Hyeong
Jin Lee
and
Geun Hyung
Kim
*
Bio/Nanofluidics Laboratory, Department of Mechanical Engineering, Chosun University, Gwangju, 501-759, Korea. E-mail: gkim@chosun.ac.kr; Fax: +82-62-236-1534; Tel: +82-62-230-7180
First published on 24th July 2012
Alginate, which can be derived from brown seaweed, is a well-known anionic linear polysaccharide. Alginate has been used extensively for tissue regeneration because it accelerates epithelialization and granular tissue formation, as well as encapsulating various growth factors due to its rapid gelation in calcium chloride. Although alginate is a good candidate as a natural tissue engineering material, difficulties in processing and its low mechanical properties as a porous structure remain important limitations. In previous work, we introduced multi-layered scaffolds using natural biomaterials, mainly collagen and chitosan, which were fabricated using a cryogenic direct-plotting process. The fabricated scaffolds showed good cellular activities; however, problems with regards to mechanical properties remained due to the presence of micropores. To overcome this limitation, we developed a new fabrication process that resulted in alginate scaffolds consisting of micropores in the shell and nanopores in the core region of a single strut. These alginate scaffolds exhibited good structural stability and a Young's modulus that was increased tenfold in the dry state in comparison to alginate scaffolds with a homogeneous micropore structure. The hierarchical scaffold showed highly viable cells in vitro, as well as sufficient alkaline phosphatase activity and calcium mineralization for bone tissue regeneration in comparison to a control alginate scaffold, which was fabricated using a conventional freeze-drying method. These results suggest that alginate scaffolds with a hierarchical structure have potential for use in hard tissue regeneration.
As another bone substitute, titanium and tantalum have been widely used as bone grafts, which are biocompatible, extremely corrosion resistant, and durable and non-biodegradable with a similar elastic modulus to that of the trabecular bone.1 Despite their favorable biological properties, bio-metals are intrinsically brittle and lack bio-degradability in biological conditions and have low processability. In particular, the high stiffness of titanium can lead to problems of stress-shielding and successive implant loosening.1 For these reasons, their clinical usage has been limited.1,2
Natural biopolymers, on the other hand, have some distinct advantages over bio-metals. Their biodegradation rates and mechanical properties can be controlled to a certain range for particular applications.2 In addition, they are easily fabricated into the desired 3D shape.3,4 For these reasons, biopolymers serve as a provisional skeleton because they gradually degrade and are replaced by new bone tissue.5 Alginate, one of the natural biopolymers, is biocompatible, hydrophilic, and biodegradable under physiological conditions.6 However, one of the major problems related with alginate is its low mechanical strength, which may limit its further application as templates for tissue regeneration and shape-controllability.2
Scaffolds are required in tissue engineering because their structures affect cellular activities.7–10 According to several researchers, two-dimensional (2D) scaffolds are unable to replicate the behaviour of cells in vivo.3 For this reason, spatially designed three-dimensional (3D) scaffolds that are structurally appropriate for 3D cell culture have been investigated. The main factors affecting cell seeding and nutrient and waste transport within the scaffold were pore size, pore shape, porosity, pore interconnectivity, tortuosity, and permeability.11–15 To develop an ideal 3D spatial scaffold, various optical, electrical, mechanical, and chemical methods have been used, such as freeze-drying, electrospinning, phase separation, gas forming, and solid-freeform fabrications (e.g., melt-plotting, printing, fused deposition modelling, stereolithography and laser sintering).16 In addition, several modified techniques (thermally induced phase-separation,17 direct laser writing (DLW),18–20 a polymer-leaching method,21 a cryogenic process combined with a freeze drying method,22,23 and electrohydrodynamic processes24,25) have been suggested. However, the original goal of these tools was to fabricate a pore-controlled structure with a high affinity for seeded cells. Cryogenic processes were used to fabricate 3D scaffolds with a controlled pore structure for various natural biopolymers.26,27 Our group used this technique to create multi-layered/pore size-controlled scaffolds using collagen, chitosan, and alginate.22,23,26,27 Although the optimum pore size and shape remain debated, cryogenically designed scaffolds with a stable pore size-controlled structure are an innovative biomaterial; they have the most uniform and controllable pore size and a completely interconnected pore structure. However, despite their dramatic pore-shape controllability and high porosity, the inadequate mechanical properties of such structures are insufficient to support tissue regeneration.
