ARTICLE pubs.acs.org/cm
Three-Dimensional Collagen/Alginate Hybrid Scaffolds Functionalized with a Drug Delivery System (DDS) for Bone Tissue Regeneration Hyeong-jin Lee, Seung-Hyun Ahn, and Geun Hyung Kim* Department of Mechanical Engineering, Bio/Nanofluidics Lab, Chosun University, 375 Seosok-dong, Dong-gu, Gwangju 501759, South Korea ABSTRACT: Biomedical scaffolds have recently evolved into various functional materials, including drug delivery systems (DDS). Here, we report the development of a new, highly porous scaffold based on a layer-by-layer collagen scaffold coated with an alginate polymer, which shows improved mechanical properties and controllable drug release without loss of the original biological function of the collagen scaffold. In particular, the scaffold (75 vol % alginate in a collagen scaffold with a porosity of 88%) attained a Young’s modulus of 30 MPa, which is ∼9 times the value for the pure collagen scaffold (porosity = 98%). Although the scaffolds are highly porous, the drug release and initial burst were wellcontrolled with an appropriate volume fraction of alginate. Osteoblast-like cells (MG63) readily proliferated and migrated into the interior of the scaffolds, and calcium and phosphate on the cell surfaces were well-formed, similarly on pure collagen and alginate/ collagen scaffolds, within only 7 days of culture. The alginate/collagen scaffolds with a drug delivery function have potential as biomedical scaffolds for clinical use in soft and hard tissue regeneration. KEYWORDS: tissue engineering, bone, biomaterials, collagen, alginate
’ INTRODUCTION To regenerate various tissues, including bone, skin, and nerve, biomedical scaffolds should provide mechanical and biological support for neo-tissue formation and survival.1 To acquire this biological function, scaffolds should be biocompatible (to improve cellular behavior) and biodegradable (to avoid the need for later surgical removal), and they must have appropriate mechanical properties. In addition to their conventional purposes, recent studies have expanded the functionality of scaffolds to assist in cell growth and differentiation through the release of various bioactive compounds, such as bone morphogenetic protein-2 (BMP-2),2,3 which has a stimulating effect on bone formation in orthopedic applications, transforming growth factor-β (TGF-β),4 and fibroblast growth factor (bFGF).5 In addition, to avoid side effects, such as contamination and inflammation, which may appear when the scaffold is implanted in the body, the scaffold can be loaded with various drugs, such as vancomycin (a hydrophilic antibiotic drug, which can be used for preventing osseous staphylococcal infections after surgery).6,7 Such functional scaffolds have been used to efficiently regenerate various tissues, including bone, cartilage, and skin.810 In normal drug delivery systems (DDS) using biodegradable polymers and/or composites, a widely accepted method involves controlling the rate of degradability of the materials.11 Drug release by the biodegradable materials can be obtained by diffusion of the injected drugs at the surface of the materials as those in contact with in vitro phosphate-buffered saline (PBS). Recently, drug release systems have been improved by the use of various porous microspheres,12 r 2011 American Chemical Society
coreshell structured nanofiber,13 and hybrid structures using rapid-prototyped structures with microfibrous/nanofibrous mats.14,15 Although various techniques have been developed for DDS using scaffolds, the ultimate goal of functional scaffolds— i.e., precise manipulation of not only the amount and release rate of biological compounds or drugs, but also the scaffold to provide good sites of cell proliferation and differentiation—has yet to be attained. Because collagen enhances cell attachment and proliferation through interactions between the Arg-Gly-Asp (RGD) domains in collagen molecules and integrin receptors in the cell membrane,16 it has been considered a highly promising biomedical material. In addition, collagen is the major protein component of the extracellular matrix and it has high water affinity and low antigenicity,17 so various scaffolds based on collagen have been studied for biomedical applications in regeneration of soft and hard tissues, such as skin, vascular grafting, and bone.16,18 To overcome the mechanical shortcomings of the collagen scaffold, it has been supplemented with various synthetic and natural materials, such as glycosaminoglycan,18 chitosan,19 alginate,20 polycaprolactone,21 and poly(lactic acid) (PLA).22 Collagen scaffolds have also been combined with other bioactive agents Special Issue: Materials for Biological Applications Received: March 13, 2011 Revised: May 16, 2011 Published: June 16, 2011 881
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to control the release rate of growth factors and increase their therapeutic effect.23 However, the porous collagen scaffolds developed for DDS have been in the shape of spongy and porous microspheres, so it has been difficult to control the amount and release rate of the drugs and various growth factors from lyophilized structures, because of their random surface mesopores/ micropores. Maeda et al.24 reported that 80% of human serum albumin injected into lyophilized collagen sponge was released within 6 h following the start of the release test, because of rapid infiltration of the release medium into the porous scaffold. In addition, although various techniques have been developed to improve the mechanical properties of collagen scaffolds, poor mechanical strength is still a problem for use in hard tissue regeneration with no loss of biological function. The goal of this research was to develop a DDS using a lyophilized collagen scaffold. For this purpose, we propose a new and simple coating system that combines alginate with our previous cryogenic plotting method.25 Alginate is a biocompatible polysaccharide that has been extensively studied as a scaffold for cell and growth factor encapsulation and also for gene delivery, because of its rapid gelation in calcium chloride.26,27 For these reasons, alginate has been widely used in tissue engineering, including the regeneration of skin,28 bone,29 and cartilage,30 because of its ability to accelerate epithelialisation and granulation tissue formation.28 Using alginate and a simple coating process with the cryogenic plotting method, we designed a three-dimensional (3D) structured collagen scaffold with a drug release system and improved mechanical properties. First, we fabricated a lyophilized collagen scaffold that was cross-linked with a cross-linking agent. After the process of drug absorption, it was coated with a thin layer of alginate. To validate the drug release characteristics, rhodamine-B, which has previously been used as an indicator of drug release,31 was used. In addition, to observe the cellular behavior of the alginate/collagen scaffolds, cell viability and alkaline phosphatase (ALP) activity were measured using osteoblast-like cells (MG63) for bone tissue regeneration.
