Fabrication and characterization of six electrospun poly(α-hydroxy ester)-based fibrous scaffolds for tissue engineering applications

Acta Biomaterialia 2 (2006) 377–385 www.actamat-journals.com

Fabrication and characterization of six electrospun poly(-hydroxy ester)-based Wbrous scaVolds for tissue engineering applications Wan-Ju Li a, James A. Cooper Jr. b, Robert L. Mauck a, Rocky S. Tuan a

a,¤

Cartilage Biology and Orthopaedics Branch, National Institute of Arthritis and Musculoskeletal and Skin Diseases, National Institutes of Health, Room 1523, Bldg 50, MSC 8022, Department of Health and Human Services, Bethesda, MD 20892, United States b Polymers Division, National Institute of Standards and Technology, Gaithersburg, MD 20899, United States Received 7 September 2005; received in revised form 16 February 2006; accepted 22 February 2006

Abstract The most common synthetic biodegradable polymers being investigated for tissue engineering applications are FDA approved, clinically used poly(-hydroxy esters). To better assess the applicability of the electrospinning technology for scaVold fabrication, six commonly used poly(-hydroxy esters) were used to prepare electrospun Wbrous scaVolds, and their physical and biological properties were also characterized. Our results suggest that speciWc, optimized fabrication parameters are required for each polymer to produce scaVolds that consist of uniform structures morphologically similar to native extracellular matrix. Scanning electron microscopy (SEM) revealed a highly porous, three-dimensional structure for all scaVolds, with average Wber diameter ranging from 300 nm to 1.5 m, depending on the polymer type used. The poly(glycolic acid) (PGA) and poly(D,L-lactic-co-glycolic acid 50:50) (PLGA5050) Wbrous structures were mechanically stiVest, whereas the poly(L-lactic acid) (PLLA) and poly(-caprolactone) (PCL) scaVolds were most compliant. Upon incubation in physiological solution, severe structural destruction due to polymer degradation was found in the PGA, poly(D,L-lactic acid) (PDLLA), PLGA5050, and poly(D,L-lactic-co-glycolic acid 85:15) (PLGA8515) Wbrous scaVolds, whereas PLLA and PCL Wbrous scaVolds maintained a robust scaVold structure during the same time period, based on macroscopic and SEM observations. In addition, PLLA scaVolds supported the highest rate of proliferation of seeded cells (chondrocytes and mesenchymal stem cells) than other polymeric scaVolds. Our Wndings showed that PLLA and PCL based Wbrous scaVolds exhibited the most optimal structural integrity and supported desirable cellular response in culture, suggesting that such scaVolds may be promising candidate biomaterials for tissue engineering applications. © 2006 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. Keywords: Electrospinning; NanoWber; Fiber; Mesenchymal stem cell; Chondrocyte

1. Introduction A tissue engineered scaVold is a three-dimensional, biomaterial matrix that may be used as a vehicle to deliver therapeutic cells or bioactive factors to the defect region or as a space Wller to recruit surrounding cells into the scaVold for the tissue repair process. The goal of scaVold design for tissue engineering is to produce a biomaterial

*

Corresponding author. Tel.: +1 301 451 6854; fax: +1 301 435 8017. E-mail address: [email protected] (R.S. Tuan).

matrix that can replace the natural extracellular matrix (ECM), until the seeded cells can produce a new natural matrix and regenerate the desired tissue structure. Critical parameters for tissue engineering scaVold thus include biocompatibility, biodegradability, optimal mechanical strength, and ability to regulate appropriate cellular activities [1]. A large number of polymeric biomaterials, including non-biodegradable and biodegradable polymers, have been tested and analyzed for tissue engineering applications [2]. Since non-biodegradable polymers would interfere with tissue turnover and remodeling, the current trend is to use biodegradable polymers in tissue engineering,

1742-7061/$ - see front matter © 2006 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. doi:10.1016/j.actbio.2006.02.005