In this study, we propose a cryogenic design/freeze-drying and cross-linking process to control the internal structure of alginate struts, aimed at increasing the mechanical strength of 3D porous alginate scaffolds. Alginate, which is derived from brown seaweed, is a well-known anionic linear polysaccharide composed of 1,4-linked β-D-mannuronic (M) and α-L-guluronic acid (G) residues.28 It has been used extensively in tissue regeneration because it accelerates epithelialization and granular tissue formation, as well as encapsulating various growth factors due to its rapid gelation in calcium chloride.29,30 By cryogenically processing the alginate solution, we obtained 3D multi-layered alginate scaffolds that contained micro-sized alginate struts consisting of two internal pore structures (the core and shell region of struts). During the cross-linking process, a hybrid pore structure consisting of micro-sized (surface region) and nano-sized (core region) pores within the strut were generated due to the different diffusion rates of the various CaCl2 weight fractions. The layered alginate scaffolds exhibited over 90% porosity and completely interconnected pores, and showed highly improved mechanical properties as compared to scaffolds fabricated using the previous cryogenic process. To assess the capacity of the scaffolds to be used as biomaterials, various physical and biological activities, including water uptake ability, mechanical properties in the wet and dry states, viable cells, total protein content, and alkaline phosphatase (ALP) activity, were investigated, and an alizarin red Ca2+ quantization assay with osteoblast-like cells (MG63) was performed.
Fig. 1 (a) A schematic of the cryogenic–freeze-drying process with two different procedures and (b) a schematic of the general freeze-drying process. (c) Cryogenically plotted alginate scaffolds at −20 °C. (d) Freeze-dried alginate scaffold, and (e) final FC-scaffold after cross-linking with 5 wt% CaCl2. (f) Cross-linked scaffold without freeze-drying and (g) dried final CF-scaffold. |
Fig. 2(a–c) show cross-sectional scanning electron microscope (SEM) images of the freeze-dried [Fig. 2(a)], FC- [Fig. 2(b)], and CF- [Fig. 2(c)] scaffolds. The FC-scaffold exhibited a highly rough surface, with struts of a slightly shrunken shape as compared to those of the cryogenically plotted alginate, while the struts of the CF-scaffolds were of a more stable shape.
Fig. 2 Cross-sectional SEM images of fabricated 3D alginate scaffolds. (a) SEM micrographs of a 3D alginate scaffold after freeze-drying. Cross-sectional SEM images of (b) FC- and (c) CF-scaffolds. In (c), the core region of the CF-scaffold shows a nano-sized pore structure, while in the shell region, micro-sized pores were observed. |
In a cross-sectional view of the CF-scaffold strut [Fig. 2(c)], two different pore structures can be seen, such as micro-sized pores in the shell region, similar to the cryogenically plotted struts [Fig. 2(a)]. However, nano-sized pores were found in the core region, while in the FC-scaffold, no hierarchical pore structure was evident. We believe that this was due to the different rates of CaCl2 absorption of the FC- and CF-scaffolds. In the case of the FC-scaffold, the completely porous structure of the strut can easily and quickly absorb the cross-linking agent, while the CF-scaffold cannot, since the iced region within the struts slows the wetting of the CaCl2 solution. This slowed absorption may cause the development of nano-sized pores in the core region due to melted iced alginate.
The effect of CaCl2 concentration on the absorption rate of CF-scaffolds was also characterized. Fig. 3(a) shows a schematic of the alginate cross-linking process for two different CaCl2 concentrations. We hypothesise two cross-linking mechanisms: at low CaCl2 concentrations, the frozen alginate strut is slowly cross-linked, such that iced alginate in the strut core region melts, while at high CaCl2 concentrations, the wetting rate into the strut is relatively high, such that the whole strut is rapidly cross-linked.
