA, Boesel LF, Neves NM, Reis RL. Bioinert, biodegradable and injectable polymeric matrix composites for hard tissue replacement: State of the art and recent developments. Compos Sci Technol 2004;64:789–817. 11. Zhang F, Sun F, Van Kan JA, Shao PG, Zheng Z, Ge RW, Watt F. Measurement of cell motility on proton beam micromachined 3D
MUNJ ET AL.
ORIGINAL RESEARCH REPORT
12. 13.
14.
15. 16.
17.
18.
19.
20.
21.
22.
23.
24.
25.
26.
27.
28.
29.
30.
31.
32.
33.
scaffolds. Nucl Instrum Methods Phys Res Sect B 2005;231:413– 418. Srouji S, Kizhner T, Livne E. 3D scaffolds for bone marrow stem cell support in bone repair. Future Med 2006;1:519–528. Liu X, Holzwarth JM, Ma PX. Functionalized synthetic biodegradable polymer scaffolds for tissue engineering. Macromol Biosci 2012;12:911–919. Liu X, Won Y, Ma PX. Porogen-induced surface modification of nano-fibrous poly(L-lactic acid) scaffolds for tissue engineering. Biomaterials 2006;27:3980–3987. Liu X, Won Y, Ma PX. Surface modification of interconnected porous scaffolds. J Biomed Mater Res A 2005;74:84–91. Sionkowska A. Current research on the blends of natural and synthetic polymers as new biomaterials: Review. Prog Polym Sci 2011;36.9:1254–1276. Pant HR, Kim CS. Electrospun gelatine/nylon-6 composite nanofibers for biomedical applications. Polym Int 2012;62:1008– 1013. Duarte ARC, Mano JF, Reis RL. Supercritical fluids in biomedical and tissue engineering applications: A review. Int Mater Rev 2009;54:214–222. Bhamidipati M, Scurto AM, Detamore MS. The future of carbon dioxide for polymer processing in tissue engineering. Tissue Eng Part B: Rev 2013;19:221–232. Duarte ARC, Santo VE, Alves A, Silva SS, Moreira-Silva J, Silva TH, Marques AP, Sousa RA, Gomes ME, Mano JF, Reis RL. Unleashing the potential of supercritical fluids for polymer processing in tissue engineering and regenerative medicine. J Supercrit Fluids 2013;79:177–185. Yanhui Y, Teja AS. Quantification of specific interactions between CO2 and the carbonyl group in polymers via ATR-FTIR measurements. J Supercrit Fluids 2011;56:208–212. Tomasko DL, Li H, Liu D, Han X, Wingert MJ, Lee LJ, Koelling KW. A Review of CO2 Applications in the processing of polymers. Ind Eng Chem Res 2003;42:6431–6456. Roshan Y, Mammucari R, Foster NR. Impregnation of ibuprofen into polycaprolactone using supercritical carbon dioxide. J Phys: Conf Ser 2010;215:012087. Yazan HA, Grant CS. Ibuprofen impregnation into submicron polymeric films in supercritical carbon dioxide. J Supercrit Fluids 2012;71:127–135. Ayodeji O, Graham E, Kniss D, Lannutti JJ, Tomasko DL. Carbon dioxide impregnation of electrospun polycaprolactone fibers. J Supercrit Fluids 2007;41:173–178. Kamiya Y, Mizoguchi K, Terada K, Fujiwara Y, Wang JS. CO2 sorption and dilation of poly(methyl methacrylate). Macromolecules 1998;31:472–478. Di Noto V, Vezzu K, Giffin GA, Conti F, Bertucco A. Effect of High Pressure CO2 on the Structure of PMMA: A FT-IR Study. J Phys Chem B 2011;115:13519–13525. Nelson MT, Munj HR, Tomasko DL, Lannutti JJ. Carbon dioxide infusion of composite electrospun fibers for tissue engineering. J Supercrit Fluids 2012;70:90–99. Kumari S, Basha SK, Sudha PN. Physicochemical and morphological evaluation of chitosan/poly (vinyl alcohol)/methylcellulose chemically cross-linked ternary blends. Polym Bull 2012;68:1387–1393. Liu H, Song W, Chen F, Guo L, Zhang J. Interaction of microstructure and interfacial adhesion on impact performance of polylactide (PLA) ternary blends. Macromolecules 2011;44: 1513–1522. Grande R, Carvalho AJ. Compatible ternary blends of chitosan/ poly (vinyl alcohol)/poly (lactic acid) produced by oil-in-water emulsion processing. Biomacromolecules 2011;12:907–914. Mudigoudra BS, Masti SP, Chougale RB. Thermal behavior of poly (vinyl alcohol)/poly (vinyl pyrrolidone)/chitosan ternary polymer blend films. Res J Recent Sci 2012;1:83–86. Gu SY, Wang ZM, Ren J, Zhang CY. Electrospinning of gelatin and gelatin/poly(L-lactide) blend and its characteristics for wound dressing. Mater Sci Eng C 2009;29:1822–1828.