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although the non-biodegradable polymers have the advantage that their properties, both chemical and mechanical, are less aVected by the cellular and tissue milieu. On the other hand, polymer biodegradation via the combined eVect of enzymatic and hydrolytic activities generates space within the scaVold to allow for cell proliferation and the deposition of newly synthesized ECM [3]. Ideally, optimal tissue regeneration occurs upon complete biodegradation of the polymeric matrix followed by restoration of biological functions. Poly(-hydroxy esters), including poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and their copolymer poly(lactic-co-glycolic acid) (PLGA), are the most commonly used synthetic polymers in tissue engineering, because of their well characterized biodegradable property and the fact that they are approved by the United States Food and Drug Administration for clinical use. PGA is soluble in highly Xuorinated organic solvents, due to its high crystallinity [4]. With the addition of a methyl group, PLA is more hydrophobic than PGA but soluble in common organic solvents [4]. The three isomers of PLA [D(¡), L(+), and D, L], that diVer based on the position of a methyl group in the lactic acid monomer, exhibit distinct properties. To expand the spectrum of applications, PGA and PLA have been copolymerized in diVerent ratios to form new polymers, PLGA. The PLGA copolymer is amorphous because the PGA and PLA polymer chains are not packed tightly [4]. Another family member in the poly(-hydroxy ester) group is poly(-caprolactone) (PCL), a semicrystalline, hydrophobic, biodegradable polymer. Compared to other polyester family members such as PLA, PGA, and PLGA, PCL has been used less frequently as a material for fabricating biomaterial scaVolds, mainly because of concern over its slower degradation kinetics. However, PCL may be suitable for applications such as long-term drug delivery [5], and in addition, its mechanical properties and degradation proWle can be modiWed by blending or copolymerizing PCL with other polyesters [4]. Electrospinning has been used to fabricate tissue engineered scaVolds comprising non-woven, three-dimensional, porous, and nanoscale Wber-based matrix. The characteristics of Wbrous scaVolds, such as high surface area to volume ratio and similar structural morphology to the Wbrillar ECM, suggest they may serve as eVective tissue engineering scaVolds [6–8]. In this study, in order to better evaluate their applicability for tissue engineering, six commonly used poly(-hydroxy esters) were electrospun into Wbrous scaVolds, including PGA, poly(L-lactic acid) (PLLA), poly(D,L-lactic acid) (PDLLA), poly(D,L-lactic-co-glycolic acid 50:50) (PLGA5050), poly(D,L-lactic-co-glycolic acid 85:15) (PLGA8515), and PCL, and their physical properties were characterized. In addition, two candidate cell types applicable for skeletal tissue engineering, chondrocytes and multipotential mesenchymal stem cells, were used to evaluate their biological responses upon seeding into the Wbrous scaVolds.

Table 1 Optimized fabrication parameters used in electrospinning process polymeric Wbersa Polymer

Solvent

Voltage (in kV)

PGA PDLLA PLLA PLGA5050 PLGA8515 PCL

40 mL HexaXuoro-2-propanol 5.7 mL THF + 5.7 mL DMF 25 mL chloroform + 2.5 mL DMF 5.7 mL THF + 5.7 mL DMF 6.7 mL THF + 6.7 mL DMF 14 mL THF + 14 mL DMF

15 10 16 12 15 12

a Poly(glycolic acid) (PGA); poly(L-lactic acid) (PLLA); poly(D,L-lactic acid) (PDLLA); poly(D,L-lactic-co-glycolic acid 50:50) (PLGA5050); poly(D,L-lactic-co-glycolic acid 85:15) (PLGA8515); poly(-caprolactone) (PCL); tetrahydrofuran (THF); N,N-dimethylformamide (DMF).

2. Materials and methods 2.1. Polymers and reagents1 The poly(-hydroxy ester) polymers and reagents in this study were obtained from the following sources: PGA (MW D 150,000), PLLA (MW D 50,000), poly(2-hydroxyethyl methacrylate) (poly-HEMA), Polysciences (Warrington, PA); PDLLA (MW D 109,000), PLGA5050 (MW D 79,000), PLGA8515 (MW D 123,000), Alkermes (Cincinnati, OH); PCL (MW D 80,000), Aldrich (Milwaukee, WI); 1,-1,-1,-3,-3,-3-hexaXuoro-2-propanol, Sigma (St. Louis, Mo); tetrahydrofuran (THF), N,N-dimethylformamide (DMF), chloroform, Fisher ScientiWc (Pittsburgh, PA); Dulbecco’s modiWed Eagle’s medium (DMEM, high glucose), Eagle’s minimum essential medium (MEM) vitamin solution, penicillin–streptomycin, phosphate-buVered saline (PBS), 0.25% trypsin, Gibco BRL Life Technologies (Grand Island, NY); fetal bovine serum (FBS), L-ascorbic acid 2-phosphate, glutaraldehyde, Sigma (St. Louis, MO); Hanks’ Balanced Salt Solution (HBSS), BioSource International (Camarillo, CA); CellTiter 96™ Aqueous One Solution Cell Proliferation Assay, Promega (Madison, WI); RPMI 1640, Invitrogen (Carlsbad, CA). 2.2. Electrospun Wbrous matrices Each polymer solution was prepared by dissolving 4 g of polymer in an optimal amount of organic solvent mixture and mixed well by vortexing overnight. The solvent type used and the concentration of each polymer solution prepared for electrospinning are listed in Table 1. For the electrospinning process, used in our previous studies [6], the polymer solution was placed in a 10 mL glass syringe Wtted with a 10 cm, 18-G blunt tipped needle. The syringe was Wxed vertically at the support in a custom-designed 1 Certain commercial materials and equipment are identiWed in this paper in order to specify adequately the experimental procedure. In no case does such identiWcation imply recommendation by the National Institute of Standards and Technology nor does it imply that the material or equipment identiWed is necessarily the best available for this purpose.

