Inverted Colloidal Crystal Scaffolds for Uniform Cartilage

Feb 16, 2010 - ACS eBooks; C&EN Global Enterprise .... Scaffolds with pores of inverted colloidal crystal (ICC) geometry could be fabricated by infilt...
1 downloads 0 Views 2MB Size
Download PDF
Biomacromolecules 2010, 11, 731–739

731

Inverted Colloidal Crystal Scaffolds for Uniform Cartilage Regeneration Yung-Chih Kuo* and Yu-Tai Tsai Department of Chemical Engineering, National Chung Cheng University, Chia-Yi, Taiwan 62102, Republic of China Received November 19, 2009; Revised Manuscript Received February 6, 2010

A uniform distribution of chondrocytes in 3D scaffolds is a critical challenge to cartilage regeneration. This study aims to resolve the problem by showing uniformly distributed chondrogenesis in chitin/chitosan matrix with pores of inverted colloidal crystal (ICC) structure. The results revealed that the effect of solvent on the regularity of colloidal crystal arrays was in the order of ethanol > ethylene glycol > acetone. When the concentration of chitin/chitosan gel was 92%. Over 4 weeks of cultivation, the percentage of biodegradation of ICC scaffolds with pure ethanol was ∼34%. The order in the produced BKCs, glycosaminoglycans (GAGs), and collagen was freeform constructs > ICC constructs with pure ethanol > ICC constructs with 95% acetone. However, the spatial distribution of BKCs in ICC constructs was more uniform than that in freeform constructs. In addition, BKCs could secrete GAGs and type II collagen in the core of ICC constructs, indicating the maintenance of phenotypic chondrocytes and their metabolism. The ICC constructs with well-controlled pore regularity and unique topography can generate uniform tissue-engineered cartilage.

Introduction Diarthrodial cartilage, expressing elastic, lubricative, and loadbearing characteristics, is one of the most important connective tissues in human.1 In a traumatic cartilage injury, the migration of neighboring chondrocytes to the wound site is slow because the normal cartilage contains sparse chondrocytes in sticky extracellular matrices.2 This indicates that the articular cartilage is deficient in the self-recovery ability. Therefore, generating neocartilage greatly attracts biomedical practitioners and researchers. From the aspect of biomaterials, the porosity, pore size, pore geometry, and surface functional group could considerably affect the formation of tissue-engineered cartilage.3-5 Although porous scaffolds might favor chondrocyte growth,6-8 random and unconnected pores in freeform scaffolds could cause a failure in cartilage substitute. Therefore, strict control over the pore structure of scaffolds became a highlighted issue. The scaffolds with a similar pore size and shape could be fabricated by compression molding and particulate leaching method.9 However, these scaffolds contained sealed pores and did not exhibit a specific pore regularity. Scaffolds with pores of inverted colloidal crystal (ICC) geometry could be fabricated by infiltrating Na2SiO3 into latticed microspheres and annealing.10 ICC scaffolds could also be prepared from biocompatible polyacrylamide hydrogel by dissolving monodispersed poly(methyl methacrylate) colloids.11 The two types of ICC scaffolds were applied to the assessment of the colony of human hepatocellular carcinoma HEP G2 and human bone marrow stromal HS-5. In addition, the effective diffusivity of nutrients in ICC scaffolds could be estimated by Brownian dynamics and Monte Carlo simulations.12 Chitin and chitosan are implantable biomaterials with complementary physicochemical properties.13 For example, chitin * To whom correspondence should be addressed. Tel: 886-5-272-0411, ext. 33459. Fax: 886-5-272-1206. E-mail: [email protected].

matrix is strong in tensile resistance and brittle.14 On the contrary, positively charged chitosan with amine groups displays low mechanical strength, high extensibility, and high cell affinity.15 Therefore, hybrid matrices of chitin and chitosan may yield novel biomedical characteristics. For instance, chitin/ chitosan scaffolds with deposited hydroxyapatite could mimic the extracellular substrate for cartilage tissue engineering.16 The goal of this study is to demonstrate uniform chondrogenesis in chitin/chitosan ICC scaffolds. Because the mechanical stress on the articular joints requires consistently sustaining, the distribution of cartilaginous components must be uniform. In addition, chondrocytes show interstitial growth. Therefore, a uniform nutrient supply over a construct is essential during the regeneration of diarthrodial cartilage. We presented the uniformity of the proliferated bovine knee chondrocytes (BKCs), secreted glycosaminoglycans (GAGs), and collagen in the ICC constructs over 4 weeks of cultivation. The distribution of BKCs, GAGs, and type II collagen in the periphery and core of the ICC constructs was studied by immunochemical staining.