Table 1. Volume Fraction of Alginate in the CAC Scaffolds and Density and Porosity of CAC Scaffolds CAC-1 scaffold collagen:alginate
a
CAC-2 scaffold
CAC-3 scaffold
71:29
43:57
(wt %) vol% of alginate
v28:72
32.4 ( 7
60.9 ( 7
75.1 ( 10
density
1.24 g/cm3
1.18 g/cm3
1.15 g/cm3
porositya
92.5% ( 4%
88.9% ( 3%
87.9% ( 5%
Porosity of pure collagen scaffold = 98.5% ( 3%.
A Fourier-transform infrared (FTIR) spectrometer (Model 6700, Nicolet, West Point, PA, USA) was used for measuring the cross-linkage of collagen and alginate. IR spectra represent the average of 30 scans between 400 cm1 and 4000 cm1 at a resolution of 8 cm1. Water absorption was calculated by weighing the scaffolds before and after soaking in distilled water for 2 h. The percent increase in water absorption was calculated as ð%Þ ¼
W2 h W0 100 W0
where W2 h is the weight of the scaffold after 2 h and W0 is the original weight of the scaffold at time zero. The mechanical properties of the scaffolds were evaluated in a dry state using the tensile mode. The scaffolds were cut into small strips (10 mm 20 mm 1 mm). For each scaffold, five samples were obtained from different sites. The size of the specimens was considered as a rectangular shape and was measured using a digital caliper micrometer (Ultracal III; Sylvac, Bern, Switzerland). To determine the size, three different parts of the specimen were measured and averaged. The Young’s modulus and maximum strength were characterized using a universal tensile machine (Top-tech 2000; Chemilab, Suwon, South Korea). The stressstrain curves of the scaffolds were recorded at a stretching speed of 0.5 mm/s, and the Young’s modulus was calculated in 1% strain. The apparent porosity of the scaffolds was determined using the following equation:
porosity ð%Þ ¼ 1
’ EXPERIMENTAL SECTION
Materials. Porcine type I atelocollagen (Matrixen-PSP; Bioland, Cheonan, South Korea) was used for scaffold fabrication. A collagen solution was prepared in 0.05 M acetic acid (pH 3.2) at a fixed concentration of 4.5% (w/v). The measured viscosity of the collagen solution was 30.2 Pa s (spindle velocity = 17 mm/s) at 25 °C. To crosslink the lyophilized collagen scaffold, it was immersed in a 50 mM 1-ethyl-(3-3-dimethylaminopropyl) carbodiimide hydrochloride (EDC, Mw = 191.7; Sigma, St. Louis, MO, USA) solution in 95% ethanol for 24 h at room temperature. Intermediate-G sodium alginate was obtained from SigmaAldrich. Sodium alginate at 5, 10, and 20 wt % in 5 wt % NaCl solution was used as a coating material. To cross-link the alginate solution, the collagen scaffold coated with alginate was immersed in 1 wt % CaCl2 solution. To characterize the drug release behavior from the scaffold, rhodamine-B (SigmaAldrich) was used as an indicator. Scaffold Characterization. The 3D pore size and roughened surface of the scaffolds were observed under an optical microscope (Model BX FM-32; Olympus, Tokyo, Japan) that was connected to a digital camera, and a scanning electron microscope (SEM; Sirion, Hillsboro, OR, USA). To observe the infiltration of the coated alginate solution into the collagen scaffold, the alginate was mixed with rhodamine, which is red in color, and images were acquired with an optical microscope.
