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In vitro assessment of cell penetration into porous hydroxyapatite scaffolds with a central aligned channel
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Biomaterials 25 (2004) 5507–5514
In vitro assessment of cell penetration into porous hydroxyapatite scaffolds with a central aligned channel Felicity R. Rose*, Lesley A. Cyster, David M. Grant, Colin A. Scotchford, Steven M. Howdle, Kevin M. Shakesheff School of Pharmacy, Biomaterials and Tissue Engineering, University of Nottingham, Boots Science Building University Campus, Nottingham, Nottinghamshire NG7 2RD, UK Received 4 September 2003; accepted 27 December 2003
Abstract There is a clinical need for synthetic scaffolds that promote bone regeneration. A common problem encountered when using scaffolds in tissue engineering is the rapid formation of tissue on the outer edge of the scaffold whilst the tissue in the centre becomes necrotic. To address this, the scaffold design should improve nutrient and cell transfer to the scaffold centre. In this study, hydroxyapatite scaffolds with random, open porosity (average pore size of 282711 mm, average interconnecting window size of 7274 mm) were manufactured using a modified slip-casting methodology with a single aligned channel inserted into the centre. By varying the aligned channel diameter, a series of scaffolds with channel diameters ranging from 170 to 421 mm were produced. These scaffolds were seeded with human osteosarcoma (HOS TE85) cells and cultured for 8 days. Analysis of cell penetration into the aligned channels revealed that cell coverage increased with increasing channel diameter; from 2273% in the 170 mm diameter channel to 3876% coverage in the 421 mm channel. Cell penetration into the middle section of the 421 mm diameter channel (average cell area coverage 121 103732 103 mm2) was significantly greater than that observed within the 170 mm channel (average cell area coverage 26 10376 103 mm2). In addition, the data presented demonstrates that the minimum channel (or pore) diameter required for cell penetration into such scaffolds is approximately 80 mm. These results will direct the development of scaffolds with aligned macroarchitecture for tissue engineering bone. r 2004 Elsevier Ltd. All rights reserved. Keywords: Bone tissue engineering; Hydroxyapatite; Scaffold; Macroarchitecture; Osteoblast
1. Introduction Bone loss due to trauma or disease is not only a major socio-economic burden on world healthcare systems but leads to a reduced quality of life for the patient. Current strategies for repair include bone autograft and allograft, both of which have inherent limitations such as limited supply, increased morbidity, and disease transmission potential. Biomaterial implants provide an alternative but these are also subject to limitations, such as inappropriate mechanical properties and poor integration with the surrounding tissue [1]. Therefore, the need to develop improved material structures that could provide better fillers for large defect reconstructive surgery and more appropriate biocompatible orthopae*Corresponding author. Tel.: +44-115-846-7856; fax: +44-115-9515122. E-mail address: [email protected] (F.R. Rose). 0142-9612/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2004.01.012
dic implants for non-union defects, replacement for diseased tissue, and maxillofacial surgery, is evident [2] . Using tissue engineering strategies [3], such as the in vitro generation of bone tissue and fabrication of biocompatible implants that would promote cell infiltration from surrounding tissue, alternative therapies with the potential to overcome these limitations could be generated [4]. Such alternative therapies would also decrease the need for donor tissue and with better integration into the surrounding implant site, the incidence of implant failure and the need for revision surgery will be reduced [5]. A number of materials are used in bone tissue engineering, including synthetic and natural polymers, bioglass, and a range of calcium phosphate ceramic biomaterials, such as hydroxyapatite (HA) and tri calcium phosphate [6]. The calcium phosphate ceramics employed are similar in composition to the mineral phase of natural bone and have been used in orthopaedic applications for many years [7]. The osteoconductive
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nature of the ceramics, in conjunction with their FDA approval for clinical use, are properties that have contributed to the suitability and popularity of these materials for use in bone engineering applications [8]. As well as material type, the architectural design of scaffolds is important to facilitate cell and tissue growth. The scaffold provides the general shape and structure of the tissue to be replaced and must promote cell adhesion and subsequent tissue growth by allowing the diffusion of nutrients and cells throughout the scaffold. The scaffold also has to provide some mechanical support for the growing tissue and be strong enough to withstand the forces within the implantation site [6]. Scaffold optimization