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electrospinning apparatus. An optimal voltage (listed in Table 1), speciWc for each polymer type, was provided by a high DC voltage power supply (Gamma High Voltage Research, Ormond Beach, FL), with the electric Weld applied at a distance of 20 cm between the copper collecting plate (cathode) and the needle tip (anode). To eVectively collect electrospun Wbers, a square of Wber-collecting area (12 cm £ 12 cm) on the copper plate was created by masking the rest of the plate with non-conductive plastic. The square was covered with aluminum foil for easy removal of the Wbrous matrix from the plate. The polymer solution was drawn from the syringe forming a pendant drop at the tip of the needle, and then a positively charged jet ejected from the drop sprayed onto the negatively charged target. The Wbrous matrix was formed on the collecting foil to yield scaVolds of approximately 1 mm thickness. 2.3. Physical characterization 2.3.1. Scanning electron microscopy (SEM) Electrospun Wbrous structures were sputter coated with gold (Desk II; Denton Vacuum, Moorestown, NJ) and their morphologies were observed by SEM (JEOL JSM5300, Japan) at an accelerating voltage of 20 kV. 2.3.2. Tensile property The Wbrous structures were cut into 0.5 cm £ 4 cm rectangular shapes and tested for tensile property with an uniaxial tensile tester. A custom mechanical testing device [9], with a 5 lb load cell (Model 31, Sensotec, Columbus, OH) Wtted with tensile grips (GF-53, Ametek, Brooklyn, NY) was used. The grips were modiWed to incorporate 80-grit sandpaper aYxed with heavy-duty double-sided tape to securely Wx polymer meshes during tensile testing. All constructs were weighed and their dimensions measured with a digital micrometer prior to testing. After mounting, the gauge length of samples was measured and a small tare load applied (»0.05 lbs) to ensure proper seating. Ten sinusoidal pre-conditioning cycles were then carried out to 1% of the gauge length at a strain rate of 0.1%/s. After pre-conditioning, a constant strain of 0.1%/s was applied until sample failure or 50% strain was achieved. In cases where samples did not fail, the non-recoverable deformation was assessed by releasing the applied deformation until the measured load became negligible at equilibrium. The Young’s modulus of samples in tension was calculated from the slope of the stress–strain curve in the linear region (i.e., below the yield stress) and the initial sample geometry. Yield stress and yield strain for each sample was determined from the intersection of the experimental data with a line parallel to linear region of the stress strain curve and oVset by +0.2% strain. Four samples were tested for each type of Wbrous scaVold. 2.3.3. Degradation evaluation Eight-mm circular specimens were cut and bathed in PBS at 37 °C for 42 d. The bathing solution was changed

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every 4 d. At Days 1, 3, 7, 14, 21, and 42, three samples of each type of scaVold were removed, vacuum-dried for 48 h, and degradation assessed macroscopically by SEM. 2.4. Biological characterization 2.4.1. Cell types 2.4.1.1. Articular chondrocytes. Primary bovine articular chondrocytes were isolated from calf knee joints [10,11] and cultured in 150 cm2 tissue culture Xasks (Corning Glass Works, Corning, NY) containing a chondrocyte growth medium and maintained at 37 °C in a humidiWed, 5% CO2 atmosphere. The chondrocyte growth medium was composed of DMEM, 10% (v/v) FBS, 1% (v/v) MEM vitamin solution, 50 g/mL L-ascorbic acid 2-phosphate, and antibiotics (50 g/mL of streptomycin, 50 IU of penicillin/mL). Cell culture medium was replaced every 3 d. Cells obtained at passages 1 or 2 were used in this study. 2.4.1.2. Mesenchymal stem cells. With approval from the Institutional Review Board of George Washington University, human mesenchymal stem cells (hMSCs) were isolated from bone marrow obtained from the femoral heads of patients undergoing total hip arthroplasty, and processed as previously described [12–14]. BrieXy, whole bone marrow was curetted from the exposed cutting plane of the femoral neck, washed extensively in DMEM, separated from contaminating trabecular bone fragments and other tissues using a 20-G needle attached to a 10 mL syringe, and cultured in DMEM, supplemented with 10% FBS from selected lots [15], and antibiotics at a density of 4 £ 105 cells/ cm2. Medium changes were carried out every 3 d. SubconXuent cell monolayers were dissociated using 0.25% trypsin and either passaged or utilized directly for study. 2.4.2. Cell–matrix interaction analysis Each of the six electrospun Wbrous mats was cut into 1 cm £ 1 cm square shapes, and both sides of the scaVold were sterilized by ultraviolet irradiation in a laminar Xow hood for 30 min. To provide a hydrophilic surface conducive for eYcient cell attachment, scaVolds were hydrated in a graded series of ethanol, rinsed with pure water, and preserved in HBSS. Bovine chondrocytes grown in 150 cm2 cell culture Xasks were trypsinized and counted. An aliquot of 10,000 cells was seeded onto the surface of a pre-wetted scaVold placed in 24-well culture plates (Corning Glass Works, Corning, NY) pre-coated with 0.3% poly-HEMA to prevent cell attachment to tissue culture polystyrene. Cellseeded scaVolds were incubated at 37 °C for 4 h to allow cells to diVuse into and adhere to the scaVold before the addition of 2 mL of chondrocyte growth medium into each well. During the 4 h of incubation, 20 L of chondrocyte growth medium was applied to each cellular scaVold every 30 min to maintain suYcient hydrolation of the constructs. Cellular constructs were harvested at Days 1, 3, and 7, Wxed with 4% glutaraldehyde for 1 h at room temperature, dehydrated through a series of graded alcohol solutions, and