Materials and Methods Preparation and Characterization of ICC Scaffolds. Polystyrene (20 mg, PS) microspheres (diameter 160 µm, 2.2% size distribution, Duke Scientific, Fremont, CA) was dehydrated and added to 0.3 mL of ethanol (Riedel-de Hae¨n, Seelze, Germany), acetone (Mallinckrodt, Hazelwood, MO), or ethylene glycol (J. T. Baker, Phillipsburg, NJ) in a container (a cylindrical glass tube glued to a microscope slide) with an inner diameter of 5 mm and height of 30 mm. The PS suspension was ultrasonically vibrated at room temperature for 10 min and placed in a bath-reciprocal shaker at 90 rpm for 1 h. The PS microspheres in ethanol and acetone were dried at 40 °C for 24 h. The PS microspheres in ethylene glycol were dried at 90 °C for 48 h. The colloidal crystals were visualized by a phase-contrast biological microscope (Motic, Richmond, BC, Canada), processed with Photoshop CS3 (Adobe, San Jose, CA), and analyzed by Image-Pro Plus (version 4.5, Media

10.1021/bm901312x  2010 American Chemical Society Published on Web 02/16/2010

732

Biomacromolecules, Vol. 11, No. 3, 2010

Figure 1. Scheme for evaluating regularity efficiency: (a) ordered array inside circular boundary, five rectangular regions for microphotography; (b) two squares in one rectangle (enlarged from part a); (c) ordered colloids (dark circles) in one square (enlarged from part b); and (d) processed image.

Cybernetics, Bethesda, MD). Figure 1 illustrates the sampling and imaging methods for estimating the regularity efficiency, which was defined as (particle area)/(total area) × 100%. The PS arrays were then sintered at 120 °C for 4 h. Chitin (Sigma, St. Louis, MO) gel was prepared by acidolyzing 1.5 to 3 g of chitin with 12 N hydrochloric acid (Riedel-de Hae¨n). The gel was filtrated, centrifuged, and dissolved in ultrapure water (Barnstead, Dubuque, IA). 1.5 to 3 g of chitosan (from crab shells, 85% deacetylation, Sigma) was dissolved in 100 mL of ultrapure water containing 1% acetic acid (Showa, Tokyo, Japan). Chitin and chitosan gels were mixed with 2% (w/w) genipin (Challenge Bioproducts, Taichung, Taiwan) for 5 min and ultrasonically vibrated for 5 min. Figure 2 shows the mechanism of synthesis and structure of chitin/ chitosan biomaterials. The chitin/chitosan gel was infiltrated into the sintered PS template by a vacuum pump (Eyela, Tokyo, Japan) until no bubbles were out from the PS arrays. The infiltrated template was placed at room temperature for 24 h and dried at 40 °C for 24 h. The infiltration and drying were repeated eight times. The template filled with solid chitin/chitosan matrix was immersed in tetrahydrofuran (THF, J. T. Baker) overnight for dissolving the PS microspheres. The matrix containing ICC pores was immersed in acetone for 1 h and in ultrapure water for 5 h, dried at 50 °C for 24 h, and trimmed by a cryostat microtome (Slee, Mainz, Germany) into 3 × 3 × 3 mm. Freeform scaffolds were fabricated by mixing chitin and chitosan gels with 2% (w/w) genipin for 5 min, pouring in a mold of 23 × 23 × 16 mm, ultrasonically vibrating at 4 °C for 1 h, storing in a refrigerator (Sanyo, Osaka, Japan) at -80 °C for 24 h, and lyophilizing in a freeze-dryer (Eyela) for 24 h. Freeform matrices were also trimmed by a cryostat microtome into 3 × 3 × 3 mm. The scaffolds were immersed in ethanol for evaluating porosity. The porosity (P (%)) is defined as P (%) ) Vp/Vs ) (Ww - Wd)/(Fe · Vs) ) [(Ww - Wd)/(Ww - W0)] × 100%, where Vp, Vs, Fe, Ww, Wd, and W0 are the pore volume, the volume of the scaffold (volume of pores and solid matrix), the density of ethanol, the wet weight (including ethanol) of the scaffold, the dry weight of the scaffold, and the weight of the scaffold in ethanol (subtracting buoyancy from Wd), respectively. The Fourier transform infrared (FTIR) spectra of ICC and freeform chitin/chitosan matrices were obtained by an IR absorption spectrophotometer (Shimadzu, Columbia, MD). The matrices were compressed with KBr powders in a ratio of 1:5.