1 bulk density of scaffold
weight of scaffold volume of scaffold assumed as a rectangular shape
!
ð1Þ The density of the composites was determined using the rule of mixtures. The bulk densities of collagen and alginate were 1.3 g/cm3 and 1.1 g/cm3, respectively. The weight percent (ωa) of alginate in the collagen scaffold was changed to a volume fraction (j), which was calculated using the following simple equation: j¼
ωa ωa ð1 λÞ + λ
ð2Þ
where ωa is the weight fraction of alginate, λ = Fa/Fc, and Fa and Fc are the densities of alginate and collagen, respectively. The density of the alginate/collagen scaffold consisting of collagen and 29 wt % (32 vol %) alginate was 1.24 g/cm3, and the density of the composite consisting of PCL and 72 wt % (75 vol %) alginate was 1.15 g/cm3. The scaffolds were weighed with a precise balance (AD-4 Autobalance; PerkinElmer, Waltham, MA, USA). The volume fraction and density for three different coated collagen scaffolds are described in detail in Table 1. Drug release results from the scaffolds using rhodamine as an indicator were determined by incubating the scaffolds in 20 mL of PBS solution at 37 °C. To measure rhodamine release, 1 mL of the drug 882
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Figure 1. Fabrication procedure for 3D alginate/collagen scaffold with drug delivery function: (a) schematic diagram of the desired scaffold using a cryogenic plotting system (T = 30 °C); (b) lyophilized and cross-linked collagen scaffold, (c) collagen scaffold with the absorption of rhodamine solution (red color) as a drug indicator; and (d) coating/suction and drying process with alginate solution and optical and SEM images of the crosssection of the final scaffold. release solution was extracted and tested for the drug and then returned to the release solution. To avoid contamination of the release solution, the pipet tip and cuvette of the spectrophotometer were washed with 70% ethanol. Rhodamine released from alginate/collagen scaffolds was monitored by spectrophotometry (Ultraspec 2000; Amersham Pharmacia Biotech, Piscataway, NJ) in phosphate-buffered saline, pH 7.4, at a wavelength of 543 nm. The cumulative amount of rhodamine released was plotted as a function of time. Setup of Scaffold Fabrication and Drug Release. A pore size controlled collagen scaffold was fabricated using a nozzle (300 μm) connected to a three-axis robot plotting system supplemented with a cryogenic plate (30 °C). To remove the frozen ice on the surface of the cryogenic plate due to the ambient temperature (20 °C), nitrogen gas was supplied continuously to the plate. To fabricate the scaffold, an approximate scaffold dimension (20 mm 20 mm 1 mm) was selected and the ranges of pore and strut sizes were set as 250350 μm and 250 300 μm, respectively. All fabrication procedures were conducted on a clean bench. The controlled pneumatic pressure was 145 ( 3 kPa to obtain the collagen scaffold in a layer-by-layer manner. In vitro Cell Culture. Scaffolds for use with cell cultures (5 mm 5 mm 1 mm) were sterilized with 70% EtOH and UV light, and placed in culture medium overnight. MG63 cells (ATCC, Manassas, VA, USA) were used to observe the cellular behavior in the scaffolds. To perform the cell culture, we used a 24-well plate and, for each well, 500 μL of medium was used. MG63 cells were cultured in Dulbecco’s modified Eagle’s medium (Hyclone, Logan, UT, USA), supplemented with 10% fetal bovine serum (Hyclone) and 1% penicillin/streptomycin (Hyclone). Cells were maintained up to passage 14 and collected by trypsinEDTA
Figure 2. Percentage shrinkage of pore and strut after the alginate coating process, compared with the initially fabricated collagen scaffold. (Asterisk (*) denotes that P < 0.05; NS = not significant.) treatment. The cells were then seeded onto the scaffolds at a density of 5 104 cells/sample and incubated in an atmosphere of 5% CO2 at 37 °C. The medium was changed every second day. To assess the morphology of cells on the scaffolds, the cells were examined by SEM after 7 days. The cell/scaffold constructs were fixed in 2.5% glutaraldehyde and dehydrated through a graded ethanol series. Dried scaffolds were coated with gold and examined under SEM. Cell growth was 883
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Figure 3. SEM images of (a) a pure collagen scaffold fabricated using the cryogenic plotting system and magnified surface and cross section of the scaffold; (b) CAC-1 scaffold; (c) CAC-2 scaffold; (d) CAC-3 scaffold.