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then air-dried overnight. The constructs were then sputter coated and observed by SEM. 2.4.3. Cell proliferation analysis The MTS [3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium] assay (CellTiter 96™ Aqueous One Solution Cell Proliferation Assay) was used to determine cell proliferation. Chondrocytes and hMSCs were seeded at diVerent seeding densities (10,000 chondrocytes and 50,000 MSCs) onto Wbrous scaVolds (1 cm £ 1 cm £ 1 mm) in 24 well culture plates, precoated with poly-HEMA, to evaluate cell proliferation. At Days 1, 7, 14, 21, and 28, culture medium was removed, and replaced with 400 L phenol red-free RPMI 1640 plus 80 L of CellTiter 96 reagent and incubated for 1 h. A 120 L aliquot of incubated medium was transferred to a 96-well plate, and A490 determined using a multiwell plate reader (VICTOR2 V; Perkin Elmer, Boston, MA). 2.5. Statistical analysis An estimate of the standard uncertainty values were expressed as mean § standard deviation. Statistical diVerences were determined by Student’s two-tailed t-test. 3. Results 3.1. Morphology of electrospun Wbrous scaVolds SEM revealed that all six electrospun matrices were composed of uniform, randomly oriented Wbers (Fig. 1(A)– (F)) whose diameters ranged from 300 nm to 1.5 m, with PGA Wbers (Fig. 1(A)) the Wnest, and PLLA Wbers (Fig. 1(E)) the largest. Three-dimensional, interconnected pores formed between the Wbers and were distributed throughout the scaVold matrix. 3.2. Tensile properties of electrospun Wbrous scaVolds Tensile testing of the six Wbrous scaVolds revealed clear diVerences (Fig. 2(A)). For each polymeric matrix, elongation of the Wbrous scaVold resulted in a linear increase in stress followed by a sharp and reproducible yield point. Calculation of the Young’s modulus of these samples showed that PLGA5050 had the highest modulus (»144 MPa), followed by PGA (»138 MPa), PLGA8515 (»114 MPa), and PDLLA (»70 MPa). PCL and PLLA scaVolds had the lowest Young’s modulus, each on the order of 8.5 MPa (Fig. 2(B)). Yield stresses (Fig. 2(C)) were also dependent on Wber polymer composition, with those scaVolds having the highest Young’s modulus generally having the higher yield stress (with yield strains all »1.8– 2.3%). Interestingly, the properties of PCL Wbrous scaVold diVered from this trend, with a relatively low Young’s modulus, but a higher yield stress due to its much increased yield strain (»7.5%) (Fig. 2(D)). Over the application of 50% strain in these studies, PLLA Wbrous scaVolds repeat-

Fig. 1. SEM micrograph of electrospun poly(-hydroxy ester) Wbrous scaVolds composed of randomly oriented ultra-Wne Wbers. (A) PGA, (B) PLGA5050, (C) PDLLA, (D) PLGA8515, (E) PLLA, and (F) PCL. Bar: 10 m.