involves altering the microarchitecture, such as the material crystallinity or the microporosity, and/or the macroarchitecture of the scaffold. Altering the macroarchitecture can be achieved by changing the porosity, pore size, and pore interconnectivity, to match the characteristics of the native tissue whilst retaining scaffold integrity [9]. A common problem encountered when using scaffolds in tissue engineering is the rapid formation of tissue on the outer edge of the scaffold at the expense of tissue in the centre, which becomes necrotic, which is thought to be due to limitations of cell penetration and nutrient and waste exchange [10,11]. This can be addressed by altering the culture conditions used to grow the tissue, for example using a flow perfusion culture system [12], but this is only relevant to tissue engineering in vitro. A further method of addressing this is to incorporate a design within the scaffold that will improve nutrient and cell transfer to the scaffold center, both in vitro and in vivo. Recent studies have incorporated aligned channels into the general structure of the scaffold to achieve this goal [13–17]. However, there have been no reports comparing cell infiltration into a range channel diameters, within porous scaffolds for bone engineering, to determine the optimum for this application. The aim of this study was to assess osteoblast-like cell penetration into central aligned channels of varying diameter situated within random porous HA scaffolds. These porous HA scaffolds were manufactured using a modified slip-casting methodology, seeded with human osteoblast-like cells and cultured for 8 days. Cell coverage in channels of each diameter was assessed using scanning electron microscopy (SEM) and subsequent image analysis to determine the optimum channel diameter to promote cell growth within the centre of these tissue engineering scaffolds.
2. Methods 2.1. Scaffold fabrication Porous HA ceramics were fabricated using a modified slip-casting route incorporating foaming of the slip and
setting using methylcellulose. HA powder, batch P214R was supplied by Plasma Biotal Ltd., Tideswell, Derbyshire. HA slips were dispersed using Decon 90 (Sigma; UK), Viscalex and Dispex A40 (Allied Colloids) at a concentration of 70 mg g1 of HA powder. Suspensions were prepared by adding the HA powder gradually and with continuous stirring to the correct volume of water, which contained the dispersants. The mixture was homogenised by manual stirring. When a uniform mass was achieved, mechanical agitation was applied. Further homogenisation was carried out for 5 min after addition of the methylcellulose setting agent (4% w/v Methocels, Fluka). Ceramic slips were foamed by vigorous stirring following the addition of the foaming agent Tergitol TMN10 (Sigma; UK) into the slurry. Foamed slip was poured into moulds (24-well tissue culture polystyrene plates) and immediately placed into a 65 C oven overnight to facilitate setting and drying. Green ceramics were removed from the moulds and stainless steel needles with differing diameters inserted into the ceramics such that a single centrally aligned channel was achieved in each scaffold. Green ceramics were fired in a high temperature furnace in air using a heating rate of 3 C min1 up to 250 C then held at temperature for 30 min for complete burnout of the organic components. The ceramics were then heated at the same ramp rate up to the sintering temperature of 1350 C which was held for 3 h followed by cooling to room temperature. Ceramics were washed by ultrasonic agitation in ultra pure water for 15 min prior to sterilisation by autoclaving at 120 C for 20 min. To determine the average porosity of the sintered ceramics the densities of at least 12 scaffolds were measured and compared to the theoretical density of HA (3.156 g cm3). Pore size and window size distributions were measured by taking the average of 100 measurements from SEM studies for each scaffold type. Channel diameters were also determined by taking measurements during SEM (from the top view of the scaffold). XRD spectra were taken of the powders and of the fabricated scaffolds using a Philips Siemens D500 Diffractometer. 2.2. Cell culture Human osteosarcoma cells (HOSTE85; supplied by ECACC) were maintained in monolayer culture, in a humidified atmosphere at 37 C, 5% CO2, for less than five passages in DMEM supplemented with 10% (v/v) foetal calf serum, 2 mm l-glutamine, 100 U/ml penicillin, 0.1 mg/ml streptomycin, 0.25 mg/ml amphotericin B, 1% (v/v) non-essential amino acids (NEAA, 100 stock solution), and 0.15 mg/ml ascorbic acid (HOS complete medium). Passaging and preparation of single cell suspensions for scaffold seeding was achieved by enzymatic digestion using a 0.25% (v/v) trypsin, 0.02% (w/v) EDTA solution in phosphate buffered