34. Kim CH, Cho KY, Choi EJ, Park JK. Effect of P(lLA-co-eCL) on the compatibility and crystallization behavior of PCL/PLLA blends. J Appl Polym Sci 2000;77:226–231. 35. Domanska U, Pobudkowska A, Pelczarska A, Gierycz P. pKa and solubility of drugs in water, ethanol, and 1-octanol. J Phys Chem B 2009;113:8941–8947. 36. Jopling DW. Stress relaxation studies of chemically crosslinked gelatin films. J Polym Sci Part A: Gen Pap 1965;3:513–526. 37. Sarti GC, Gostoli C, Masoni S. Diffusion of alcohols and relaxation in poly(methyl methacrylate): Effect of thermal history. J Membr Sci 1983;15:181–192. 38. Matthews RG, Ajji A, Dumoulin MM, Prud’homme RE. The effects of stress relaxation on the structure and orientation of tensile drawn poly(ethylene terephthalate). Polymer 2000;41: 7139–7145. 39. Shen Z, McHugh MA, Xu J, Belardi J, Kilic S, Mesiano A, Bane S, Karnikas C, Beckman E, Enick R. CO2-solubility of oligomers and polymers that contain the carbonyl group. Polymer 2003;44:1491– 1498. 40. Shieh YT, Yang HS. Morphological changes of polycaprolactone with high-pressure CO2 treatment. J Supercrit Fluids 2005;33:183– 192. 41. Liu D, Li H, Noon MS, Tomasko DL. CO2-Induced PMMA swelling and multiple thermodynamic property analysis using sanchez lacombe EOS. Macromolecules 2005;38:4416–4424. 42. Tan EPS, Ng SY, Lim CT. Tensile testing of a single ultrafine polymeric fiber. Biomaterials 2005;26:1453–1456. 43. Mousa WF, Kobayashi M, Shinzato S, Kamimura M, Neo M, Yoshihara S, Nakamura T. Biological and mechanical properties of PMMA-based bioactive bone cements. Biomaterials 2000;21: 2137–2146. 44. Yakimets I, Wellner N, Smith AC, Wilson RH, Farhat I, Mitchell J. Mechanical properties with respect to water content of gelatin films in glassy state. Polymer 2005;46:12577–12585. 45. Zhang YZ, Venugopal J, Huang ZM, Lim CT, Ramakrishna S. Crosslinking of the electrospun gelatin nanofibers. Polymer 2006; 47:2911–2911. 46. Kikic I, Vecchione F. Supercritical impregnation of polymers. Curr Opin Solid State Mater Sci 2003;7:399–405. 47. Freiberg S, Zhu XX. Polymer microspheres for controlled drug release. Int J Pharm 2004;282:1–18. 48. Varnell DF, Coleman MM. FTIR studies of polymer blends: V. Further observations on polyester-poly(vinyl chloride) blends. Polymer 1981;22:1324–1328. 49. Singh YP, Singh RP. Compatibility studies on solutions of polymer blends by viscometric and ultrasonic techniques. Eur Polym J 1983;19:535–541. 50. Hancock BC, York P, Rowe RC. The use of solubility parameters in pharmaceutical dosage form design. Int J Pharm 1997;148:1–21. 51. Stefanis E, Panayiotou C. Prediction of hansen solubility parameters with a new group-contribution method. Int J Thermophys 2008;29:568–585. 52. Shirahase T, Komatsu Y, Tominaga Y, Asai S, Sumita M. Miscibility and hydrolytic degradation in alkaline solution of poly(l-lactide) and poly(methyl methacrylate) blends. Polymer 2006;47: 4839–4844. 53. Li WJ, Cooper Jr JA, Mauck RL, Tuan RS. Fabrication and characterization of six electrospun poly(a-hydroxy ester)-based fibrous scaffolds for tissue engineering applications. Acta Biomater 2006; 2:377–385. 54. Bigi A, Cojazzi G, Panzavolta S, Rubini K, Roveri N. Mechanical and thermal properties of gelatin films at different degrees of glutaraldehyde crosslinking. Biomaterials 2001;22:763–768. 55. Wissinger RG, Paulaitis ME. Swelling and sorption in polymer– CO2 mixtures at elevated pressures. J Polym Sci Part B: Polym Phys 1987;25:2497–2510. 56. Ritger PL, Peppas NA. A simple equation for description of solute release II. Fickian and anomalous release from swellable devices. J Controlled Release 1987;5:37–42.
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
11
Department of Chemical and Biomolecular Engineering, Ohio State University, Columbus, Ohio 43210 Department of Biomedical Engineering, Ohio State University, Columbus, Ohio 43210 3 Materials Science and Engineering, Ohio State University, Columbus, Ohio 43210 2
Received 26 July 2013; revised 26 January 2014; accepted 18 February 2014 Published online 00 Month 2014 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.b.33132 Abstract: Blends of natural and synthetic polymers have received considerable attention as biomaterials due to the potential to optimize both mechanical and bioactive properties. Electrospinning of biocompatible polymers is an efficient method producing biomimetic topographies suited to various applications. In the ultimate application, electrospun scaffolds must also incorporate drug/protein delivery for effective cell growth and tissue repair. This study explored the suitability of a ternary Polymethylmethacrylate-Polycaprolactone-gelatin blend in the preparation of electrospun scaffolds for biomedical applications. Tuning the blend composition allows control over scaffold mechanical properties and degradation rate. Significant improvements were observed in the mechanical properties of the blend compared with the individual components. In order to study drug delivery potential, triblends were impregnated with the model compound Rhodamine-B
using sub/supercritical CO2 infusion under benign conditions. Results show significantly distinct release profiles of the impregnated dye from the triblends. Specific factors such as porosity, degradation rate, stress relaxation, dye-polymer interactions, play key roles in impregnation and release. Each polymer component of the triblends shows distinct behavior during impregnation and release process. This affects the aforementioned factors and the release profiles of the dye. Careful control over blend composition and infusion conditions creates the flexibility needed to produce biocompatible electrospun scaffolds for a variety of biomedical applications. C 2014 Wiley Periodicals, Inc. J Biomed Mater Res Part B: Appl BioV
mater 00B:000–000, 2014.