edly failed in the vicinity of 20–25% strain. One of four PLGA5050 scaVolds failed at 38% strain. Of those samples that did not fail completely, the percentage of non-recoverable elongation ranged from 48% for PCL to 60–70% for PDLLA, PLGA5050 and PLGA8515. 3.3. Degradation of Wbrous scaVolds Degradation proWles of the Wbrous scaVolds were assessed macroscopically by SEM. PGA Wbrous scaVolds were able to maintain the original size in the Wrst 14 d and obvious size shrinkage was found at Day 21 (Fig. 3(A)). By Day 42, PGA Wbrous scaVolds degraded completely into polymer particles. PDLLA (Fig. 3(B)), PLGA5050 (Fig. 3(C)), and PLGA8515 (Fig. 3(D)) Wbrous scaVolds all showed signiWcant degradation, resulting in severe structural shrinkage. PLGA5050 Wbrous scaVolds shrank almost 90% in the Wrst 3 d, followed by PLGA8515 (75%), and PDLLA (60%). The structural shrinkage resulting from polymer degradation was conWrmed by SEM analysis. In contrast, PLLA (Fig. 3(E)) and PCL (Fig. 3(F)) Wbrous matrices were both well maintained during the 42 d of degradation evaluation. The structural architecture of these two Wbrous scaVolds was also preserved. SEM revealed that in the PGA (Fig. 4(A)–(C)), PLLA (Fig. 4(G)–(I)), and PCL (Fig. 4(P)–(R)) matrices, both Wbrous and pore structures remained during polymer degradation. In the PLGA8515 matrix, the Wber appeared swollen but pores were still detectable after 1 d of degradation (Fig. 4(N)). However, by Day 3, more degradation

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Fig. 2. (A) Tensile stress–strain curve of mechanical testing of electrospun Wbrous scaVolds. PLGA5050 (black line) Wbrous scaVolds had the highest yield stress, followed by PGA (orange line), PLGA8515 (green line), PDLLA (red line), PCL (blue line), and PLLA (gray line). (B) Calculated Young’s modulus. The results indicate that the PLGA5050 and PGA scaVolds are stiVest, whereas the PLLA and PCL scaVolds are softest. (C) Yield stress. PLGA5050 and PGA had the highest yield stress values. (D) Yield strain. PCL had the highest yield strain value and the other Wve had similar yield strain values. All values are an estimate of the standard uncertainty expressed as mean § S.D. ¤: p < 0.05, n D 4.

lytic degradation of the polymer, resulting in severe structural destruction. However, the nature of the degradation taking place in these two polymers diVered slightly. PDLLA degradation resulted in Wber swelling (Fig. 4(E)), whereas PLGA5050 degradation resulted in the fusion of Wbers (Fig. 4(K)). Both modes of degradation resulted in the destruction of pores (Fig. 4(F) and (L)). 3.4. Cell–matrix interaction

Fig. 3. Degradation of poly(-hydroxy ester) Wbrous scaVolds maintained in PBS up to 42 d. (A) PGA scaVolds maintained their shape and size up to 14 d, shrank at Day 21, and disintegrated at Day 42. (B) PDLLA scaVolds shrank after 3 d. (C) PLGA5050 scaVolds showed the most severe shrinkage at Day 3 as compared to other scaVolds. (D) PLGA8515 scaVolds also shrank after 3 d. (E) PLLA scaVolds maintained their structural shape and size during the 42 d. (F) PCL scaVolds behaved similarly to PLLA scaVolds and maintained their structure during the 42 d of incubation.

caused the pores to disappear (Fig. 4(O)). SEM analysis revealed that PDLLA (Fig. 4(E) and (F)) and PLGA5050 (Fig. 4(K) and (L)) Wbers were most vulnerable to hydro-

Cell–matrix interactions between cells and Wbers were studied in vitro by seeding bovine chondrocytes on six diVerent electrospun Wbrous scaVolds for 7 d. The extent and nature of cell–matrix interaction was highly inXuenced by the degradation of the Wbrous scaVold. At Day 3, rapid polymer degradation resulted in the destruction of the three-dimensional, porous structure, and closure of the interconnecting pores particularly in the case of PDLLA, PLGA5050, and PLGA8515 scaVolds (Fig. 5(C), (G) and (I)), thus precluding cell penetration. On the other hand, PGA, PLLA, and PCL Wbrous scaVolds maintained their three-dimensional, porous structure, and cells adhered to the Wbers, and started to migrate through the pores and grew within layers of Wbrous network (Fig. 5(A), (E) and (K)). By Day 7, cells merely covered the surface of PGA, PDLLA, PLGA5050, and PLGA8515 scaVolds (Fig. 5(B), (D), (H) and (J)). On the other hand, chondrocytes interacted and integrated well with the surrounding Wbers in the PLLA and PCL scaVolds (Fig. 5(E), (F), (K) and (L)). Cell growth also appeared to be guided by the Wber architecture, extending along the direction of Wber orientation, giving rise to a three-dimensional and multi-cellular network