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saline (PBS) pH 7.4. Cell counts were assessed using Trypan blue exclusion and a haemocytometer; live cell numbers were used to determine the final cell concentration for scaffold seeding. 2.3. Scaffold seeding optimization HA scaffolds (with no aligned channel) were soaked overnight in HOS complete medium before being transferred to 24-well tissue culture plastic plates containing 1 ml HOS cell suspension (passage number 90; 4 106 cells/ml). The cell suspension was pipetted through the scaffold matrix and the plates agitated on an orbital shaker (IKAs Schuttler . MTS4, Germany) at 0, 100, or 200 rpm overnight at 5% CO2 and 37 C. Control scaffolds were agitated in the same way, but with media alone. Scaffolds were assayed for cell viability using the Alamar Bluet Assay, washed three times in PBS, and stored frozen for DNA analysis to quantify cell number. 2.4. Relative cell viability determination Relative cell viability following seeding was determined using the Alamar Bluet assay, using a method adapted from that reported by Fields and Lancaster [18]. Scaffolds were removed from culture, washed three times in PBS and incubated with 10% Alamar Bluet in Hank’s balanced salt solution without phenol red (HBSS; pH 7) for 90 min (Alamar Bluet working solution). Aliquots (100 ml) of Alamar Bluet working solution were placed in a 96-well plate and the fluorescence measured at excitation wavelength 530 nm and emission wavelength 595 nm using a fluorescence plate reader (F2 Microplate Fluorescence, Absorbance and Luminescence System, Labtech, UK). Using cell number data generated with the Hoescht DNA assay, the viability per cell was determined.
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related to cell number using the standard curve of cell number versus average fluorescence (nm). 2.6. Scaffold seeding and subsequent culture for image analysis HA scaffolds were soaked overnight in HOS complete medium before being transferred to a 24-well tissue culture plastic plate containing 1 ml HOS cell suspension (passage number 87; 4 106 cells/ml). The cell suspension was pipetted through the scaffold prior to agitation at 100 rpm on an orbital shaker (IKAs Schuttler . MTS4, Germany), overnight at 37 C, 5% CO2. The seeded scaffolds were cultured statically in 6-well tissue culture plastic plates for a period of 8 days; scaffolds were turned over every day to encourage cell growth on both sides of the scaffold. Constructs (quadruplicate) were sacrificed for image analysis at day 8; scaffolds were washed three times in PBS and fixed in 3% (v/v) glutaraldehyde at 4 C. 2.7. SEM analysis Preparation of cultured scaffolds for SEM was carried out according to the method described by Robinson and Gray [20]. In brief, fixed scaffolds were washed in PBS, cut longitudinally through the central aligned channel, and fixed further in 1 ml 1% (v/v) osmium tetroxide in PBS for 2 h. Scaffolds were washed with distilled water, dehydrated through a series of ethanol concentrations (25%, 50%, 75%, 90%, 95%, 100% (v/v) in distilled water), and chemically dried using hexamethyldisilazane (HMDS; Sigma, Poole UK). Dried scaffolds were sputter-coated with gold prior to analysis using a scanning electron microscope (Philips XL30; 10 kV). Images were taken along the length of the scaffold channel at 300 and 400 magnification for all scaffold diameters. The length of each channel was determined using the measurement facility on the SEM software.
2.5. Cell number determination 2.8. Image analysis and calculations For the determination of the number of cells within scaffolds, scaffolds were washed three times in PBS, frozen and lyophilized. HOSTE85 cell pellets (4 106 cells) were also frozen and lyophilized for use as a standard curve. Scaffolds and cell pellets were digested overnight in 1 ml papain solution (1.06 mg/ml) at 60 C as described by Kim and colleagues [19]. Serial dilutions of the digested cell pellet were prepared as standards (0, 1.25 105, 2.5 105, 5 105, 1 106, 2 106, 4 106 cells/ml). The amount of DNA within the scaffolds and standards was determined using Hoechst 33258 dye [19]. Fluorescence was measured at excitation wavelength 355 nm and emission wavelength 460 nm on a fluorescence plate reader. The amount of DNA was
Image analysis was carried out using Leica QWin image analysis software. The pixel to micron ratio was established by measuring the micron bar on a representative SEM image and this ratio used to calibrate the image analysis software. Using the channel diameter and length measurements taken from the SEM images, a measurement frame was imposed on the image to represent the area of the channel for measurement (Fig. 1B). The SEM images of the channel are a 2-D projection of the curved hemi-cylinder of the channel in cross-section. When calculating the relative cell coverage, the bottom of the channel surface should be used since this surface is perpendicular to the SEM electron
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Fig. 1. A random porous HA scaffold with a single aligned channel (421 mm) following cell culture and preparation for SEM (A). SEM pictures were taken along the length of the channel and using the channel radius and length measurements taken during SEM, a measurement frame was imposed on each image to represent the area of the channel for cell coverage measurements (B).