Kuo and Tsai Images of colloidal crystals, ICC, and freeform scaffolds and cultivated constructs were obtained by a scanning electron microscope (SEM, Jeol, Tokyo, Japan). The colloidal crystals were vacuum-dried and sputter-coated with gold. The scaffolds were sliced by a cryostat microtome, washed with Dulbecco’s phosphate-buffered saline (DPBS, Sigma), vacuum-dried, and sputter-coated with gold. The constructs were sliced, washed, treated with 2.5% glutaraldehyde (Fluka, Buchs, Switzerland) for 4 h, desiccated stepwise with ethanol from 70 to 100%, vacuum-dried, and sputter-coated with gold. BKCs. Forearm knees of calves were harvested at a local abattoir within 30 min of animal death and transported to the laboratory in ice-bathed DPBS containing 1% antibiotic-antimycotic solution (Sigma) and 50 mM 2-[4-(2-hydroethyl)-1-piperazine] ethane-sulfonic acid (HEPES, J. T. Baker). Cartilage was sliced into cubes ∼1 mm3 and digested with 0.18% type II collagenase (Sigma) in a humidified CO2 incubator (NuAire, Plymouth, MN) at 37 °C for 24 h. Supernatant was centrifuged at 420g for 5 min. Pellet-containing BKCs were resuspended in Dulbecco’s modified Eagle’s medium (DMEM, Sigma). The concentration of BKCs was determined by trypan blue exclusion using a hemocytometer (Neubauer, Marienfeld, Germany) under a phasecontrast biological microscope. Viability, Biodegradation, and Cultivation. The scaffolds were treated with 70% ethanol for 30 min and dried in a biological safety cabin at room temperature. The suspension of BKCs was injected into the scaffolds with a density of 3 × 104 cells/construct using a syringe. The constructs were incubated in a humidified CO2 incubator at 37 °C for 4 h. After seeding, the constructs were placed in a 24-well plate (Corning, Horseheads, NY) and cultured with DMEM, which was supplemented with 1% antibiotic antimycotic solution, 10 mM HEPES, 50 µg/mL L-(+)-ascorbic acid (J. T. Baker), 10% fetal bovine serum (Sigma), 0.1 mM MEM nonessential amino acid (Sigma), and 0.4 mM L-proline (Sigma) in a CO2 incubator at 37 °C for 8 h. The culture medium was discarded. The constructs were washed with DPBS and centrifuged at 130g for 5 min. One construct was reacted with 50 µL of 3-(4,5-dimethyl thiazol-2-yl)-2,5-diphenyl-tetrazolium bromide (MTT, 5 mg/mL, Sigma) in a CO2 incubator at 37 °C for 4 h. MTT was removed, and 500 µL of MTT solubilization solution (Sigma) was applied in darkness. The solution (200 µL) in a 96-well MicroWell polystyrene plate (Nalge Nunc, Rochester, NY) was detected by a UV-vis spectrophotometer (Bio-Tek Instruments, Winooski, VT). The viability of BKCs is defined as (Ac/As) × 100%, where Ac and As are the absorbance at 570 nm after cultivation and after seeding, respectively. One BKC-seeded construct was placed in one well of a 12-well plate (BD, Franklin Lakes, NJ) and cultivated with 5 mL of the culture medium in a CO2 incubator at 37 °C over 28 days. The culture medium was replaced every two days. The constructs were digested with 95.24 µg papain (Sigma) in 2 mL of tris-hydroxyl methyl-amino-methane (Riedel-de Hae¨n) containing 55 mM trisodium citrate (Hanawa, Osaka, Japan), 150 mM sodium chloride (J. T. Baker), 5 mM cysteine (Fluka), and 5 mM ethylene-diamine-tetra-acetic acid (Riedel-de Hae¨n) at 60 °C for 24 h. The percentage of biodegradation (D (%)) is defined as D (%) ) [(Wd - Wr)/Wd] × 100%, where Wd and Wr are the initial dry weight of a scaffold and the residual dry weight of a scaffold after cultivation, respectively. The papain-digested solution was used for assaying cartilaginous components in the cultured constructs. The amount of BKCs in a construct (ABKC) was determined by a fluorescence spectrophotometer (F-4500, Hitachi, Tokyo, Japan) at excitation of 365 nm and emission of 458 nm with Hoechst no. 33258 (0.1 µg/mL, Sigma) and calibrated by freshly isolated BKCs. The Hoechst no. 33258 staining was widely adopted as a standard method to determine the amount of DNA in cultured cells.17-19 Because the average weight of a chondrocyte is 0.1 ng,20 the weight percentage of BKCs in a construct (PBKC (%)) is defined as PBKC (%) ) (10|10 × ABKC/Wd) × 100%. The weight of GAGs in a construct (WGAG) was analyzed by a UV-visible spectrophotometer at 525 nm with 1,9-dimethyl-methylene blue (Sigma) and calibrated by chondroitin sulfate (Sigma). The weight percentage of GAGs in a