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Fig. 4. SEM analysis of the degradation of electrospun Wbrous scaVolds. PGA (A–C), PLLA (G–I), and PCL (P–R) scaVolds maintained their Wbrous and porous structure. PDLLA (D–F) and PLGA5050 (J–L) scaVolds degraded after 1 d and all the pores in both structures closed after 3 d. PLGA85151 (M–O) scaVolds showed signiWcant degradation at Day 3. (A, D, G, J, M, P), without PBS; (B, E, H, K, N, Q), 1 d in PBS; and (C, F, I, L, O, R), 3 d in PBS. Bar: 1 m.

guided by the architecture of the Wbrous scaVold (Fig. 5(F) and (L)). 3.5. Cellular proliferation MTS analysis showed that chondrocytes cultured in all polymeric scaVolds exhibited a similar proliferation pattern, i.e., cells proliferated rapidly in the Wrst 21 d then reached a growth plateau (Fig. 6(A)). Although overall cell number increased in all scaVolds during the culture, PLLA scaVold supported signiWcantly higher cell proliferation compared to other scaVolds. In contrast, undiVerentiated hMSCs showed three distinct proliferation patterns in various scaVold cultures. In PGA, PDLLA, and PCL scaVolds, cells proliferated during the Wrst week, whereas in PLGA5050, PLGA8515, and PLLA, proliferation began after 1 week (Fig. 6(B)). The overall cell numbers in all scaVold cultures, except PGA, increased during the culture. Notably, similar to chondrocytes, hMSCs also proliferated most rapidly in the PLLA scaVold. Taken together, these results indicate that, in all polymeric scaVolds, chondrocyte proliferation was supported more than hMSC proliferation.

Fig. 5. SEM analysis of cellular interaction with electrospun Wbrous scaVolds at Days 3 and 7. On both Days 3 (C, G, I) and 7 (D, H, J), cells spread and covered the surface of PGA (A, B), PDLLA (C, D), PLGA5050 (G, H), and PLGA8515 (I, J) scaVolds. In comparison, cells in PLLA (E, F) and PCL (K, L) scaVolds adhered to the Wbers, and formed a three-dimensional cell–matrix network. The electrospun Wbrous scaVolds appeared to support cell attachment, and the cells showed guided growth according to Wber orientation. Bar: 10 m.

4. Discussion In this study, we have assessed the potential application of Wbrous matrix as tissue engineering cell scaVold by examining the physical and biological properties of electrospun Wbrous scaVolds fabricated using six commercially available biodegradable polymers. To produce uniform Wbrous scaVolds, we have optimized the fabrication process for each biodegradable polymer, including using diVerent solvents and Wne-tuning electrospinning parameters, such as polymer viscosity and strength and distance of the electric Weld. Our results showed that the properties of the electrospun Wbers are polymer type-dependent. PLLA and PCL based Wbrous scaVolds exhibit favorable physical and biological characteristics, such as the maintenance of scaVold structure in physiological buVer and the support of cell

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Fig. 6. Kinetics of cell proliferation after seeding into electrospun Wbrous scaVolds. (A) Bovine chondrocytes (10,000 in 1 cm £ 1 cm £ 1 mm scaVold); and (B) hMSCs (50,000 in 1 cm £ 1 cm £ 1 mm). Cells seeded on diVerent poly(-hydroxy ester) scaVolds exhibited diVerent proliferation kinetics. PLLA culture exhibited the highest proliferation rate, for both bovine chondrocytes and hMSCs. All values are an estimate of the standard uncertainty expressed as mean § S.D. ¤: p < 0.05, n D 4.

proliferation, and may be more promising for the engineering of more structurally stable connective tissues. On the other hand, PGA, PLGA5050, PLGA8515, and PDLLA Wbrous scaVolds showed rapid polymer degradation, resulting in structural degradation likely to negatively aVect cellular activities. It is noteworthy that the degradation and mechanical properties of polymeric scaVolds are a function of the molecular weight of the polymers. In this study, our goal is to directly compare the fabrication process and properties of electrospun Wbrous polymeric matrix using six representative biodegradable polymers, and the issue of molecular weight will be investigated in a subsequent study. Optimal Wber fabrication parameters include solvent type, voltage, distance, and viscosity, and are speciWc for a particular poly(-hydroxy ester) [16]. An ideal organic