beam (a line parallel to the bottom of the channel). However, to maximise the data available the line was broadened to two parallel lines, at a distance equal to the radius of the channel, from the line parallel to the bottom of the channel. The area between these two lines gave a suitable compromise between a more accurate representation of cell coverage while still maintaining less than 5% error between the actual surface area and the projected area. Whilst every attempt was made to section the channels into perfect hemi-cylinders this was difficult to achieve. By using the method above, which analyses an area dependent on the actual radius of the channel rather than an arbitary projection of a sectioned channel, consistent comparisons between different channel diameters were made as this method ensures that the same ratio of total channel surface area was always analysed. The channel length was divided into three sections to allow cell coverage in the upper (uppermost during seeding), middle, and lower sections of the channel to be determined (the sum of which giving the total cell coverage). Using the image analysis software, cell coverage was measured by drawing around the cells within the measurement frame; the software then determines the cell coverage area marked by the operator. The cell area coverage is also represented as a percentage of the total channel area within the measurement frame.
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Fig. 2. XRD spectra HA powder where (—) represents pre-sintering and ( ) represents post-sintering. Both spectra diffraction patterns correspond exactly to the JCPDS reference standard 9-432 for hydroxylapatite. The broad peaks at the low 2y range correspond to the presence of some amorphous/nanocrystalline hydroxylapatite.
porosity of 85% (Fig. 1A). The pore size variability was approximately gaussian with a standard deviation of approximately 100 mm; the average pore size was 282711 mm (7SEM). The size of the interconnecting windows was measured to be 7274 mm (7SEM). Inserting needles of different diameter (800, 500, 330 and 250 mm) into the HA scaffolds in the green state generated a series of scaffolds after sintering with a single central aligned channel of diameters 421721.5, 314749.1, 19877.5, and 170730.5 mm (7SEM), respectively. XRD spectra of the HA powder taken before and after sintering confirmed the presence of crystalline HA with no evidence of decomposition of the powder to bTCP (Fig. 2). Both spectra diffraction patterns correspond exactly to the JCPDS reference standard 9432 for hydroxylapatite. The broad peaks at the low 2y range correspond to the presence of some amorphous/ nanocrystalline hydroxylapatite. Very slight changes in peak intensity can be observed. 3.2. Number of cells seeded and relative cell viability
3. Results
Optimum seeding conditions for these HA scaffolds (without an aligned channel) were determined by comparing the average number of cells seeded with the viability per cell data at 0, 100, and 200 rpm (Fig. 3). The number of cells seeded increased with increasing agitation speed but in contrast, cell viability decreased with increasing agitation, over the speeds analysed. The optimum seeding speed for this study was chosen to be 100 rpm, where the average number of cells seeded was 1.5 106 cells and the viability per cell was 0.00770.002 nm/cell (7SEM).
3.1. Scaffold specifications
3.3. Total cell coverage in central aligned channels
Porous HA scaffolds (approximately 9 mm diameter 5 mm height) were fabricated with an average
The total percentage cell coverage generally increased with increasing channel diameter ranging from 3876%
2.9. Statistical analysis Statistical significance of results was assessed using GraphPad InStat v.3.00 (GraphPad Software Inc, San Diego, CA, USA). The Tukey–Kramer multiple comparison post-test was performed.
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Fig. 3. Cell seeding optimization of random porous HA scaffolds (no aligned channel). Average number of cells seeded (—) and corresponding relative viability per cell (nm) (- - -) following seeding with agitation at 0, 100, and 200 rpm (error bars indicate SEM; n ¼ 3).