Key Words: electrospinning, polymer blends, impregnation, supercritical CO2
How to cite this article: Ramesh Munj H, Tyler Nelson M, Sadanand Karandikar P, Joseph Lannutti J, Lane Tomasko D. 2014. Biocompatible electrospun polymer blends for biomedical applications. J Biomed Mater Res Part B 2014:00B:000–000.
INTRODUCTION
Many challenges are faced in the development of tissue engineering scaffolds capable of controlled drug delivery, the desired cellular response, mechanical stability and tailored in vivo degradation.1 Electrospinning is a simple process to prepare polymeric fibrous scaffolds with high surface area and porosity required for drug delivery and tissue engineering.2 Electrospinning presents a threedimensional, porous microenvironment closely mimicking the extracellular matrix thus providing a suitable platform for the proliferation and differentiation of many different cell types.3 The efficacy of such scaffolds is based on several parameters such as cellular adhesion and proliferation in addition to structural integrity, void volume and fiber diameter.4,5 The design of the optimal scaffold continues to present a series of complex challenges due to these competing demands. Biocompatible polymers are widely used in tissue engineering and controlled drug release systems.6 Polycaprolactone (PCL) is an attractive option for tissue
engineered implants due to its mechanical properties and intermediate resistance to degradation.7 Electrospun PCLbased scaffolds have been used as tissue engineered scaffolds for in vitro development of blood vessels and nerve regeneration.8,9 Poly-methyl methacrylate (PMMA), on the other hand, is a synthetic, nondegradable biocompatible polymer historically useful in bone repair and hard tissue regeneration.10 Electrospun PMMA scaffolds have demonstrated good cellular adhesion, proliferation, and viability for in vitro and in vivo tissue engineering applications.11,12 Scaffold biofunctionality is essential to efficient cell adhesion and proliferation. A biofunctionalized scaffold has active sites recognizable by cells providing enhanced cell attachment.13 It has been shown that use of active molecules derived from natural extra cellular matrix proteins can supply these functionalized scaffolds. As an attempt to improve upon the biochemical properties of the two synthetic materials, gelatin has also been chosen for incorporation into these electrospun composite fiber scaffolds, creating improved cellular adhesion and proliferation.14,15
Correspondence to: D. Tomasko; e-mail: [email protected] Contract grant sponsor: National Science Foundation; contract grant number: EEC-0914790
C 2014 WILEY PERIODICALS, INC. V
1
Such precise compositional adjustments give control over blend mechanical and biological properties. Blends of natural polymers with synthetic polymers and these have proven to perform better in tissue engineering applications than their purely synthetic counterparts .16 Gelatin acts as a biological cue for cells and provides better biocompatibility than a synthetic polymer. It can also moderate the hydrophobic nature of some synthetic polymers to provide an added advantage.17 Supercritical fluids (SCF) have easily tunable physical properties (e.g. density, viscosity, etc.) useful in a wide range of applications.18 Carbon dioxide (CO2) is one of the most popular SCF due to its benign critical point conditions (7.38 MPa at 31.1 C). Recent studies have explored the potential of CO2 in sustainable and inexpensive processing of many polymers during foaming, sterilization, solvent extraction and infusion.19,20 CO2 has specific molecular interactions with many polymers to result in altered polymer properties that ease plasticization.21 In the presence of CO2, properties of many polymers (such as glass transition temperature Tg, etc.) are significantly lower than the pure form and can be further tuned by adjusting CO2 pressure and temperature.22 CO2 offers a green alternative way for the infusion of drugs and biomolecules into polymer substrates for tissue engineering. Various studies have already considered the plasticization and impregnation of PCL and PMMA in high pressure CO2.23,24 Although subcritical gaseous CO2 can reversibly swell PCL during impregnation, it results in fast release of the infused additive due its low solubility in dense CO2. Liquid CO2 can swell pure PCL reversibly without permanent deformation but has low diffusivity. Supercritical CO2 (SCCO2) has moderate diffusivity in polymers and adequate solubility for additives compared with that of dense gas or liquid CO2. Unfortunately, if pure PCL is subjected to SCCO2, it melts and loses mechanical integrity.25 Similarly, PMMA shows an irreversible change in the structure if processed under certain SCCO2 conditions.26,27 Preserving the initial structure of a scaffold (the porosity of foams and the fiber structure characteristic of fibrous scaffolds) is a factor essential to cell proliferation. In our recent work, we showed that if PCL is blended with gelatin the resulting blend exhibits distinct properties. When subjected to high pressure CO2, the blend can be swelled reversibly by dense, liquid and SCCO2.28 Our recent results have confirmed that pure PCL swells under high pressure CO2 whereas pure gelatin compresses due to dehydration in the presence of CO2. Simultaneous swelling of PCL and shrinking of gelatin stabilize the blend under SCCO2 without deformation. This discovery opens new avenue to explore CO2 assisted infusion into biocompatible blends containing gelatin for drug delivery and tissue engineering applications. The addition of PMMA to the existing PCL-Gelatin binary blend increases loading capacity of the blend due to amorphous nature of PMMA. Recently, several studies have investigated potential of polymeric ternary blends in the biomedical applications.29–32
2
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