temperatures for collagen were described in detail in our previous work.32 The frozen diameter of the struts was closely associated with the dispenser nozzle movement speed and the temperature of the cryogenic plate. The diameter of the frozen struts decreased linearly as the nozzle movement speed increased, and a stable size of frozen struts was acquired at temperatures below 10 °C. To stably fabricate frozen collagen struts, we selected a temperature of 30 °C for the cryogenic plate, because thermal conduction from the freezing plate may be sufficient to fabricate a stable shape of the scaffold with a thickness of 2 mm. A nozzle movement speed of 5 mm/s was selected because this processing speed can minimize the size difference between the designed and experimentally obtained size of struts (see Figure 1a). The pneumatic pressure used to extrude the collagen solution was fixed at 145 ( 3 kPa. The size of the nozzle tip was 300 μm. After a layer-by-layer fabrication of struts, the collagen scaffolds were cross-linked with EDC solution and lyophilized at 76 °C (see Figure 1b). In the next step (Figure 1c), the cross-linked collagen scaffold was immersed in rhodamine solution to allow the absorption of rhodamine as a drug indicator and again lyophilized at 76 °C. After absorption of the drug, the collagen scaffolds were coated with different alginate solutions (5, 10, 20 wt % of alginate) using a syringe pump (see Figure 1d). After coating the alginate solution, the coated solution in the scaffolds was suctioned, and the suction pressure (∼20 kPa) was carefully selected to avoid the clogging of pores and the breakup of collagen struts (see Figure 1d). To cross-link the coated alginate, the alginate/collagen scaffolds were immersed in 1 wt % CaCl2 for 1 h. After cross-linking the coated alginate, the scaffolds were dried at room temperature for 24 h. The final samples were
determined by the MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide] assay (Cell Proliferation Kit I; Boehringer Mannheim, Mannheim, Germany). This assay is based on cleavage of the yellow tetrazolium salt MTT by mitochondrial dehydrogenases in viable cells to produce purple formazan crystals. Cells on the scaffold were incubated with 0.5 mg mL1 MTT for 4 h at 37 °C and the absorbance at 570 nm was measured using a microplate reader (EL800; Bio-Tek Instruments, Winnooski, VT, USA). Five samples were tested for each incubation period, and each test was performed in triplicate. Alkaline Phosphatase Activity. For MG63 cells seeded in the scaffolds for 1 and 7 days, alkaline phosphatase (ALP), which is a marker of osteoblast activity, was assayed by measuring the release of p-nitrophenol from p-nitrophenyl phosphate (p-NPP). The scaffolds seeded with osteoblast-like cells were rinsed gently with PBS and incubated in Tris buffer (10 mM, pH 7.5) containing 0.1% Triton X-100 for 10 min. Then, 100 μL of the lysate was added to 96-well tissue-culture plates containing 100 μL of p-NPP solution, which was prepared using an alkaline phosphatase kit (Procedure No. ALP-10; Sigma). In the presence of ALP, p-NPP is transformed to p-nitrophenol and inorganic phosphate. The ALP activity was determined from the absorbance at 405 nm, using a microplate reader (Spectra III; SLT-Lab Instruments, Salzburg, Austria). Statistical Analysis. All data are presented as mean values ( the standard deviation (mean ( SD). Statistical analyses consisted of singlefactor analyses of variance (ANOVA). In all analyses, P < 0.05 was taken to indicate statistical significance.
’ RESULTS AND DISCUSSION Cryogenic and Coating Process. The size differences of microsized collagen struts with various cryogenic processing 884
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washed five times with distilled water. Figure 1d shows crosssectional optical and SEM images of the alginate-coated collagen scaffold. The detailed compositions (alginate and collagen) of the final scaffolds are listed in Table 1. In the table, CAC-1, CAC-2, and CAC-3 indicate three different collagen scaffolds coated with alginate solution at weights of 5, 10, and 20 wt %, respectively. Shrinkage is an important processing parameter in the fabrication of 3D scaffolds. As shown in Figure 2, by comparing the sizes of the pure collagen and CAC scaffolds before and after the coating process, we determined that the dimensions of the coated CAC scaffolds shrank by ∼10%40%, depending on the volume fraction of coated alginate solution. The shrinkage