in the largest channel diameter (421 mm) to 2273% in the smallest (170 mm; 7SEM). There was no statistical significant difference in total percentage cell coverage over the range of channel diameters (Fig. 4A). Total cell area coverage decreased from 376 103740 103 to 88 10379 103 mm2 (7SEM) from the smallest to the largest channel diameter, respectively. The total cell area coverage in the largest channel was statistically significantly greater when compared with the other channel diameters (Fig. 4B). The relationship between total cell area coverage and channel diameter follows a linear relationship (Fig. 4C) which may be extrapolated to show that the minimum channel diameter required for cell penetration is 82720 mm. 3.4. Cell coverage in channel sections Cell coverage in the upper, middle and lower sections of each scaffold was assessed. In general, both the percentage cell coverage and the average cell area coverage in each section of the channel increased with increasing channel diameter (Fig. 5A and B, respectively). Cell coverage in the middle section of the channel was consistently lower than that seen in the upper and lower sections of the channel. It is important to note that the cell coverage in the middle section of the channel was greatest in the 421 mm channel (3679%; 121 103732 103 mm2; 7SEM) and was at it’s lowest in the 170 mm diameter channel (1974%; 26 1037 6 103 mm2; 7SEM). Data from each section was analysed for statistical significance, comparing like with like. There was no statistical significant difference in the average percentage cell coverage between channel sections. Using cell area data, the average cell area coverage in the upper section was statistically greater in the largest channel (163 10377 103 mm2; 7SEM) compared with that in the upper section of the other three channel diameters (98 10372 103 mm2, po0:05; 50 10377 103 mm2, po0:001; and 3 33 10 78 103 mm2, po0:001 respectively; 7SEM).
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Fig. 4. Average total percentage cell coverage (A) and average total cell area coverage (B and C) within each aligned channel of different diameter. Statistical significance using the Tukey–Kramer multiple comparison post-test po0:05; po0:01; po0:001 (error bars indicate SEM; n ¼ 4).
In the middle section, average cell coverage area in the 421 mm diameter channel (121 103732 103 mm2; 7SEM) was significantly greater than that in the middle section of the 170 mm channel only (26 1037 6 103 mm2, po0:05; 7SEM). No statistically significant difference in the average cell coverage between the lower sections was observed. This data is represented schematically in Fig. 6A and B.
4. Discussion A common problem encountered with tissue engineering strategies using scaffolds is the rapid formation of tissue on the outer edge of the scaffold whilst the tissue in the centre becomes necrotic. A way of addressing this problem is to incorporate a specific macroarchitecture design into the scaffold to improve nutrient and cell transfer to the scaffold centre. This may be achieved by inserting aligned channels into the body of porous scaffolds creating a scaffold with both a random and controlled architecture. One concern when fabricating
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Fig. 5. Average percentage cell coverage (A) and average cell area coverage (B) in the upper ( ), middle ( ), and lower ( ) sections within each aligned channel of different diameter. Statistical significance using the Tukey–Kramer multiple comparison post-test comparing average cell area coverage comparing like with like. No statistically significant difference was observed when comparing average percentage cell coverage data. Comparing upper sections, average cell area coverage (mm2) was significantly greater in the 421 mm with that in all other channel diameters po0:05 and po0:001: Comparing middle sections, average cell area coverage was significantly greater in the 421 mm than that in the smallest channel diameter po0:05: No significant differences were observed when comparing cell coverage in the lower sections of the channels (error bars indicate SEM; n ¼ 4).
such scaffolds where part of the main body of the scaffold is removed is the potential loss of mechanical strength. Subsequent studies have demonstrated that the insertion of 13 channels within these scaffolds (432 mm in diameter) actually enhanced the mechanical strength of the scaffolds (when tested in the direction of the channels) almost two-fold when compared with scaffolds with no channels (unpublished observations). It is postulated that this increase in mechanical strength is due to strut formation around the channel on insertion of the needles during fabrication. The purpose of this study was to assess osteoblast-like cell penetration into central aligned channels of different diameter, situated within random porous HA scaffolds, to determine the optimum channel diameter for this application and subsequent bone engineering projects. Initial experiments determined the optimum seeding speed for these
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Fig. 6. Schematic representation of the average percentage cell coverage (A) and average cell area coverage (B) in the upper, middle, and lower sections within each aligned channel of different diameter (Fig. 5). Shaded boxes represent percentage cell coverage ( ) 50– 41%, ( ) 40–31%, ( ) 30–21%, ( ) 20–11% and cell coverage area ( ) 200 103–151 103 mm2, ( ) 150 103–101 103, ( ) 100 103–51 103 mm2, ( ) 50,000–0 mm2. Representative SEM images from the middle section of both the 421 mm (a) and the 170 mm (b) diameter channel. Arrow heads indicate location of cells.