ICC Scaffolds for Uniform Cartilage Regeneration

Biomacromolecules, Vol. 11, No. 3, 2010

733

Figure 2. Reaction mechanism of chitin/chitosan matrices: (A) genipin, (B) chitosan, (C) chitin, and (D) product.

construct (PGAG (%)) is defined as PGAG (%) ) (WGAG/Wd) × 100%. N-Chloro-p-toluene-sulfonamide (Sigma) was used to oxidize collagen after acidolysis. The weight of hydroxyproline in a construct (Whp) was obtained by a UV-visible spectrophotometer at 550 nm with 4-dimethylamino-benzaldehyde (Sigma) and calibrated by L-4-hydroxyproline (Fluka). Because the ratio of collagen to hydroxyproline is ∼10,21 the weight percentage of collagen in a construct (Pcollagen (%)) is defined as Pcollagen (%) ) (10 × Whp/Wd) × 100%. Histological and Immunofluorescent Staining. The constructs were fixed with 10% formaldehyde (Sigma) for 1 h, sliced by a cryostat microtome into samples of 4 µm, and treated with o-xylene (Fluka) for 2 h. The samples were immersed in hematoxylin (Sigma) for 2 min, rinsed with 0.5% hydrochloric acid (Riedel-de Hae¨n), immersed in eosin (Sigma) for 2 min, dehydrated in 100% ethanol, and treated with o-xylene for 5 min. In addition, the microtomic samples were immersed in safranin-O (Sigma) for 1 h, dehydrated with ethanol, and treated with o-xylene. BKCs and GAGs in the constructs were visualized by a phase-contrast biological microscope. Immunostaining against type II collagen in the constructs was obtained by immersing in 1% hydrogen peroxide (Sigma) for 5 min, reacting with pepsin (Lab. Vision, Fremont, CA) at 37 °C for 10 min, incubating with serum blocking solution (nonbiotin amplification system, NBATM kit, Zymed Lab., South San Francisco, CA) for 1 h, incubating with collagen II Ab-2 (clone 2B1.5, mouse monoclonal antibody, Lab. Vision) in a dilution ratio of 1:100 for 1 h, and incubating with antichicken IgY fluorescein isothiocyanate (FITC) conjugate (Abcam, Cambridge, U.K.) for 1 h in darkness. Fluorescent microstructures of type II collagen in

the constructs were obtained by a phase contrast fluoromicroscope (Axioskop 2 plus, Zeiss, Munchen-Hallbergmoos, Germany) using an argon laser with a filter at excitation of 458 nm and emission of 488 nm. The surface density of type II collagen was calibrated by collagen type II-FITC conjugate (from bovine articular cartilage, Sigma). The fluorescence density is determined by Image-Pro Plus in density (sum) mode concerning the green area and intensity in the red-green-blue model. The samples from the construct boundary to 200 µm in depth were defined as the peripheral samples. The samples in the core region were obtained from 1400-1600 µm in depth (from construct boundary).

Results and Discussion Regularity, Porosity, Functional Group, and Morphology. Figure 3 shows the regularity efficiency of colloidal crystal templates. Application of ultrasonic vibration to a particle dispersion in solvent could yield a self-assembly array. As revealed in Figure 3, an increased concentration of ethanol and ethylene glycol enhanced the regularity efficiency. On the contrary, an increased concentration of acetone reduced the regularity efficiency. The regularity efficiency was in the order of ethanol > ethylene glycol > acetone. These results were mainly because PS particles are hydrophobic (nonpolar), whereas ethanol and ethylene glycol are polar solvents, and acetone is relatively nonpolar. Because the affinity of acetone to PS particles was high, it could be difficult to disperse the

734

Biomacromolecules, Vol. 11, No. 3, 2010

Figure 3. Effect of solvent concentration on regularity efficiency of colloidal crystal arrays: 4, ethanol; O, ethylene glycol; 0, acetone. n ) 20.