Fig. 3 (a) A schematic of the effects of cross-linking agent concentration on nanopore and micropore development. Cross-sectional SEM images of the CF-scaffolds after treatment with (b) 1, (c) 3, (d) 5, and (e) 10 wt% CaCl2. (f) Region: An/At ratio. An and At represent the cross-sectional area of the nano-sized pores and the total cross-sectional area of the struts, respectively. (g) Pore size, which was defined as the distance between the alginate struts, and the strut diameter of the final alginate scaffolds. |
To evaluate the effect of CaCl2 concentration on nano-sized pores in the alginate struts, we fabricated several CF-scaffolds that were cross-linked using CaCl2 solutions of various weight fractions (1, 3, 5, and 10 wt%). Fig. 3(b–e) show cross-sectional SEM images of 3D alginate scaffolds cured using various CaCl2 weight fractions. Fig. 3(b) shows a CF-scaffold cross-linked using 1 wt% CaCl2. A multi-layered alginate structure was sustained; however, the struts were extremely shrunken and the micro-sized pores on the struts were completely removed due to the presence of solubilized alginate during the cross-linking process. These effects were abrogated with 3, 5, and 10 wt% CaCl2 weight fractions [Fig. 3(c–e)]. Compared to the image in Fig. 3(b), microporous structures on strut surfaces were present, but in the core area micropore-sized alginate had melted and nano-sized pores developed. However, increasing the concentration of the cross-linking agent lessened this phenomenon. In general, the process of cross-linking alginate can occur via two competing phenomena, i.e. the dissolution of alginate and desolubilization by the formation of cross-links between Ca2+ and alginate carboxyl groups.32 At low CaCl2 concentrations, alginate dissolution is dominant in the strut core-region, while at high CaCl2 concentrations, dissolution by the cross-linking process is completely blocked. Based on this reasoning, as the concentration of CaCl2 was increased, the melted region within the alginate struts was decreased. Furthermore, in the melted region, micro-sized structures can be reorganized into nano-sized pores [Fig. 3(e)]. In Fig. 3(f), At and An are the total cross-sectional area and the area of the nanopore structures in the struts, respectively.
Detailed pore and strut sizes of the fabricated alginate scaffolds, which were cross-linked with various CaCl2 weight fractions, are shown in Fig. 3(g). The pore sizes and strut diameters of the CF-scaffolds remained constant above 5 wt% of CaCl2. For this reason, we used 5 wt% CaCl2 as the cross-linking agent to compare the physical and biological properties of the FC- and CF-scaffolds.
Fig. 4 shows Fourier transform infrared spectroscopy (FTIR) spectra before and after cross-linking. The FTIR spectra exhibited alginate carboxyl peaks near 1631 cm−1, which represent symmetric COO− stretching vibrations, and near 1419 cm−1, which represent asymmetric COO− stretching vibrations. According to Mohan and Nair,33 when alginate is cross-linked with CaCl2, the symmetric COO− stretching vibration peak can be slightly shifted because of the ionic cross-linking of the COOH groups. We observed that the COO− peak was altered slightly from 1631 to 1587 cm−1. In addition, although the surface-pore structure of the CF-scaffolds cross-linked with 1 wt% CaCl2 was completely dissolved, the remaining structure was well cross-linked.
Fig. 4 Infrared (IR) spectra of the alginate scaffolds cross-linked with various cross-linking agents (1, 3, 5, and 10 wt%). Lower graph: IR data showing alginate scaffolds before and after treatment with 5 wt% CaCl2. |
In this work, our target was to regenerate trabecular bone, which has a modulus from 38 to 130 MPa.35 The modulus is highly dependent on applied forces due to the anisotropic nature of the bone. To observe the mechanical properties of the FC- and CF-scaffolds, uniaxial tensile measurements were performed. Fig. 5(a) and (b) show the stress–strain curves for the dry and wet state conditions.