HA scaffolds using HOSTE85 cells. The number of cells seeded into the scaffolds increased with increasing agitation speed in contrast with cell viability, which was compromised at the higher speeds. Although cell viability was highest with a static seeding, there were concerns that not enough cells would be present for subsequent image analysis. For this reason, 100 rpm was selected as the optimum seeding speed for this study (over the speeds analyzed) and although initial cell viability was compromised, it was hypothesized that this would increase with time in culture. Overall, total cell coverage of the channel walls increased with increasing channel diameter. If channel diameter were the sole influencing factor on cell coverage, then the cell coverage area would be linearly proportional to channel diameter. This is shown in Fig. 4C. However, the graph
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line does not pass through the origin, intercepting the Xaxis at 82 mm. This suggests that below a critical channel diameter, or even pore size, of 82 mm, no cell coverage within the scaffold would be observed. A further factor that affects cell survival within tissue engineering scaffolds is nutrient flow. However, the linear relationship between cell coverage and channel diameter suggests that this was not a limiting factor in this study, as a deviation from this linear relationship would have been observed. Other studies have demonstrated that osteoblasts do not require high media flow rates [21,22] as compared with other cell types, for example chondrocytes [23]. This would explain why no such deviation was observed as the channel diameter decreased and gives confidence that the extrapolation to the X-axis is valid. Additional confirmation is provided by our observation that no cell penetration was detected within identical scaffolds containing no aligned channels. The average window size of these scaffolds was 72 mm in diameter, which is below the 82 mm cut off suggested in this study. This would explain the lack of cell infiltration into the scaffold from the central aligned channels themselves (unpublished observations). Importantly, cell coverage in the middle section of the channel increased with increasing channel diameter; the cell coverage area in this section of the largest channel being significantly greater than that in the smallest channel diameter assessed. Since Fig. 4C suggests that access of nutrients is not a rate limiting step with the conditions investigated, then the increased cell coverage in the middle section of the largest channel diameter is probably due to the initial increased cell seeding into the channel. This study may be compared to that published by Frosch and colleagues [24] where they assessed the infiltration of human osteoblast cells into a range of channel diameters (from 300 to 1000 mm) within titanium implants containing no other form of porosity. They found that cell ingrowth and subsequent matrix formation increased with increasing channel diameter up to 600 mm; however matrix production was suboptimal in the 1000 mm channel. There have been a number of reports published where aligned channels have been used for bone tissue engineering scaffolds. Lin and colleagues [13] used steel rods coated with poly (l-lactide-co-dl-lactide) that, upon removal of the rods, generated an interconnected porous polymer scaffold with axially orientated channels measuring 100 mm in diameter. Using osteoblastlike MC3T3-E1 cells to study in vitro attachment, viable cells were largely seen on the periphery of the scaffold although some viable cells were observed in the interior. However, these cell attachment studies were carried out to assess the biocompatibility of the material and did not quantify cell penetration into the scaffold. In another study, porous HA scaffolds with channels
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arranged in an orthogonal design (444 mm diameter channels; 44% porosity) and in a radial design (a central channel of 3 mm from which 366 mm diameter channels radiated to the scaffold periphery; 38% porosity) were evaluated in terms of bone regeneration using an in vivo model. Bone regeneration was observed in both designs after 9 weeks and new tissue formation followed the design of the scaffold. In the orthogonal design, new tissue formed an interconnected matrix with the scaffold material whereas tissue formation was largely seen in the central channel of the radial design. However, on assessing bone penetration into different zones of the scaffold they found that tissue production was largely seen throughout either end of both scaffold designs but was restricted to the periphery in central sections. Although bone penetration was generally higher in the orthogonal design, this was not statistically greater than that seen in the radial design [14]. Ma and Zhang [16] generated porous poly(l-lactic acid) scaffolds made using a solid–liquid phase separation technique at either a fixed temperature or uniaxial temperature gradient to produce scaffolds with isotropic and orientated macroarchitecture (with channel diameters in the range of 50–100 mm). These polymer scaffolds were assessed for in vitro tissue formation and cell growth over a 4-week period using MC3T3-E1 osteoblast-like cells. They reported that cell distribution followed the architectural features of the scaffold and that initial tissue formation was enhanced in scaffolds with orientated architecture compared to that with a random pore network. They do not comment on the extent of cell/tissue penetration into the scaffolds but as these were only 1.5 mm thick the effectiveness of such channels in a larger scaffold cannot be assessed. Aligned channels have been used in a range of other areas of tissue engineering. A common application is in nerve regeneration where such channels are used to provide spatial guidance and increased surface area for neuron growth [25–31]. In such applications, scaffolds with channel diameters ranging from 60 mm to 1.35 mm have been fabricated. Whilst the optimal pore diameter has yet to be determined for peripheral and spinal cord nerve regeneration, an optimal channel diameter for sciatic nerve regeneration has been reported to be 100 mm [31]. Survival of hepatocytes on three-dimensional synthetic polymer scaffolds manufactured with an internal network of aligned channels (800 mm diameter) has also been published [32]. Previous studies have reported the positive influence of aligned architecture within tissue engineering scaffolds on cell and tissue growth. It is likely that this occurs due to the initial increased access of cells to such areas and subsequent enhanced exchange of nutrients. As such, there will be a threshold diameter below which no such influence is observed. To date, many studies
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lack the comparison of different channel diameters within scaffolds and the effect on cell penetration into the middle of the construct. This is an aspect that will vary depending on the dimensions of the scaffold and the cell type being used. It is therefore essential to address this when planning to use such macroarchitecture to influence tissue formation.