of total dimension and pore size of final scaffolds may result from the accumulation of shrinkage in the struts within the scaffold. This local shrinkage of struts is the result of the absorption of deposited alginate solution on the surface of the highly porous collagen scaffold. The shrinkage values for the various volume percentages of alginate were used to fabricate uniform pore size and final dimension of CAC scaffolds, which should have similar pore size and final dimensions to the pure collagen scaffold. Characterization of Fabricated Scaffolds. Karageorgiou et al.33 reported that pore size was an important factor for the regeneration of bone tissue, and they suggested that a pore size above 300 μm should be used to increase bone formation through vascularization and pore sizes of 100400 μm should be used for osteoconduction.34 Thus, we designed scaffolds with pore sizes in the range of 300 ( 50 μm. Pure collagen scaffold and three different types (CAC-1, CAC-2, CAC-3) of collagen scaffolds coated with various concentrations of alginate solution are presented in Figure 3. The SEM images indicate that the cryogenically plotted collagen scaffold and CAC scaffolds were stable structures with a uniform pore size (312 ( 30 μm) and strut diameter (296 ( 25 μm). Generally, the roughened surface of the scaffold can influence initial cell attachment and proliferation.35 Wan et al.36 reported that osteoblast cells showed better adhesion to poly(L-lactic acid) (PLLA) and polystyrene (PS) surfaces roughened with nanoscale and microscale roughness (surface patterns with diameters of 0.45 and 2.2 μm) than to smooth PLLA and PS. Figure 3 shows SEM images of the surface and cross section of pure collagen and CAC scaffolds. As shown in these images, the surface roughness of CAC scaffolds decreased as the alginate volume percentage increased, because of the absorption of alginate solution into the roughened collagen struts. However, as shown in these images, although the coated alginate on the surface partially filled the mesopores/micropores of the scaffolds, the surface was still rough. The coated alginate influenced the porosity of the final CAC scaffolds. As shown in Table 1, the porosity of pure collagen scaffold was 98.5% ( 3%, while those of CAC-1, CAC-2, and CAC-3 were 92.5% ( 4%, 88.9% ( 3%, and 87.9% ( 5%, respectively. The decrease in porosity was linearly related to the increase in volume percentage of coated alginate. Figure 4a shows the Fourier transform infrared spectroscopy (FTIR) spectra before and after cross-linkage of the collagen scaffold under 50 mM EDC solution in 95% ethanol for 24 h at room temperature. In the collagen spectrum, the NH stretching vibration peak for the amide A was ∼3324 cm1 and amide I band (∼1630 cm1), amide II bands (∼1543 and 1452 cm1), and amide III bands (∼1235 and 1268 cm1) can be seen. From the spectra, the positions of peaks before and after collagen crosslinkage are located at the same points, but the areas of the
Figure 4. (a) FTIR spectra before and after crosslinkage of a collagen scaffold in EDC solution. (b) IR spectra for pure alginate and crosslinked CAC-3 scaffold with 1 wt % CaCl2 solution.
Figure 5. Increased water absorption of pure collagen and CAC scaffolds. (NS = not significant.)
infrared (IR) bands for cross-linked collagen are much smaller than those of non-cross-linked collagen. Sinokowska et al.37 reported that the positions of IR spectra for collagen cross-linked with EDC were the same as those of non-cross-linked collagen, because the secondary structure of collagen was not destroyed. In addition, because cross-linked collagen loses water bonded to the collagen, the coupled area of amide A and OH bands (3700 3100 cm1) is much smaller for cross-linked collagen. From these IR results, we concluded that the collagen was cross-linked well in the EDC solution. The cross-linked collagen scaffold was coated with alginate solution and was cross-linked with 1 wt % of CaCl2 in triple-distilled water. The IR spectra shown in Figure 4b indicate alginate carboxyl peaks near 1637 cm1 (symmetric COO stretching vibration) and 1419 cm1 (asymmetric COO stretching vibration). Mohan et al.38 reported that, to cross-link 885
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Figure 6. (a) Stressstrain curves for pure collagen scaffold and CAC scaffolds with a stretching rate of 0.5 mm/s. (b) Comparison of Young’s modulus and maximum tensile strength between pure collagen scaffold and CAC scaffolds (n = 5).