Kuo and Tsai

was freeform scaffolds > ICC scaffolds with pure ethanol > ICC scaffolds with 95% acetone. These results were explained below. The pores in freeform scaffolds originated from lyophilization of nonuniform ice crystals in chitin/chitosan matrices. An increase in the gel concentration could decrease water content, ice crystal density, and porosity. Freeform scaffolds with porosity >85% were regarded as proper milieus for cell proliferation.23,24 For example, the porosity of freeform polylactide-co-glycolide scaffolds could be >90%.25 The chitin/ chitosan gel filled the space among colloidal PS crystals, and the dissolution of the PS particles produced the pores with ICC structure. Therefore, the template particles determined the porosity. The gel concentration could scarcely alter the porosity. Because chitin and chitosan were abundant in hydrogen bonds, an increased gel concentration enhanced the viscosity.26 When 3% gel was used, the viscosity was about 1500-1700 cP. This viscous gel could not completely occupy the space among the PS particles after infiltration, rendering residual voids and a promoted porosity. In addition, the order in the porosity with pure ethanol and 95% acetone was consistent with the regularity efficiency of colloidal crystal template (data shown in Figure 3). Lower regularity efficiency denoted fewer particles in unit volume, more voids, and easier infiltration. Therefore, when 3% gel was introduced, ICC scaffolds with 95% acetone produced a smaller porosity increase than ICC scaffolds with pure ethanol, as indicated in Figure 4. Although freeform scaffolds favored cell growth, the diffusion of nutrients in ICC scaffolds could be more efficient than that in freeform scaffolds.9,12 Hence, the only information from porosity is not sufficient to judge the ability of a scaffold in cell culture. In a study on hydroxyapatite scaffolds, microcomputer tomography was used to evaluate the size distribution and interconnectivity of pores.27 However, the pore size and geometry in ICC scaffolds were controlled by arrayed PS microspheres. The ratio of pore surface area to apparent volume of an ICC scaffold can be estimated by

[ Figure 4. Effect of gel concentration on porosity of freeform and ICC scaffolds. 4, freeform; ]: ICC, pure ethanol; O: ICC, 95% acetone. n ) 3.

particles by ultrasonic vibration. No close-packed array with pure acetone was observed because acetone might dissolve PS chains on the surface, rendering a partially solvated external layer and irreversible aggregation. In addition, the dispersion of the ultrasound-treated PS particles in ethanol and ethylene glycol was more uniform than that in acetone. Ethylene glycol could produce a looser particle arrangement than ethanol via repelling the floating PS particles to the container surfaces. Therefore, ethylene glycol led to a relatively disorderly template. However, ethanol assisted the particle assembly by appropriate vapor pressure and density.11,22 Figure 4 shows the porosity of freeform and ICC scaffolds. As indicated in this Figure, the porosity of freeform scaffolds decreased when the gel concentration increased. The porosity of freeform scaffolds was >89%. When the gel concentration was ICC scaffolds with pure ethanol > ICC scaffolds with 95% acetone. This order was the same as the order in porosity (data shown in Figure 4), and this order was inverted for dry scaffold weight. A light scaffold was more sensitive to the weight loss than a heavy scaffold, in general. Therefore, a lighter scaffold with higher porosity yielded faster biodegradation. In addition, hydrolysis and enzymatic degradation were the two mechanisms of weight loss in chitin/chitosan matrix.39,40 A freeform scaffold accelerated the biodegradation in the first two weeks because its high porosity promoted hydrolysis. A mitigating increase

Figure 8. Chondrogenesis in freeform and ICC constructs over 4 weeks of cultivation. (a) proliferated BKCs; (b) excreted GAGs; (c) produced collagen. 2.5% gel. 4, freeform; ], ICC, pure ethanol; O, ICC, 95% acetone. n ) 3.

Figure 7. Biodegradation of freeform and ICC scaffolds. 2.5% gel. 4, freeform; ]: ICC, pure ethanol; O: ICC, 95% acetone. n ) 3.

in the weight loss followed in the third and fourth week. ICC scaffolds accelerated the biodegradation in the last two weeks. This was because the distribution of BKCs and ECM was uniform over an ICC construct (images shown in Figure 9), stimulating enzymatic degradation. From the third to

ICC Scaffolds for Uniform Cartilage Regeneration

Biomacromolecules, Vol. 11, No. 3, 2010

737

Table 2. Surface Density of Type II Collagen in Constructsa density of type II collagen (µg/cm2) freeform conctruct (periphery) freeform conctruct (center) ICC conctruct (periphery) ICC conctruct (center) a

Figure 9. Staining of freeform and ICC constructs over 4 weeks of cultivation. 2.5% gel. ICC constructs with pure ethanol: (a) H&E, freeform, periphery; (b) H&E, freeform, core; (c) H&E, ICC, periphery; (d) H&E, ICC, core; (e) safranin O, freeform, periphery; (f) safranin O, freeform, core; (g) safranin O, ICC, periphery; (h) safranin O, ICC, core; (i) type II collagen, freeform, periphery; (j) type II collagen, freeform, core; (k) type II collagen, ICC, periphery; and (l) type II collagen, ICC, core.