Fig. 5 Stress–strain curves of the FC- and CF-alginate scaffolds in the (a) dry and (b) wet states. (c) The Young's modulus and maximum strength of the scaffolds. (d) The shrinkage and porosity of the FC- and CF-scaffolds. |
As shown in Fig. 5(a) and (b), the maximum tensile strength and modulus of the CF-scaffold dramatically increased compared to the FC-alginate scaffold. Table 1 provides a summary of the tensile properties under the same test conditions (ambient temperature = 28 °C and stretching speed = 2 mm s−1). As shown in Fig. 5(c), most of the mechanical properties were higher in the CF-scaffold as compared with the FC-scaffold. In particular, the Young's modulus of the CF-scaffold in the dry state was around tenfold greater (2.5 in the wet state) than the FC-scaffold [Fig. 5(c)]. We hypothesise that this was due to the different pore structure of the scaffolds. In general, the mechanical properties of the structured scaffolds are affected by pore structure, strut size, pore size, and porosity.36,37 In particular, porosity can markedly affect the mechanical properties. According to Ishai and Cohen,38 porosity can be related to the elastic modulus as follows,
E(ϕ) = Eo(1 − ϕ2/3) |
Young's modulus (MPa) | Maximum tensile strength (MPa) | |||
---|---|---|---|---|
FC-scaffold | CF-scaffold | FC-scaffold | CF-scaffold | |
Dry state | 2.21 ± 0.11 | 22.69 ± 0.56 | 0.07 ± 0.02 | 0.33 ± 0.03 |
Wet state | 0.19 ± 0.01 | 0.45 ± 0.09 | 0.07 ± 0.03 | 0.23 ± 0.02 |
The mechanical properties of the CF-scaffold may be sufficient for soft tissue regeneration, but remain low compared to the modulus of trabecular bone. However, the mechanical properties may be further improved through the use of a composite system supplemented with various bioceramics, such as beta-tricalcium phosphate or hydroxyapatite. In the near future, we will fabricate mechanically improved scaffolds composed of alginate and bioceramics.
Scaffold shrinkage (%) is described in Fig. 5(d). The CF-scaffold provided a more stable size change that was similar to our initial design concept.
The FC-scaffold exhibited a higher water absorption and retention capacity than the CF-scaffold [Fig. 6(a)]. However, the differences were less than 10 and 15% for absorption and retention, respectively. Although the CF-scaffold showed a slightly lower water absorption and retention capacity, the absolute values of these properties indicate its ability to contain more than its own weight in water. Fig. 6(b and c) show the wetting behaviours of the FC- and CF-scaffolds. Water mixed with red dye (5 mL) was used to determine the absorption rate. Both the FC- and CF-scaffolds showed similarly high rates of wetting.
Fig. 6 (a) Increased water absorption (%) and retention by FC- and CF-scaffolds, and optical images of the water wettability of the (b) FC-, and (c) CF-scaffolds vs. time. |
Fig. 7 (a) MG63 viable cells after culture on the C-, FC-, and CF-scaffolds. (b) Total protein contents of the MG63 cells extracted from the alginate scaffolds over time. Asterisk (*) denotes P < 0.05; NS = not significant. |
The total protein content of the cells is shown in Fig. 7(b). There was a significant difference in the total protein content of the FC- and C- and CF-scaffolds. This trend was similar to the optical density of the viable cells at seven days, and we believe that the difference may depend on the area to which the cells can attach and proliferate.
Fig. 8(a) shows the ALP activity on the scaffolds, which is a marker of osteoblastic differentiation.42 The ALP activity was normalized to total protein content, and the value of the C-scaffold was set as 100%. The relative ALP levels of the proliferated cells on all the scaffolds were similar at seven days. However, activities on the FC- and CF-scaffolds at 11 days were markedly increased as compared with the C-scaffold. Although we are unable to explain this difference, it may be because the FC- and CF-scaffolds can provide a larger area for contact between the proliferating cells and the osteogenic factors in the culture media.