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solvent is one that allows full extension of the polymer. Most biodegradable poly(-hydroxy esters) can be dissolved in organic solvents such as chloroform, THF and DMF, whereas PGA is only soluble in highly Xuorinated organic solvents. Another requirement for a solvent of choice is that it should evaporate completely after the Wber formation process, leaving no trapped solvent residue to redissolve the formed Wbers. By also optimizing other parameters, including electric Weld strength and distance, and polymer viscosity for each polymer, we showed here that consistent uniformed-size Wbers can be produced. It should be pointed out that if even one of these parameters was less than optimal, irregular, bead-containing Wbers formed, suggesting their coordinated requirement during the electrospinning process. The primary factor aVecting the structural stability of a scaVold is polymer degradation, mainly controlled by polymer chemistry and physical features of a scaVold. In general, in an aqueous environment, hydrophilic or amorphous polymers are more readily degraded than hydrophobic or crystalline polymers. PGA exhibiting high crystallinity (45%) [4] is the only hydrophilic biodegradable polymers among these six polymers. Owing to the hydrophilic nature of PGA, this polymer is sensitive to its environment and tends to degrade in aqueous solution. With the addition of a methyl group, PLA is more hydrophobic and degrades more slowly than PGA [4]. The three isoforms of PLA, diVering based on the position of a methyl group in the lactic acid monomer, exhibit distinctive properties. For example, PDLLA, composed of L- and D-stereoisomers of lactic acid, degrades faster than PLLA composed solely of L-isomers. In addition, PLGA copolymers, regardless of the subtypes, are amorphous and exhibit faster degradation rates. In this study, the degradation proWle of a Wbrous scaVold is primarily regulated by the degradation of the constituent polymer. In this manner, the higher degradability of PGA, PLGA5050, PLGA8515, and PDLLA renders the Wbrous scaVolds composed of these polymers to degrade more rapidly than the PLLA and PCL based Wbrous scaVolds. We also observed that a Wbrous scaVold, featuring ultraWne Wbers with high surface area to volume ratio, is more sensitive to hydrolytic attack compared to a bulk scaVold. Our results showed that only PLLA and PCL based Wbrous scaVolds were hydrolysis-resistant during the culture period of 42 d, suggesting that application of Wbrous scaVolds for tissue engineering requires proper polymer chemistry to optimally compensate for their increased biodegradability. Interestingly, hydrophilic PGA Wbrous scaVolds exhibit a distinct degradation proWle compared to other poly(hydroxy ester) Wbrous scaVolds. Upon hydrolysis, the long PGA Wbers degrade into sectioned, short Wbers and, eventually, decompose into polymer powders, whereas PLGA5050, PLGA8515, and PDLLA Wbers swell and are shortened to stubby, Wbers. Microscopically, swelling of the Wbers forces the inter-Wber to close the pores, and macroscopically, the entire structure of the scaVold shrinks because of Wber shortening. We propose a scheme in Fig. 7 to illustrate the

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Fig. 7. Schematic illustration of diVerent modes of degradation for poly(-hydroxy ester) Wbrous scaVolds. (A) Long PGA Wbers break down into short, sectioned Wbers. Fiber diameter remains relatively unchanged during the degradation process. (B) PDLLA, PLGA5050, and PLGA8515 Wbers swell and become larger Wbers, closing pores between Wbers. The entire scaVold undergoes gross shrinkage due to the shortening of Wbers.

diVerences between these two modes of degradation. A similar experimental result was also reported by Zong et al. [17], who showed that PLGA7525 and PDLLA electrospun membranes shrink more than 80%, and that the porosity of these two membranes decreases dramatically during degradation. The degradation process also alters the mechanical properties of the scaVolds. PGA Wbrous scaVold gradually loses mechanical strength with degradation, whereas PLGA5050, PLGA8515, and PDLLA scaVolds increase in stiVness as a result of hydrolysis (data not shown). Most likely, a crystalline structure with stronger stiVness and less structural Xexibility forms as the amorphous polymer resulting from hydrolysis repacks the macromolecular structure of the scaVold. Our results showed that cellular activities are highly dependent on the stability of the scaVold structure. The rapid degradation of PGA, PDLLA, PLGA5050, and PLGA8515 Wbrous scaVolds destroys the porous structure, and negatively aVects cellular adhesion and ingrowth. Unfavorable cell–matrix interactions thus result from the physical instability of the scaVolds. On the other hand, PLLA and PCL Wbrous scaVolds, which maintain sound Wbrous architecture throughout the cell culture period, perform better than the other poly(-hydroxy ester) scaVolds, in terms of supporting cell–matrix interaction and cellular proliferation. These eVects are similar for both chondrocytes and hMSCs, suggesting the generality of matrix inXuence on cells. In our previous study [6], we have shown that in culture bovine chondrocytes proliferate rapidly and quickly reach a growth plateau. Therefore, in this study, we seeded a lower number of chondrocytes, compared to hMSCs, in the Wbrous scaVolds to allow space for cellular proliferation. Chondrocytes in fact exhibit a much more rapid proliferation rate than hMSCs, and similar cell number are reached at end of the 4-week culture period. Upon seeding onto a polymeric biomaterial, cell adhesion to the polymer is the Wrst cellular event to occur, and