Acknowledgements The authors would like to acknowledge the Foresight LINK consortium for funding.
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[17] Charri"ere E, Lemaitre J, Zysset Ph. Hydroxyapatite cement scaffolds with controlled macroporosity: fabrication protocol and mechanical properties. Biomaterials 2003;24:809–17. [18] Fields RD, Lancaster MV. Dual-attribute continuous monitoring of cell-proliferation cytotoxicity. Am Biotechnol Lab 1993;11:48. [19] Kim YJ, Sah RLY, Doong JYH, Grodzinsky AJ. Fluorometric assay of DNA in cartilage explants using Hoechst-33258. Anal Biochem 1988;174:168–76. [20] Robinson G, Gray T. Electron microscopy 2: practical procedures. In: Bancroft JD, Stevens A, editors. Theory and practice of histological techniques. London: Churchill Livingstone; 1996. p. 585–626. [21] Bancroft GN, Sikavitsas VI, van den Dolder J, Sheffield TL, Ambrose CG, Jansen JA, Mikos AG. Fluid flow increases mineralized matrix deposition in 3D perfusion culture of marrow stromal osteoblasts in a dose-dependent manner. PNAS 2002;99(20):12600–5. [22] Goldstein AS, Juarez TM, Helmke CD, Gustin MC, Mikos AG. Effect of convection on osteoblastic cell growth and function in biodegradable polymer foam scaffolds. Biomaterials 2001;22:1279–88. [23] Neves AA, Medcalf N, Brindle K. Functional assessment of tissue-engineered meniscal cartilage by magnetic resonance imaging and spectroscopy. Tissue Eng 2003;9(1):51–62. [24] Frosch KH, Barvencik F, Lohmann CH, Viereck V, Siggelkow H, Breme J, Dresing K, Strurmer . KM. Migration, matrix production and lamellar bone formation of human osteoblast-like cells in porous titanium implants. Cell Tissues Organs 2002;170:214–27. # [25] Schugens C, Maquet V, Grandfils C, J!erome R, Teyssie P. Biodegradable and macroporous polylactide implants for cell transplantation: 1. Preparation of macroporous polylactide supports by solid–liquid phase separation. Polymer 1996;37(6):1027–38. [26] Sundback C, Hadlock T, Cheney M, Vacanti J. Manufacture of porous polymer nerve conduits by a novel low-pressure injection molding process. Biomaterials 2003;24:819–30. [27] Maquet V, Martin D, Malgrange B, Franzen R, Schoenen J, # Moonen G, J!erome R. Peripheral nerve regeneration using bioresorbable macroporous polylactide scaffolds. J Biomed Mater Res 2000;52:639–51. [28] Maquet V, Martin D, Scholtes F, Franzen R, Schoenen J, # Moonen G, J!erome R. Poly(d,l-lactide) foams modified by poly(ethylene oxide)-block-poly(d,l-lactide) copolymers and aFGF: in vitro and in vivo evaluation for spinal cord regeneration. Biomaterials 2001;22:1137–46. [29] Hadlock T, Sundback C, Hunter D, Cheney M, Vacanti J. A polymer foam conduit seeded with Schwann cells promotes guided peripheral nerve regeneration. Tissue Eng 2000;6(2): 119–27. [30] Widmer MS, Gupta PK, Lu L, Meszlenyi RK, Evans GRD, Brandt K, Savel T, Gurlek A, Patrick CW, Mikos AG. Manufacture of porous biodegradable polymer conduits by an extrusion process for guided tissue regeneration. Biomaterials 1998;19:1945–55. . K, Bjursten [31] Zhao Q, Droot J, Laurell T, Wallman L, Lindstrom LM, Lundborg G, Montelius M, Danielsen N. Rat sciatic nerve regeneration through a micromachined silicon chip. Biomaterials 1997;18:75–80. [32] Kim SS, Utsunomiya H, Koski JA, Wu BM, Cima MJ, Sohn J, Mukai K, Griffith LG, Vacanti JP. Survival and function of hepatocytes on a novel three-dimensional synthetic biodegradable polymer scaffold with an intrinsic network of channels. Ann Surg 1998;228(1):8–13.