In this article, we have explored the postinfusion properties of PMMA-PCL-gelatin ternary blends for biomedical applications. The major goal of this investigation is to understand the influence of blend composition and processing factors simultaneously on mechanical properties, impregnation and release from complex electrospun polymer matrices. Adjusting the blend ratio of the individual components allows control over mechanical properties and degradation rate. High pressure CO2 was used to assist in the infusion of Rhodamine B dye. Dye release from these scaffolds was studied over specific periods of time. Results indicate that PMMAPCL-gelatin ternary blends have promising properties for various biomedical applications. METHOD AND MATERIALS
Electrospinning Initially, 5% w/w polycaprolactone (PCL) (Mn 70,000– 90,000; Sigma-Aldrich, St. Louis, MO), 3% polymethyl methacrylate (PMMA) (Mn 350,000; Sigma-Aldrich, St. Louis, MO) and 6.7% w/w porcine gelatin type A (300 Bloom; SigmaAldrich, St. Louis, MO) solutions were prepared in 1,1,1,3,3,3-hexafluoro-2-propanol (HFP) (>99% purity; Sigma-Aldrich, St. Louis, MO) by stirring at room temperature (ffi25 C) for 24 h. Weight % concentrations were chosen based on electrospinning of pure polymer to achieve required fiber morphology. Different ternary blend compositions were obtained on a volume/volume basis by adjusting the amount of each component in the blended solution. After complete mixing of the solutions, Tri-PCL (PCL 50%, PMMA 25%, Gelatin 25%), Tri-PMMA (PCL 25%, PMMA 50%, Gelatin 25%), and Tri-Gelatin (PCL 25%, PMMA 25%, Gelatin 50%) blends were then poured individually into separate 20 mL syringes, fitted with a 20-gauge blunt tip needle, and electrospun using a DC high voltage power supply (Glassman High Voltage, Inc., High Bridge, NJ) at positive 20 kV, 20 cm needle-to-collector distance, 10 mL/hr flow rate, for 45 to 60 min; 8 cm 3 8 cm randomly oriented electrospun fiber mats of Tri-PCL, Tri-PMMA and Tri-Gelatin were produced on aluminum non-stick foil. The as-spun mats were then placed in a vacuum oven (99.5% purity, 200 Proof; Sigma-Aldrich, St. Louis, MO) solution was placed uniformly onto each of the scaffolds. These scaffolds were placed in chemical hood to dry for 2 h until all the ethanol was evaporated. Scaffolds were moved using tweezers in order to avoid their adhesion to the well plate surface. In the case of adsorption-only study, scaffolds were washed three times with pure ethanol after 30 min. For CO2 infusion, dried scaffolds were then placed into a high pressure stainless steel vessel. Temperature of the vessel was controlled using heat tape and monitored using a temperature sensor (Omega CSC 32, Stamford, CT). Bonedry CO2 (99.9% purity, Praxair, Columbus, OH) was allowed to flow in the pressure vessel and pressure was controlled with the syringe pump (500 D; Teledyne ISCO, Inc., Lincoln, NE). Equilibrium conditions were maintained for 2 h. Subcritical CO2 infusion was conducted at 6.20 MPa and 25 C whereas supercritical CO2 conditions were maintained at 8.27 MPa and 37 C. In either case, pressure was released by reversing syringe pump slowly (0.5–1 mL/min) over >12 h. Scaffolds were then washed with pure ethanol three times and dried in the chemical hood. Rhodamine B release was carried out in 24-well plates for all three impregnation conditions (adsorption, subcritical CO2 and supercritical CO2) by immersing each scaffold in the 2 mL of phosphate buffer saline (PBS) and keeping it in the incubator (Wiseven, Witeg Labortechnik GmbH, Germany) at 37 C. At each time point, two aliquots were taken from each sample and stored in 96-well plates. PBS was replaced with fresh PBS in each well after every time point. All 96-well plates containing aliquots were stored in the refrigerator. A UV-Vis 96-well plate spectrometer (Spectra Max 190 Absorbance UV-VIS plate reader, Sunnyvale, CA) was used to measure the absorbance of released Rhodamine B solutions at a wavelength of 535 nm. Spectrometer readings were then converted to concentration values (lg/mL) using a polynomial calibration curve produced by making serial dilutions of Rhodamine B in PBS. Weight loss study Similar to the infusion study, 13 mm diameter discs of TriPMMA, Tri-PCL and Tri-Gelatin fibers were weighed and placed in 24-well plates. Each scaffold was immersed in 2 mL of PBS and stored in the incubator (Wiseven, Witeg Labortechnik GmbH, Germany) at 37 C. Scaffolds were then removed after specific time interval to wash with distilled water in order to get rid of salts. These scaffolds were then dried overnight in a vacuum oven to remove all water before they were weighed again to analyze weight loss. Statistical analysis Statistical analysis was performed using JMP software. Each treatment was represented by at least three independent experiments. Comparisons between multiple treatments were made by fitting model for response against different variable and then comparing means for different pairs using Tuckey HSD. A p value of PMMA > gelatin. Among the different blends, a similar trend is observed as that of pure polymers i.e., the elongation of Tri-PCL >TriPMMA > Tri-Gelatin. Figure 6(C) shows ultimate tensile strength (UTS) measurements for pure components and ternary blends. There is notable enhancement in the UTS of Tri-PMMA and Tri-Gelatin blends when evaluated against their pure states. Although Tri-PCL blend shows some improvement over pure PCL it is not as high as PMMA and gelatin. Mechanical testing shows that a remarkable change in the properties of the blends compared with the individual components. PCL is a semi-crystalline polymer with lower modulus and higher elongation.42 Being above the glass
FIGURE 6. Comparison among different mechanical properties of different compositions of PMMA-PCL-gelatin blends against PCL, PMMA, and gelatin in their pure state A) modulus, B) elongation, C) ultimate tensile strength.
maintaining the fiber structure. However, for Tri-PMMA and Tri-Gelatin, PCL is a minor component. DSC and XRD confirmed that PCL is unable to form crystals in these blends.