the alginate with CaCl2, the COO peak can be shifted, because of the ionic crosslinkage of COOH groups, and so the peak of COO was slightly changed, from 1637 cm1 to 1587 cm1, in our IR peaks. Water Uptake Ability of Fabricated Scaffolds. The water uptake ability of the scaffold has been referred to as its swelling ability. This has been shown to be an important factor for the absorption of body fluids and for the transfer of cell nutrients and metabolites within the scaffold. Figure 5 shows the water absorption of pure collagen and CAC scaffolds. Generally, highly porous collagen has very high water absorption capabilities, compared with other synthetic biopolymers. As shown in Figure 5, the water absorption of pure collagen scaffold was 5-fold greater than its original weight, because of its hydrophilic properties and porous structure, which can greatly influence the material’s water absorption characteristics.39 However, the degree of water uptake in the collagen scaffold was decreased as the volume fraction of the coating solution (alginate) increased. This was likely because the coated alginate solution partially filled the mesopores/micropores of the porous collagen scaffold, thus reducing the high water uptake properties of the collagen scaffold. The decrease in water absorption with increasing volume fraction of alginate in the CAC scaffolds followed an exponentially decaying curve. However, because excess water absorption could lead to a loss of mechanical integrity and the designed pore size of scaffold cannot be stably maintained during the cell culture process, 100% of water uptake for the most CAC scaffolds may be a reasonable value. Tensile Properties. Although collagen scaffolds have been used for various biomaterials, because of their excellent biological properties, their poor mechanical strength has been a major limitation restricting their use in hard tissue regeneration. To overcome this, various composite systems supplemented with high-strength synthetic polymers have been suggested and these show good mechanical strength and stiffness characteristics using small amounts of synthetic polymers. However, the use of synthetic polymers can adversely affect the biological functions of collagen, because of the low cell recognition signals of synthetic polymers. We selected alginate as a reinforcing material to improve the mechanical properties of the pure collagen scaffold, because it has good mechanical strength characteristics and shows good biocompatibility. To examine the mechanical properties of these alginate/collagen scaffolds, we measured uniaxial tensile strength (Figure 6a) in the
Table 2. Young’s Modulus of Pure Collagen and CAC Scaffolds in Dry and Wet State Young's Modulus (MPa)
a
scaffold
collagen
CAC-1
CAC-2
CAC-3
dry state
3.4 ( 0.2
14 ( 2
20 ( 2
27 ( 1
wet statea
0.1 ( 0.1
0.25 ( 0.2
0.3 ( 0.1
0.38 ( 0.1
The wetted scaffold was measured under the PBS solution at 37 °C.
dry state. The Young’s modulus and maximum tensile strength for a pure collagen scaffold and alginate/collagen scaffolds (CAC-1, CAC-2, CAC-3) were compared (Figure 6b). Young’s moduli of the pure collagen, and the CAC-1, CAC-2, and CAC-3 scaffolds were 3.4 ( 3, 14 ( 2, 20 ( 2, and 27 ( 1 MPa, respectively. The maximum tensile strength and modulus of the alginate/collagen scaffolds were increased, compared with those of the pure collagen scaffold, and the mechanical properties showed marked improvement with increasing alginate composition in the scaffold. The Young’s modulus of the alginate-coated CAC-3 scaffold increased by a maximum of ∼9-fold, compared to the pure collagen scaffold. The moduli of the scaffolds were analyzed using the simple rule of mixtures shown below, which represents a simple approximation for predicting the modulus of mixed composites. E ¼ Ec Vc + Ea Va
ð3Þ
where Ec and Ea are the Young’s moduli of pure collagen and alginate, respectively, and Vc and Va are the volume fractions of each respective component. From this equation, we can predict the Young’s modulus of the alginate/collagen scaffold. To measure the modulus of pure alginate, we fabricated a thin alginate film (thickness = 0.3 ( 0.03 mm) and cross-linked under 1 wt % of CaCl2 solution in triple-distilled water. The measured Young’s modulus of the pure alginate film was 39 ( 3 MPa. The volume fraction of alginate ranged from 32% to 75%. According to eq 3, the calculated moduli were 14.8, 24.8, and 30.1 MPa for the CAC-1, CAC-2, and CAC-3 scaffolds, respectively. The theoretically predicted moduli were consistent with the measured values, although eq 3 does not consider the dispersion uniformity of alginate or the surface interaction between alginate and collagen. 886
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Figure 7. (a) Cumulative percentage release of rhodamine from pure collagen and CAC scaffolds over time. (b) Percentage of initial burst versus CAC scaffold porosity.
To observe the mechanical properties of pure collagen and CAC scaffolds in the wet state, we tested the mechanical properties of the scaffolds after soaking them in PBS solution at 37 °C for 24 h. As shown in Table 2, the Young’s moduli were compared with those obtained under the dry state. The mechanical properties for both the pure collagen and CAC scaffolds were dramatically decreased when the scaffolds were wetted by PBS solution. However, although the Young’s modulus was highly decreased in wet-state measurement, higher alginate content in CAC scaffolds resulted in the higher Young’s modulus compared to that of the pure collagen scaffold. From these results, we can observe that the alginate content in the CAC scaffolds can be an increasing factor of mechanical properties in both dry and wet state. Release of Rhodamine from Alginate/Collagen Scaffold. Generally, scaffolds implanted in damaged areas provide a stable structure to support neo-tissue formation and good cellular responses. However, in some cases, the implanted area can be contaminated, and thus inflammatory reactions in skin and bone tissues have been major problems in the clinical application of synthetic scaffolds. There have been a number of studies using mixtures of matrix material and dispersive drug particles to address these problems. However, the drug release behavior is related to the degradation of matrix biomaterials; therefore, the release is difficult to control. Moreover, highly porous lyophilized collagen has been used as a promising biomaterial, but it was difficult to control drug release from the mixture of porous collagen and drug, because of the porous structure and ease of infiltration by the release medium.24 Here, we used an alginate coating method to overcome this rapid drug release behavior in porous collagen scaffold. Generally, although the drug size, charge and potential interactions between drugs and matrix polymers can influence drug release behavior,40 we used rhodamine-B as a drug indicator to observe qualitative drug release behavior from the alginate/collagen scaffold. The release of rhodamine from the scaffold (10 mm 10 mm 1 mm) was characterized by spectrophotometry in PBS (pH 7.4) at a wavelength of 543 nm, as a function of time. In the scaffolds, various alginate coating compositions were used (CAC-1, CAC-2, CAC-3). The rhodamine release test from the scaffolds was performed over a period of 15 days in PBS at 37 °C. Figure 7a shows the results of the rhodamine release test for pure collagen and CAC scaffolds. As shown in the figure, the
Figure 8. Cell viability for pure collagen and CAC scaffolds versus days. (Asterisk (*) denotes P < 0.05.)