fourth week, ICC scaffolds with pure ethanol produced a greater increase in the weight loss than those with 95% acetone. This was because the number of BKCs in the former was larger than that in the latter (data shown in Figure 8a). Over 4 weeks of cultivation, the biodegradation of ICC scaffolds with pure ethanol was ∼34%. The percentage of biodegradation of freeform chitin/chitosan scaffolds over 4 weeks of dynamic cultivation was between 30 and 60%,

2.41 ( 0.20 0.27 ( 0.18 2.02 ( 0.17 1.53 ( 0.26

ICC constructs with pure ethanol. n ) 3.

depending on the prefreezing temperature and the ratio of chitin to chitosan.16 As compared with dynamic cultivation, static culture slightly retarded the biodegradation. Chondrogenesis in ICC Constructs. Figure 8 shows the production of cartilaginous components in the cultured constructs. As revealed in Figure 8a, BKCs could proliferate in the constructs, and the order in the weight percentage of BKCs was freeform constructs > ICC constructs with pure ethanol > ICC constructs with 95% acetone. This order was the same as the order in porosity; that is, a construct with higher porosity yielded a larger amount of BKCs. In addition, the proliferation of BKCs was fast in the first two weeks. The proliferation was retarded in the third and fourth week. The order in the growth retardation was freeform constructs > ICC constructs with 95% acetone > ICC constructs with pure ethanol. This retardation order could be explained below. First, the sealed pores in freeform constructs restrained the migration of BKCs. Hence, BKCs in freeform constructs were stacked up and formed large cell clusters in the last two weeks. Second, ICC constructs with interconnected pores enhanced cell ingrowth (images shown in Figure 9). In addition, ICC constructs with pure ethanol produced more effective surfaces for the colony of BKCs than ICC constructs with 95% acetone. As exhibited in Figure 8b,c, BKCs could also secrete GAGs and collagen in the constructs. The order in the weight percentage of GAGs and collagen was freeform constructs > ICC constructs with pure ethanol > ICC constructs with 95% acetone. In the first two weeks, the synthesis of ECM could be closely correlated with the proliferation of BKCs. In the third and fourth week, the production of ECM was not inhibited, in general. This was because welladapted BKCs in the constructs could stably secrete ECM. However, the synthesis of collagen in freeform scaffolds was retarded in the fourth week. This exception suggested that ICC constructs could be more efficient in the synthesis of collagen than freeform scaffolds. In a study on oxygen consumption in freeform gels, it was concluded that the chondrocytic activity reduced in the internal portion of the biomaterials.30 However, the supply of oxygen to the core of ICC constructs could be adequate. The transport of oxygen could explain why ICC constructs exhibited a less significant retardation effect on chondrogenesis than freeform constructs. Figure 9 shows H&E staining, Safranin O staining, and staining against type II collagen of freeform and ICC constructs. The purple dots in Figure 9a-d denoted basophilic neucli of BKCs. As revealed in Figure 9a, the periphery of freeform constructs was abundant in proliferated BKCs. These BKCs aggregated into cell clusters and existed in the pore center. As displayed in Figure 9b, BKCs scattered in the core of freeform constructs. BKCs might encounter difficulty in colonizing confined compartments, leading to fewer cells in the core region. As indicated in Figure 9c,d, BKCs grew along the pore walls of ICC constructs. The pores in these samples were not precisely sectioned in half. Therefore, the circle of pore rim in the images was not the projection of the void space in ICC scaffolds. In addition, the number of BKCs in the periphery was larger than that in the core.

738

Biomacromolecules, Vol. 11, No. 3, 2010

Kuo and Tsai

Table 3. Qualitative Summary about the Scaffolds fabrication (day)