FIGURE 7. Different factors affecting various blend components during impregnation (ethanol and high pressure CO2) and release (release media diffusion and degradation) processes.
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
7
FIGURE 8. Comparison of release profiles of Rhodamine B dye from ternary blends after different impregnation conditions.
transition temperature under ambient conditions, low stress is required to achieve large strain in PCL. PMMA is brittle material with lower elongation and high compressive strength at room temperature.43 The mechanical properties of gelatin are closely related to water content.44 As water
8
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
content is reduced, it changes the gel behavior to that of a rigid polymer. Electrospun gelatin fibers in solid state at ambient conditions contain about 10 to 15% moisture.45 Thus, pure gelatin fibers exhibit higher moduli and lower elongation under ambient conditions. Although blends
MUNJ ET AL.
ORIGINAL RESEARCH REPORT
prepared for this study are heterogeneous blends, each component contributes to the improvement of overall mechanical properties. Gelatin increases the modulus of the blends, whereas PCL enhances their elongation and UTS. Release of a solute from polymer blends is a complex process. Each polymer in the blend has a unique capacity for solute release based on polymer-solute interactions as well as infusion conditions.46 Apart from these factors, degradation and morphology of the polymer matrix in the release medium can play an essential role in the release.47 If we consider solute-polymer interactions for ternary blends, comparing solubility parameters for different polymers with Rhodamine B can provide a basis for comparison. The solubility parameters of PCL, PMMA and gelatin are 18.25, 18.2, and 24.3 MPa5 respectively.48–50 The solubility parameter of Rhodamine B estimated using group contribution theory is about 27.4 MPa0.5.51 Thus, gelatin should have higher compatibility for Rhodamine B as compared with PCL and PMMA. The blends have mixtures of semi-crystalline PCL, amorphous PMMA and gelatin all of which show different behavior when subjected to high pressure CO2. As discussed earlier, PMMA and PCL show significant increases in volume (swelling) when subjected to high pressure CO2. However, gelatin shows compression due to dehydration in the presence of CO2. Also these polymers differ in terms of their stability in aqueous release media. PMMA is nonbiodegradable polymer which cannot be degraded hydrolytically on a relevant time scale.52 PCL, a poly-hydroxy ester, shows slow hydrolysis of ester groups in PBS which proceeds via bulk degradation.53 Uncrosslinked gelatin is extremely unstable in aqueous media and shows rapid degradation due to hydrolytic or enzymatic degradation.54 Furthermore, relaxation dynamics of polymers play an important role in the impregnation and release. PMMA and gelatin electrospun fibers in their pure state show prominent reduction in the voidage of electrospun mat when subjected to 100% ethanol. This phenomenon is explained by the relaxation of stresses built up during the extremely rapid electrospinning process. As shown in the SEM analysis [Figure 1(E,F)], both Tri-PMMA and Tri-Gelatin blends show decrease in the void volume of an electrospun mat when subjected to Rhodamine B in ethanol. For Tri-gelatin fibers, a large degree of melting was observed, whereas, Tri-PMMA fibers show no melting but reduction in the void volume and change in fiber structure. Tri-PCL scaffolds show no change in structure after pure ethanol wetting. Figure 7 shows how different factors affect each component of the ternary blend during impregnation and release. Release from the blended electrospun fiber mats is shown in Figure 8. All release curves are expressed in terms of the cumulative mass of Rhodamine B per mass of fiber to compare different blends and different impregnation conditions. Figure 8(A) shows release from Tri-Gelatin fibers infused via simple adsorption, subcritical CO2 and supercritical CO2 conditions. Subcritical CO2 infusion releases significantly more dye with time than adsorption and supercritical CO2 infused samples that are not significantly different from each other. Release curves from Tri-PCL fibers [Figure 8(B)]
FIGURE 9. Percentage weight loss of electrospun ternary blends over time in PBS.
exhibit an obvious trend for the mass of released dye, where adsorption < subcritical CO2 < supercritical CO2. In the case of Tri-PMMA fibers [Figure 8(C)], all three infusion conditions show similar release profiles without notable change. Subcritical CO2 infused samples show a small but not prominent increase in the amount of released dye as compared with supercritical CO2 and adsorption which overlap with each other. Figure 8(D–F) show different blend compositions subjected to a specific impregnation condition. All three blend compositions show significantly different cumulative amounts released for each time point for adsorption [Figure 8(D)]. Tri-Gelatin shows the highest release followed by Tri-PMMA and finally Tri-PCL. Subcritical and supercritical conditions show Tri-PMMA and Tri-PCL are significantly different from Tri-Gelatin but not from each other [Figure 8(E,F)]. As discussed earlier, we see that release of a solute from polymer blends is a complex process and each polymer in the blend can show a unique response towards solute release rate. Among all three blends, Tri-Gelatin samples show highest amount of release of dye (per mg of fibers) for all impregnation conditions for the following reasons: (1) during impregnation large amount of Rhodamine B is retained in the TriGelatin scaffolds due to hydrophilic interactions between the dye and the gelatin; (2) although melting of scaffolds results in lower diffusion of release medium in the scaffolds, faster degradation of gelatin causes quick release of dye; (3) pure gelatin is compressed in the presence of high pressure CO2 preventing significant depth of infusion within the fiber. Nevertheless, subcritical CO2 shows smaller degree of compression for gelatin as compared with supercritical CO2 swelling. In the case of Tri-PCL fibers, a similar trend was observed as in the PCL-gelatin binary blend.28 Supercritical CO2 causes more swelling for PCL in contrast to subcritical CO2. Although gelatin is present in the blend, a compression effect is not dominant in the case of Tri-PCL fibers. All release curves of Tri-PMMA almost overlap for each of the impregnation conditions. Tri-PMMA fibers show no
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
9
TABLE I. Diffusion Coefficients (D) of Rhodamine B from Ternary Blends Infused Using Different Impregnation Conditions in a Static Release Medium (PBS).