rhodamine in the pure collagen scaffold was released completely within 12 h. These observations indicated that the PBS solution can readily infiltrate the porous collagen scaffold, and the initial rate of drug release was very high. However, the initial burst was markedly reduced by increasing the composition of alginate. For CAC-3, the initial burst of drug release was reduced by 85%, compared with that from the pure collagen scaffold. These observations were similar to the drug release from biodegradable microspheres. Batycky et al.41,42 reported that drug release from porous microspheres occurred in three stages: the initial burst, induction time, and continuous release. The initial burst was from the drug initially contained on the surface and in mesopores connected to the external surface of the microsphere. The release behavior of the drug in the initial burst area occurs according to the following equation:42 Md ðtÞ ¼ Md ð0Þ 1 ξ 1 expð kd t Þ ð4Þ where Md is the mass of drug present in the polymer; ξ the mass fraction of drug involved in the burst, relative to the initial drug; and kd the drug desorption rate constant. After the initial burst of drug release, the release behavior can be delayed by an induction time (ti) to allow Fickian diffusion of drug through the micropores, allowing the release of the mass of 887
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Figure 9. (a) SEM image and EDS spectra for the CAC-3 scaffold after being incubated in the medium for 7 days without cells. (b) SEM micrograph of MG63 cells cultured on a pure collagen scaffold. (c) SEM micrograph of MG63 cells cultured on the CAC-3 scaffold after 3 days. Also shwon are SEM images after 7 days of cell culture, showing proliferated cells on (d) the collagen scaffold, (e) CAC-1, (f) CAC-2, and (g) and CAC-3 scaffolds. EDS spectra show the Ca and P compositions of the cell surfaces. The spectra were taken from the magnified SEM images over the spectra.
Table 3. Calcium and Phosphate Compositions for the Cell Surface of Scaffolds
remnant drug. According to Batycky et al., the diffusion of drug after induction time (t > ti) can be determined according to the following equation:42 Md ðtÞ ¼ Md ð0Þ expAðt ti Þ ð5Þ + ξ expð kd t Þ exp Aðt ti Þ
Composition (%) collagen scaffold CAC-1 scaffold CAC-2 scaffold CAC-3 scaffold
where parameter A is directly proportional to the effective drug diffusivity and inversely proportional to the square of microparticle geometry. From eq 5, when effective drug diffusivity is zero for times t e ti, eq 5 became eq 4. With this theoretical background, the release behavior of porous particles has two bursts, but when slow desorption or rapid degradation of microparticles can cause simultaneous desorption and diffusion
P
55.6 ( 2.2
66.9 ( 5.1
70.2 ( 0.7
62.3 ( 0.5
Ca
44.4 ( 2.2
33.0 ( 5.1
29.8 ( 0.7
37.7 ( 0.5
of remnant drug, the two bursts cannot be distinguished.42 We found two clear bursts of drug release from the CAC scaffolds, which was likely because the alginate coating on the porous collagen blocked the mesopores on the surface of collagen struts and sporadically covered the drug on the surface and thus reduced 888
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Figure 10. SEM images of cross sections of cellscaffold constructs on day 7: (a) CAC-1, (b) CAC-2, and (c) CAC-3. Cells and the tissue sheets are indicated by arrows.