Fig. 8 (a) The alkaline phosphatase activity of the MG63 cells on the alginate scaffolds from 7 to 11 days. (b) The calcium mineralization of the scaffolds at 11 days (n = 5). (c–e) ARS staining of the mineralization of the C-, FC-, and CF-scaffolds, respectively, on day 11. Asterisk (*) denotes P < 0.05. NS = not significant. |
Fig. 8(b) shows the calcium deposition on the scaffolds after 11 days, which was characterized using Alizarin Red dye and normalized to the total protein content. The calcium deposition showed a similar trend to the ALP activity. Calcium deposition on the FC- and CF-scaffolds was higher than that on the control scaffold. Fig. 8(c–e) shows the Alizarin Red calcium assay results for the C-, FC-, and CF-scaffolds at 11 days; more intense (similar to black) Alizarin Red-S staining indicates higher calcium concentrations. In both the FC- and CF-scaffolds, sufficient calcium mineralization was achieved, while in the C-scaffold calcium mineralization was evident at the surface, but not in the cross-sectional view, due to low cell migration or proliferation to the thickest portion of the C-scaffold.
This process has three setup procedures: (1) design of a pore structure-controlled and multi-layered scaffold under cryogenic conditions (−20 °C); (2) cross-linking of this scaffold in calcium chloride solution; and (3) freeze-drying. To fabricate multi-layered alginate scaffolds, we used a cryogenic process as described previously,22,23 combined with a 3-axis robot system (DTR3-2210-T-SG; DASA Robot, Bucheon, South Korea) and dispensing system (AD-3000C; Ugin-tech, Siheung, South Korea). The pore structure was controlled by manipulating system parameters, such as the nozzle moving speed, pneumatic pressure, and the temperature (−20 °C) of the cryogenic stage. Alginate struts were obtained using a 410 μm (inner diameter) and 710 μm (outer diameter) plotting nozzle, and the moving speed of the nozzle was set at 10 mm s−1 under pneumatic pressure (100 kPa) in an extrusion system. Alginate scaffolds were placed in a freeze-dryer (SFDSM06; Samwon, Busan, South Korea) at −78 °C for three days.
Cross-linking of the alginate scaffolds was observed by Fourier transform infrared (FTIR) spectrometry (model 6700; Nicolet, West Point, PA, USA). The IR spectra represent the average of 30 scans between 500 and 4000 cm−1 at a resolution of 8 cm−1.
Water absorption ability was measured by weighing the scaffolds before and after soaking in distilled water for 2 h. The increase in water absorption was calculated as (%) = (W12h − Wo)/Wo × 100, where W12h is the weight of the scaffolds after 12 h and Wo is the original weight of the scaffold. To measure water retention, scaffolds immersed for 12 h were placed in a centrifuge tube that contained filter paper at the bottom, and then centrifuged at 500 rpm for 3 min and immediately weighed. The water retention percentage was calculated as (WR-12h − Wo)/Wo × 100, where WR-12h is the weight of the centrifuged scaffolds after 12 h.
To observe the wetting properties of the scaffolds, one droplet (5 μL) of water mixed with a red dye was carefully dropped on the surface of the scaffold, and water wetting images were captured with a digital camera.
The mechanical properties of the alginate scaffolds were evaluated by measuring the Young's modulus and maximum tensile strength.43–45 The tests were performed using scaffolds in ‘dry’ and ‘wet’ states. The specimens were cut into small strips (15 × 10 mm) and the mechanical properties of five samples were measured. The dry test was conducted in a dry state at RT. For the wetting test, the specimens were pre-wetted in PBS for 12 h and then measured in a wet state at 35 °C. Tensile strength was tested using a universal tensile machine (Top-tech 2000, Chemilab, Seoul, Korea). The stress–strain curves of the scaffolds were recorded at a stretching speed of 2 mm s−1.
The porosity of the scaffolds was obtained according to the equation,
Porosity (%) = (1 − M/ρV) × 100 |
Mineralization levels were determined by Alizarin Red-S staining in 24-well plates. MG63 cells were cultured in DMEM containing 50 μg mL−1 vitamin C and 10 mM β-glycerophosphate. The cells were then washed three times with PBS, fixed in 70% (v/v) cold ethanol (4 °C) for 1 h, and air-dried. Ethanol-fixed specimens were stained with 40 mM Alizarin Red-S (pH 4.2) for 1 h and washed three times with purified water. Specimens were then destained with 10% cetylpyridium chloride in 10 mM sodium phosphate buffer (pH 7.0) for 15 min. The optical density at 562 nm was measured using a Spectra III UV microplate reader. The ALP activity and calcium deposition were normalized to the total protein content.
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