Adsorption Subcritical CO2 Supercritical CO2
DTri-PMMA (mm2/sec)
DTri-PCL (mm2/sec)
DTri-Gelatin (mm2/sec)
4.8 3 10210 6.4 3 10210 7.7 3 10210
6.9 3 10210 9.92 3 10210 9.85 3 10210
9.3 3 10210 18.4 3 10210 20.6 3 10210
significant change in the dye released under different impregnation conditions. The most likely explanation for this abnormal behavior is restricted diffusion of release medium in the Tri-PMMA samples due to reduction in void volume. Also subcritical CO2 and supercritical CO2 show similar swelling for PMMA when subjected below 40 C.55 Figure 9 shows results from the weight loss study indicating Tri-Gelatin degrades rapidly in PBS followed by Tri-PCL and finally Tri-PMMA. If we model these scaffolds by considering a thin film, we can calculate Fickian diffusion coefficients of dye release using procedure explained by Ritger et al.56 Diffusion coefficients (Table I) indicate that all three compositions have different limiting factors in the release of dye. Tri-PMMA fibers show slowest release with lowest diffusion coefficients. This is the result of formation of fibrous membrane of Tri-PMMA scaffolds after stress relaxation. Considering Tri-PCL scaffolds can maintain the porosity during the release, higher diffusion coefficients are obtained. In the case of Tri-Gelatin fibers, rapid dissolution of gelatin releases dye quickly compared with Tri-PCL and Tri-PMMA scaffolds (Figure 9). CONCLUSION
In this study, standard electrospinning procedure was used to prepare ternary blend scaffolds of Tri-PCL (PCL 50%, PMMA 25%, gelatin 25%), Tri-PMMA (PCL 25%, PMMA 50%, gelatin 25%), and Tri-Gelatin (PCL 25%, PMMA 25%, GELATIN 50%) blends. We have explored a polymer blend with three components having unique characteristics. PCL boosts the elongation capacity of the blend and aids in maintaining structural properties during further processing of blend. Gelatin imparts higher modulus and the ability to achieve rapid release of the infused drug. Additionally, it makes the blend bioactive and allows efficient adhesion of proteins and cells. PMMA can contribute in compressive strength of scaffold and allows longer degradation periods. Stress relaxation of PMMA fibers lead to loss of interfiber porosity. However, it maintains the fiber structure and thus has high surface area. This can be used in dermal and transdermal patches for long-term drug delivery. Significant improvement was obtained in the mechanical properties of ternary blends as compared with individual components of the blend. DSC and XRD studies show noteworthy changes in the crystalline properties of PCL in the ternary blend. CO2 assisted impregnation of drugs and biomolecules in polymers is simple, benign and “green” alternative to current impregnation processes. A dye release study was carried out from all three blend compositions under various
10
BIOCOMPATIBLE ELECTROSPUN POLYMER BLENDS
impregnation conditions. Stress relaxation of PMMA and gelatin in the presence of ethanol during impregnation process changes morphological properties of Tri-PMMA and TriGelatin blend but Tri-PCL remains unchanged due to high crystalline content. Three major factors were found to have a significant impact on the dye release profiles from ternary blends: (1) effect of CO2; (2) degradation rate of scaffold; (3) porosity of scaffold. The Tri-PCL blended scaffolds display long degradation rates and maintain porosity of the scaffold in the release medium. The effect of CO2 was dominant in the case of Tri-PCL blends due to melting and recrystallization on PCL in the blend after CO2 treatment. For Tri-Gelatin scaffolds, degradation rate was found to be a key factor in rapid release of the dye from scaffolds. Also ethanol treatment shows significant melting of Tri-Gelatin fibers which increases the loading of the dye in the blend. In the case of Tri-PMMA, stress relaxation in the presence of ethanol reduces the porosity of the scaffold. Although PMMA shows an increase in swelling under high pressure CO2, release of Rhodamine B is limited by porosity of the scaffold and diffusion of release medium in the scaffold. In summary, all three polymers show unique contribution in altering the properties of the ternary blends. Composition of the ternary blend was found to be a vital component in deciding several properties of these electrospun scaffolds. Adjusting the composition of PMMA-PCL-gelatin ternary blends and tuning impregnation conditions, several tissue engineered scaffolds and controlled drug delivery systems can be designed for efficient biomedical applications. REFERENCES 1. Lee K, Silva EA, Mooney DJ. Growth factor delivery-based tissue engineering: general approaches and a review of recent developments. J R Soc Interface 2011;8:153–170. 2. Nisbet DR, Forsythe JS, Shen W, Finkelstein DI, Horne MK. Review paper: A review of the cellular response on electrospun nanofibers for tissue engineering. J Biomater Appl 2009;24:7–29. 3. Jagur-Grodzinski J. Polymers for tissue engineering, medical devices, and regenerative medicine. Concise general review of recent studies. Polym Adv Technol 2006;17:395–418. 