the initial burst. In addition, for CAC-1, CAC-2, and CAC-3, the drug induction times had similar ranges (33.5 days). This was likely because, although the percentages of the initial burst of drugs were different, because of the amount of alginate coating, the diffusion behavior of drug through the structure of micropores of collagen struts was similar for the struts of CAC-1, CAC2, and CAC-3. As shown in Figure 7b, the percentage of drug release included in the initial burst was related to the square of the porosity of the alginate/collagen scaffold. In Vitro Tests of MG63 Cells. SEM, energy-dispersive spectroscopy (EDS) analysis, and MTT assays were performed to compare the cellular responses of alginate/collagen scaffolds with pure collagen scaffold for bone tissue engineering. To observe the quantitative cell viability of the scaffolds, we performed MTT assays. Figure 8 shows the cell viability results characterized by the MTT assay for pure collagen and alginate/collagen scaffolds after 7 days. To obtain the MTT results, we used five collagen and CAC scaffolds. The optical density wavelength value in the MTT assay through which the number of viable cells could be evaluated was ∼570 nm. The cells on pure collagen and CAC scaffolds proliferated with time, but the initial cell attachment of CAC scaffolds was much lower than that of the pure collagen scaffold, because of the decreased surface roughness, as seen in the SEM images (Figure 3). However, the rate of increase of viable cells over time was much higher in the CAC scaffolds. In particular, for the CAC-3 scaffold, the number of viable cells at 7 days was greater than that on the pure collagen scaffold. Figure 9 shows SEM images of MG63 cells that proliferated on the pure collagen and CAC scaffolds after 3 and 7 days of cell culture. To observe the effect of cell-culture medium on the precipitation of calcium phosphate, we conducted control experiments in which the scaffold was incubated in the medium for 7 days without cells under the completely same procedure of the cell-culturing process. As shown in Figure 9a, the calcium and phosphate contents after the scaffold was incubated for 7 days were negligible values. Figures 9b and 9c show the SEM images of MG63 cells on the pure collagen scaffold and CAC-3 scaffold after 3 days of cell culture. For both scaffolds, the cells were sporadically filled over the pores between struts. Figures 9dg
show SEM images of cells on the pure collagen, CAC-1, CAC-2, and CAC-3 scaffolds after 7 days of cell culture, respectively. As shown in the SEM images, MG63 cells were well spread on the surfaces of struts and the proliferating cells were sufficiently packed over pores between the struts, demonstrating that both the pure collagen and CAC scaffolds provided good cell adhesion and proliferation sites. To observe the morphology of individual MG63 cells on the pure collagen and CAC scaffolds, magnified SEM images are shown on the right of the figures. Cells of similar shape were seen on the pure collagen scaffold and on the CAC scaffolds. Most of the cells on the scaffolds were aggregated, although, in some cases, the cells were sporadically attached to the strut surface. However, analysis of the cell morphology on the CAC scaffolds (CAC-2, CAC-3) showed a layer of small particles covering the cells. These particles were calcium-rich minerals, as identified by EDS. The surface chemical compositions of MG63 cells on the scaffolds were determined by EDS. The EDS spectra shown in Figure 9 were taken from the magnified cell surfaces of the SEM images. As shown in the EDS cell surface spectra of the MG63 cells on the pure collagen scaffold and CAC scaffolds, meaningful signals of calcium and phosphate for the scaffolds after 7 days of culture are shown. These results suggest that both the pure collagen and the CAC scaffolds contributed to the production of calcium and phosphate. The detailed percentages of the elements are described in Table 3. In biomedical scaffolds, 100% pore interconnectivity and high porosity are important factors to provide paths for cellular infiltration and migration in the thickness direction of the scaffold, to avoid cell adhesion only to the scaffold surface. Figure 10 shows SEM images of cross sections of CAC scaffolds, showing the cell distributions within the interior of the scaffolds after 7 days of cell culture. As seen in the images, the cells were fully infiltrated in the thickness direction of the CAC scaffolds, and some pores were evenly filled with migrating or proliferating cells. The results showed that the cells proliferated sufficiently on the scaffolds, indicating similar cellular behavior on the CAC scaffolds compared with the pure collagen scaffold. This result indicated that the CAC scaffolds provide outstanding drug delivery function, as well as good biological function, compared with the pure collagen scaffold. 889
dx.doi.org/10.1021/cm200733s |Chem. Mater. 2012, 24, 881–891
Chemistry of Materials
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substituting for various natural biomaterials for use in bone tissue engineering.
’ AUTHOR INFORMATION Corresponding Author
*Tel.: +82-62-230-7180. Fax: +82-62-236-1534. E-mail: gkim@ chosun.ac.kr.
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Figure 11. Optical density as an indicator of ALP activity of MG63 cells on pure and CAC scaffolds. (Asterisk (*) denotes P < 0.05; NS = not significant.)
ALP activity was measured to compare the effects of the scaffolds on osteoblastic differentiation by culturing MG63 cells for 1 and 7 days on pure collagen and CAC scaffolds. Similar to the results of the MTT assay, ALP activity increased gradually with time. As shown in Figure 11, the ALP activities of the MG63 cells cultured on pure collagen and CAC scaffolds did not show a significant difference with 1 day of culture, whereas after 7 days of incubation, the ALP activity of the pure collagen was significantly higher than that on the CAC scaffolds. However, for the CAC-3 scaffold, the difference in ALP activity was