pore size

pore regularity

porosity

ECM

BKC ingrowth

3 20 20

nonuniform uniform uniform

disorderly ordered partially ordered

high medium low

high medium low

poor good

freeform scaffold ICC scaffold with pure ethanol ICC scaffold with 95% acetone

However, the reduced amount of BKCs in the core of ICC constructs was minor. As compared with Figure 9a, Figure 9c did not present the clusters of BKCs in the pores. On the basis of this observation, a conclusion could be drawn that ICC constructs generated less steric hindrance for mass transfer than freeform constructs. In a theoretical analysis, it was found that cells were liable to reside in the exterior zone of a scaffold, and the distribution of cells could be correlated with the depletion of nutrient supply.41 The distribution of GAGs (red spots) in the cultured constructs was shown in Figure 9e-h. As displayed in Figure 9e,f, GAGs could be distributed over the cavity in freeform constructs. On the contrary, GAGs spread mainly on the pore surface in ICC constructs, as indicated in Figure 9g,h. In addition, the number of GAGs in the periphery was larger than that in the core. GAGs in the core of freeform constructs reduced more seriously than those of ICC constructs. This suggested that the distribution of GAGs in ICC constructs was more uniform than that in freeform constructs. The green patches in Figure 9i-l were the expressed type II collagen, demonstrating the phenotypic chondrocytes in the regenerated neocartilage.42 As revealed in Figure 9i, the green patches could fill the cavity in freeform constructs. This implied that BKCs could be stacked up in the pores. In addition, the pores in freeform scaffolds displayed tubular morphology (image shown in Figure 6c), and the efficiency in the diffusion of gases and nutrients from the pore center to the pore wall could be poor. Therefore, only a little type II collagen adhered to the pore surface. Moreover, a drastic decrease in the green patches was observed in the core of freeform constructs, as displayed in Figure 9j. Figure 9k,l exhibited localized type II collagen on the pore surface of ICC constructs. The green patches decreased slightly in the core region. As compared with Figure 9j, Figure 9l revealed higher intensity of the green patches. Table 2 lists the density of type II collagen on the construct surface. Figure 9i-l are consistent with Table 2. Table 3 summarizes the general characteristics of the scaffolds and their effect on chondrogenesis. As shown in this Table, an ICC scaffold required 6.7 times the preparation time of a freeform scaffold. However, ICC scaffolds generated higher pore uniformity and regularity than freeform scaffolds. Although freeform scaffolds produced higher porosity than ICC scaffolds, the latter yielded better ingrowth of BKCs than the former.

Conclusions The chitin/chitosan scaffolds with pores of ICC geometry were fabricated for generating uniform neocartilage. Because chitin/chitosan gel was very viscous, fabricating ICC scaffolds required a complicated technique, high cost, and long processing time. As a result, ICC scaffolds yielded regular and interconnected cavities with identical size. In freeform constructs, the produced BKCs, GAGs, and collagen decreased significantly from the periphery to the core region. However, BKCs tended toward migration to the core of ICC constructs, avoiding formation of the pores without cells or large cell clusters. The ICC topography can enhance the ingrowth of chondrocytes and

homogeneous secretion of ECM over a construct, leading to uniform cartilage regeneration. Acknowledgment. This work is supported by the National Science Council of the Republic of China.

Notation Cg: concentration of chitin/chitosan gel (%) Cs: concentration of solvent for arraying colloidal crystals (%) D: percentage of weight loss of a scaffold in biodegradation (%) E: regularity efficiency of colloidal crystal arrays (%) P: porosity of a scaffold (%) PBKC: weight percentage of BKCs in a construct (%) PGAG: weight percentage of GAGs in a construct (%) Pcollagen: weight percentage of collagen in a construct (%)

References and Notes (1) Mobasheri, A.; Carter, S. D.; Martı´n-Vasallo, P.; Shakibaei, M. Cell Biol. Int. 2002, 26, 1–18. (2) O’Driscoll, S. W. J. Bone Joint Surg. 1998, 80-A, 1795–1812. (3) Yamane, S.; Iwasaki, N.; Kasahara, Y. J. Biomed. Mater. Res. 2007, 81A, 586–593. (4) Griffon, D. J.; Sedighi, M. R.; Schaeffer, D. V.; Eurell, J. A.; Johnson, A. L. Acta Biomater. 2006, 2, 313–320. (5) Wan, Y.; Tu, C.; Yang, J.; Bei, J.; Wang, S. Biomaterials 2006, 27, 2699–2704. (6) Raghunath, J.; Rollo, J.; Sales, K. M. Biotechnol. Appl. Biochem. 2007, 46, 73–84. (7) Lee, C. T.; Huang, C. P.; Lee, Y. D. Biomacromolecules 2006, 7, 2200–2209. (8) Kuo, Y. C.; Ku, I. N. Biomacromolecules 2008, 9, 2662–2669. (9) Hou, Q.; Grijpma, D. W.; Feijen, J. Biomaterials 2003, 24, 1937– 1947. (10) Kotov, N. A.; Liu, Y.; Wang, S.; Cumming, C.; Eghtedari, M.; Vargas, G.; Motamedi, M.; Nichols, J.; Cortiella, J. Langmuir 2004, 20, 7887– 7892. (11) Zhang, Y.; Wang, S.; Eghtedari, M.; Motamedi, M.; Kotov, N. A. AdV. Funct. Mater. 2005, 15, 725–731. (12) Shanbhag, S.; Leeb, J. W.; Kotov, N. A. Biomaterials 2005, 26, 5581– 5585. (13) Khor, E.; Lim, L. Y. Biomaterials 2003, 24, 2339–2349. (14) Vachoud, L.; Zydowicz, N.; Domard, A. Carbohydr. Res. 2000, 326, 295–304. (15) Suh, J. K. F.; Matthew, H. W. T. Biomaterials 2000, 21, 2589–2598. (16) Kuo, Y. C.; Lin, C. Y. Biotechnol. Bioeng. 2006, 95, 132–144. (17) Murphy, C. M.; Haugh, M. G.; O’Brien, F. J. Biomaterials 2010, 31, 461–466. (18) Lien, S. M.; Ko, L. Y.; Huang, T. J. Acta Biomater. 2009, 5, 670– 679. (19) Tierney, C. M.; Haugh, M. G.; Liedl, J.; Mulcahy, F.; Hayes, B.; O’Brien, F. J. J. Mech. BehaV. Biomed. Mater. 2009, 2, 202–209. (20) Freed, L. E.; Vunjak-Novakovic, G.; Biron, R. J.; Eagles, D.; Lesnoy, D.; Barlow, S. K.; Langer, R. Bio/Technology 1994, 20, 689–693. (21) Hollander, A. P.; Heathfield, T. F.; Webber, C.; Iwata, Y.; Bourne, R.; Rorabeck, C.; Poole, R. A. J. Clin. InVest. 1994, 93, 1722– 1732. (22) Liu, Y.; Wang, S.; Lee, J. W.; Kotov, N. A. Chem. Mater. 2005, 17, 4918–4924. (23) Hutmacher, D. W. Biomaterials 2000, 21, 2529–2543. (24) Mao, J. S.; Zhao, L. G.; Yin, Y. J. Biomaterials 2003, 24, 1067– 1074. (25) Whang, K.; Thomas, C. H.; Healy, K. E.; Nuber, G. Polymer 1995, 36, 837–842.