4. Gloria A, De Santis R, Ambrosio L. Polymer-based composite scaffolds for tissue engineering. J Appl Biomater Biomechanics 2010;8:57–67. 5. Place ES, George JH, Williams CK, Stevens MM. Synthetic polymer scaffolds for tissue engineering. Chem Soc Rev 2009;38: 1139–1151. 6. Ulery BD, Nair LS, Laurencin CT. Biomedical applications of biodegradable polymers. J Polym Sci Part B: Polym Phys 2011;49: 832–864. 7. Dash TK, Badireenath Konkimalla V. Poly-¾-caprolactone based formulations for drug delivery and tissue engineering: A review. J Controlled Release 2012;158:15–33. 8. Vaz CM, Van Tuijl S, Bouten CVC, Baaijens FPT. Design of scaffolds for blood vessel tissue engineering using a multi-layering electrospinning technique. Acta Biomaterialia 2005;1:575–582. 9. Ghasemi-Mobarakeh L, Prabhakaran MP, Morshed M, NasrEsfahani MH, Ramakrishna S. Electrospun poly(e-caprolactone)/ gelatin nanofibrous scaffolds for nerve tissue engineering. Biomater 2008;29:4532–4539. 10. Mano JF, Sousa RA, Boesel LF, Neves NM, Reis RL. Bioinert, biodegradable and injectable polymeric matrix composites for hard tissue replacement: State of the art and recent developments. Compos Sci Technol 2004;64:789–817. 11. Zhang F, Sun F, Van Kan JA, Shao PG, Zheng Z, Ge RW, Watt F. Measurement of cell motility on proton beam micromachined 3D
MUNJ ET AL.
ORIGINAL RESEARCH REPORT
12. 13.
14.
15. 16.
17.
18.
19.
20.
21.
22.
23.
24.
25.
26.
27.
28.
29.
30.
31.
32.
33.
scaffolds. Nucl Instrum Methods Phys Res Sect B 2005;231:413– 418. Srouji S, Kizhner T, Livne E. 3D scaffolds for bone marrow stem cell support in bone repair. Future Med 2006;1:519–528. Liu X, Holzwarth JM, Ma PX. Functionalized synthetic biodegradable polymer scaffolds for tissue engineering. Macromol Biosci 2012;12:911–919. Liu X, Won Y, Ma PX. Porogen-induced surface modification of nano-fibrous poly(L-lactic acid) scaffolds for tissue engineering. Biomaterials 2006;27:3980–3987. Liu X, Won Y, Ma PX. Surface modification of interconnected porous scaffolds. J Biomed Mater Res A 2005;74:84–91. Sionkowska A. Current research on the blends of natural and synthetic polymers as new biomaterials: Review. Prog Polym Sci 2011;36.9:1254–1276. Pant HR, Kim CS. Electrospun gelatine/nylon-6 composite nanofibers for biomedical applications. Polym Int 2012;62:1008– 1013. Duarte ARC, Mano JF, Reis RL. Supercritical fluids in biomedical and tissue engineering applications: A review. Int Mater Rev 2009;54:214–222. Bhamidipati M, Scurto AM, Detamore MS. The future of carbon dioxide for polymer processing in tissue engineering. Tissue Eng Part B: Rev 2013;19:221–232. Duarte ARC, Santo VE, Alves A, Silva SS, Moreira-Silva J, Silva TH, Marques AP, Sousa RA, Gomes ME, Mano JF, Reis RL. Unleashing the potential of supercritical fluids for polymer processing in tissue engineering and regenerative medicine. J Supercrit Fluids 2013;79:177–185. Yanhui Y, Teja AS. Quantification of specific interactions between CO2 and the carbonyl group in polymers via ATR-FTIR measurements. J Supercrit Fluids 2011;56:208–212. Tomasko DL, Li H, Liu D, Han X, Wingert MJ, Lee LJ, Koelling KW. A Review of CO2 Applications in the processing of polymers. Ind Eng Chem Res 2003;42:6431–6456. Roshan Y, Mammucari R, Foster NR. Impregnation of ibuprofen into polycaprolactone using supercritical carbon dioxide. J Phys: Conf Ser 2010;215:012087. Yazan HA, Grant CS. Ibuprofen impregnation into submicron polymeric films in supercritical carbon dioxide. J Supercrit Fluids 2012;71:127–135. Ayodeji O, Graham E, Kniss D, Lannutti JJ, Tomasko DL. Carbon dioxide impregnation of electrospun polycaprolactone fibers. J Supercrit Fluids 2007;41:173–178. Kamiya Y, Mizoguchi K, Terada K, Fujiwara Y, Wang JS. CO2 sorption and dilation of poly(methyl methacrylate). Macromolecules 1998;31:472–478. Di Noto V, Vezzu K, Giffin GA, Conti F, Bertucco A. Effect of High Pressure CO2 on the Structure of PMMA: A FT-IR Study. J Phys Chem B 2011;115:13519–13525. Nelson MT, Munj HR, Tomasko DL, Lannutti JJ. Carbon dioxide infusion of composite electrospun fibers for tissue engineering. J Supercrit Fluids 2012;70:90–99. Kumari S, Basha SK, Sudha PN. Physicochemical and morphological evaluation of chitosan/poly (vinyl alcohol)/methylcellulose chemically cross-linked ternary blends. Polym Bull 2012;68:1387–1393. Liu H, Song W, Chen F, Guo L, Zhang J. Interaction of microstructure and interfacial adhesion on impact performance of polylactide (PLA) ternary blends. Macromolecules 2011;44: 1513–1522. Grande R, Carvalho AJ. Compatible ternary blends of chitosan/ poly (vinyl alcohol)/poly (lactic acid) produced by oil-in-water emulsion processing. Biomacromolecules 2011;12:907–914. Mudigoudra BS, Masti SP, Chougale RB. Thermal behavior of poly (vinyl alcohol)/poly (vinyl pyrrolidone)/chitosan ternary polymer blend films. Res J Recent Sci 2012;1:83–86. Gu SY, Wang ZM, Ren J, Zhang CY. Electrospinning of gelatin and gelatin/poly(L-lactide) blend and its characteristics for wound dressing. Mater Sci Eng C 2009;29:1822–1828.