cell migration, proliferation, and diVerentiation take place only after cells are securely attached [18]. Therefore, optimal cell adhesion is critical for favorable cellular response. Synthetic biodegradable polymers, unlike natural ECMs, do not have speciWc cell-binding molecular moieties. Cell adhesion to polymers is therefore mediated by serum and/ or ECM proteins adsorbed onto the polymer surface. Protein adsorption is aVected by the hydrophilicity [19] and the surface energy properties [20] of the polymer. Nikolovski et al. showed that vitronectin, compared to Wbronectin, was the predominant matrix protein adsorbed from serum-containing medium onto PGA and PLA [21]. Compared to other polymeric scaVold, an electrospun Wbrous scaVold, because of its higher surface area to volume ratio, is thus able to adsorb more vitronectin and Wbronectin molecules for cell adhesion. Woo et al. reported that Wbronectin and vitronectin preferentially adsorbed to a nanoWbrous scaVold at approximately 2–4 times higher than a solid-walled scaVold [22]. Although the speciWc mechanism(s) responsible for the selective substrate property of nanoWbrous scaVolds remains to be determined, it is clear that the adsorption of cell adhesion matrix molecules enhances cell adhesion. Results from mechanical testing suggest that the mechanical properties of a scaVold are determined by Wber packing density and individual Wber mechanics, i.e., when tensile loading is applied to a Wbrous scaVold, the mechanical response of a scaVold is the sum of the eVects on both factors. In the initial stage of tensile stress–strain testing, non-woven Wbers in a scaVold start to reorient along the direction of the load, suggesting that Wber packing density is a primary factor to determine mechanical response of a scaVold. A scaVold with a higher Wber packing density exhibits stronger mechanical properties, because more force resisting Wber reorientation is generated. When loading continues, Wbers are aligned along the direction of loading, bringing it into the second phase of mechanical response. At this stage, Wber mechanics, controlled by polymer chem-

W.-J. Li et al. / Acta Biomaterialia 2 (2006) 377–385

istry, determines the mechanical response of a scaVold. Engelberg and Kohn compared the mechanical properties of diVerent biodegradable polymers, and showed that PDLLA and PLLA exhibited a higher tensile modulus than PCL, whereas PCL exhibited a much higher percentage of elongation at break [23]. Although the structure of their test material is diVerent from the Wbers studied here, our results conWrmed the response of the composing Wbers to mechanical loading, with both PGA and PLGA Wbrous scaVolds exhibiting stiVer properties and PCL Wbrous scaVold being more compliant. It is thus reasonable to assume that, by electrospinning polymer blends into multi-Wbrous scaVolds, the mechanical properties of a Wbrous scaVold can be varied in a predictable manner. 5. Conclusion In this study, six commonly used, commercially available biodegradable polymers are used to fabricate three-dimensional Wbrous scaVolds using electrospinning technology. Optimized fabrication parameters have been developed to electrospin poly(-hydroxy esters) into Wbrous scaVolds. These uniform Wbrous scaVolds are characterized with respect to their degradation and mechanical properties to assess their applicability as tissue engineering scaVolds. PLLA and PCL Wbrous scaVolds maintain their structural integrity, whereas PGA, PDLLA, PLGA5050, and PLGA8515 scaVolds are readily degraded upon culture incubation. The PLLA and PCL based Wbrous scaVolds show promise as biologically preferred scaVolds/substrates for tissue engineering. Acknowledgements This research was supported by NIH Z01 AR 41113, Intramural Research Program of NIH, NIAMS. The authors thank Dr. Paul Manner, George Washington University, for orthopaedic surgical specimens. References [1] Hutmacher DW. ScaVold design and fabrication technologies for engineering tissues—state of the art and future perspectives. J Biomater Sci Polym Ed 2001;12:107–24. [2] Seal BL, Otero TC, Panitch A. Polymeric biomaterials for tissue and organ regeneration. Mater Sci Eng 2001;34:147–230. [3] Peter SJ, Miller MJ, Yasko AW, Yaszemski MJ, Mikos AG. Polymer concepts in tissue engineering. J Biomed Mater Res 1998;43:422–7. [4] Middleton JC, Tipton AJ. Synthetic biodegradable polymers as orthopedic devices. Biomaterials 2000;21:2335–46.

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