ICC Scaffolds for Uniform Cartilage Regeneration (26) Ravi Kumar, M. N. V. React. Funct. Polym. 2000, 46, 1–27. (27) Jones, A. C.; Arns, C. H.; Hutmacher, D. W.; Milthorpe, B. K.; Sheppard, A. P.; Knackstedt, M. A. Biomaterials 2009, 23, 1440– 1451. (28) Tangsadthakun, C.; Kanokpanont, S.; Sanchavanakit, N.; Pichyangkura, R.; Banaprasert, T.; Tabata, Y.; Damrongsakkul, S. J. Biomater. Sci., Polym. Ed. 2007, 18, 147–163. (29) Wang, T.; Turhan, M.; Gunasekaran, S. Polym. Int. 2004, 53, 911– 918. (30) Guaccio, A.; Borselli, C.; Oliviero, O.; Netti, P. A. Biomaterials 2008, 29, 1484–1493. (31) Jianga, T.; Abdel-Fattahb, W. I.; Laurencin, C. T. Biomaterials 2006, 31, 4894–4903. (32) Wu, H.; Wan, Y.; Cao, X.; Wu, Q. Acta Biomater. 2008, 4, 76–87. (33) Deng, Y.; Zhao, K.; Zhang, X. F.; Hu, P.; Chen, G. Q. Biomaterials 2002, 23, 4049–4056. (34) Kuo, Y. C.; Ku, I. N. Biotechnol. Prog. 2007, 23, 238–245.

Biomacromolecules, Vol. 11, No. 3, 2010

739

(35) Woodfield, T. B. F.; Malda, J.; de Wijn, J.; Peters, F.; Riesle, J.; van Blitterswijk, C. A. Biomaterials 2004, 25, 4149–4161. (36) Kafienah, W.; Jakob, M.; Demarteau, O.; Frazer, A.; Barker, M. D.; Martin, I.; Hollander, A. P. Tissue Eng. 2002, 8, 817–826. (37) Klein, T. J.; Malda, J.; Sah, R. L.; Hutmacher, D. W. Tissue Eng. B 2009, 15, 143–157. (38) Blewis, M. E.; Schumacher, B. L.; Klein, T. J.; Schmidt, T. A.; Voegtline, M. S.; Sah, R. L. J. Orthop. Res. 2007, 25, 685–695. (39) Freier, T.; Koh, H. S.; Kazazian, K.; Shoichet, M. S. Biomaterials 2005, 26, 5872–5878. (40) Kurita, K.; Kaji, Y.; Mori, T.; Nishiyama, Y. Carbohydr. Polym. 2000, 42, 19–21. (41) Shanbhag, S.; Wang, S.; Kotov, N. A. Small 2005, 1, 1208–1214. (42) Grad, S.; Kupcsik, L.; Gorna, K.; Gogolewski, S.; Alini, M. Biomaterials 2003, 24, 5163–5171.

BM901312X