34. Kim CH, Cho KY, Choi EJ, Park JK. Effect of P(lLA-co-eCL) on the compatibility and crystallization behavior of PCL/PLLA blends. J Appl Polym Sci 2000;77:226–231. 35. Domanska U, Pobudkowska A, Pelczarska A, Gierycz P. pKa and solubility of drugs in water, ethanol, and 1-octanol. J Phys Chem B 2009;113:8941–8947. 36. Jopling DW. Stress relaxation studies of chemically crosslinked gelatin films. J Polym Sci Part A: Gen Pap 1965;3:513–526. 37. Sarti GC, Gostoli C, Masoni S. Diffusion of alcohols and relaxation in poly(methyl methacrylate): Effect of thermal history. J Membr Sci 1983;15:181–192. 38. Matthews RG, Ajji A, Dumoulin MM, Prud’homme RE. The effects of stress relaxation on the structure and orientation of tensile drawn poly(ethylene terephthalate). Polymer 2000;41: 7139–7145. 39. Shen Z, McHugh MA, Xu J, Belardi J, Kilic S, Mesiano A, Bane S, Karnikas C, Beckman E, Enick R. CO2-solubility of oligomers and polymers that contain the carbonyl group. Polymer 2003;44:1491– 1498. 40. Shieh YT, Yang HS. Morphological changes of polycaprolactone with high-pressure CO2 treatment. J Supercrit Fluids 2005;33:183– 192. 41. Liu D, Li H, Noon MS, Tomasko DL. CO2-Induced PMMA swelling and multiple thermodynamic property analysis using sanchez lacombe EOS. Macromolecules 2005;38:4416–4424. 42. Tan EPS, Ng SY, Lim CT. Tensile testing of a single ultrafine polymeric fiber. Biomaterials 2005;26:1453–1456. 43. Mousa WF, Kobayashi M, Shinzato S, Kamimura M, Neo M, Yoshihara S, Nakamura T. Biological and mechanical properties of PMMA-based bioactive bone cements. Biomaterials 2000;21: 2137–2146. 44. Yakimets I, Wellner N, Smith AC, Wilson RH, Farhat I, Mitchell J. Mechanical properties with respect to water content of gelatin films in glassy state. Polymer 2005;46:12577–12585. 45. Zhang YZ, Venugopal J, Huang ZM, Lim CT, Ramakrishna S. Crosslinking of the electrospun gelatin nanofibers. Polymer 2006; 47:2911–2911. 46. Kikic I, Vecchione F. Supercritical impregnation of polymers. Curr Opin Solid State Mater Sci 2003;7:399–405. 47. Freiberg S, Zhu XX. Polymer microspheres for controlled drug release. Int J Pharm 2004;282:1–18. 48. Varnell DF, Coleman MM. FTIR studies of polymer blends: V. Further observations on polyester-poly(vinyl chloride) blends. Polymer 1981;22:1324–1328. 49. Singh YP, Singh RP. Compatibility studies on solutions of polymer blends by viscometric and ultrasonic techniques. Eur Polym J 1983;19:535–541. 50. Hancock BC, York P, Rowe RC. The use of solubility parameters in pharmaceutical dosage form design. Int J Pharm 1997;148:1–21. 51. Stefanis E, Panayiotou C. Prediction of hansen solubility parameters with a new group-contribution method. Int J Thermophys 2008;29:568–585. 52. Shirahase T, Komatsu Y, Tominaga Y, Asai S, Sumita M. Miscibility and hydrolytic degradation in alkaline solution of poly(l-lactide) and poly(methyl methacrylate) blends. Polymer 2006;47: 4839–4844. 53. Li WJ, Cooper Jr JA, Mauck RL, Tuan RS. Fabrication and characterization of six electrospun poly(a-hydroxy ester)-based fibrous scaffolds for tissue engineering applications. Acta Biomater 2006; 2:377–385. 54. Bigi A, Cojazzi G, Panzavolta S, Rubini K, Roveri N. Mechanical and thermal properties of gelatin films at different degrees of glutaraldehyde crosslinking. Biomaterials 2001;22:763–768. 55. Wissinger RG, Paulaitis ME. Swelling and sorption in polymer– CO2 mixtures at elevated pressures. J Polym Sci Part B: Polym Phys 1987;25:2497–2510. 56. Ritger PL, Peppas NA. A simple equation for description of solute release II. Fickian and anomalous release from swellable devices. J Controlled Release 1987;5:37–42.
JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2014 VOL 00B, ISSUE 00
11