Biomaterials 25 (2004) 5507–5514
In vitro assessment of cell penetration into porous hydroxyapatite scaffolds with a central aligned channel Felicity R. Rose*, Lesley A. Cyster, David M. Grant, Colin A. Scotchford, Steven M. Howdle, Kevin M. Shakesheff School of Pharmacy, Biomaterials and Tissue Engineering, University of Nottingham, Boots Science Building University Campus, Nottingham, Nottinghamshire NG7 2RD, UK Received 4 September 2003; accepted 27 December 2003
Abstract There is a clinical need for synthetic scaffolds that promote bone regeneration. A common problem encountered when using scaffolds in tissue engineering is the rapid formation of tissue on the outer edge of the scaffold whilst the tissue in the centre becomes necrotic. To address this, the scaffold design should improve nutrient and cell transfer to the scaffold centre. In this study, hydroxyapatite scaffolds with random, open porosity (average pore size of 282711 mm, average interconnecting window size of 7274 mm) were manufactured using a modified slip-casting methodology with a single aligned channel inserted into the centre. By varying the aligned channel diameter, a series of scaffolds with channel diameters ranging from 170 to 421 mm were produced. These scaffolds were seeded with human osteosarcoma (HOS TE85) cells and cultured for 8 days. Analysis of cell penetration into the aligned channels revealed that cell coverage increased with increasing channel diameter; from 2273% in the 170 mm diameter channel to 3876% coverage in the 421 mm channel. Cell penetration into the middle section of the 421 mm diameter channel (average cell area coverage 121 103732 103 mm2) was significantly greater than that observed within the 170 mm channel (average cell area coverage 26 10376 103 mm2). In addition, the data presented demonstrates that the minimum channel (or pore) diameter required for cell penetration into such scaffolds is approximately 80 mm. These results will direct the development of scaffolds with aligned macroarchitecture for tissue engineering bone. r 2004 Elsevier Ltd. All rights reserved. Keywords: Bone tissue engineering; Hydroxyapatite; Scaffold; Macroarchitecture; Osteoblast
1. Introduction Bone loss due to trauma or disease is not only a major socio-economic burden on world healthcare systems but leads to a reduced quality of life for the patient. Current strategies for repair include bone autograft and allograft, both of which have inherent limitations such as limited supply, increased morbidity, and disease transmission potential. Biomaterial implants provide an alternative but these are also subject to limitations, such as inappropriate mechanical properties and poor integration with the surrounding tissue [1]. Therefore, the need to develop improved material structures that could provide better fillers for large defect reconstructive surgery and more appropriate biocompatible orthopae*Corresponding author. Tel.: +44-115-846-7856; fax: +44-115-9515122. E-mail address: [email protected] (F.R. Rose). 0142-9612/$ - see front matter r 2004 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2004.01.012
dic implants for non-union defects, replacement for diseased tissue, and maxillofacial surgery, is evident [2] . Using tissue engineering strategies [3], such as the in vitro generation of bone tissue and fabrication of biocompatible implants that would promote cell infiltration from surrounding tissue, alternative therapies with the potential to overcome these limitations could be generated [4]. Such alternative therapies would also decrease the need for donor tissue and with better integration into the surrounding implant site, the incidence of implant failure and the need for revision surgery will be reduced [5]. A number of materials are used in bone tissue engineering, including synthetic and natural polymers, bioglass, and a range of calcium phosphate ceramic biomaterials, such as hydroxyapatite (HA) and tri calcium phosphate [6]. The calcium phosphate ceramics employed are similar in composition to the mineral phase of natural bone and have been used in orthopaedic applications for many years [7]. The osteoconductive
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nature of the ceramics, in conjunction with their FDA approval for clinical use, are properties that have contributed to the suitability and popularity of these materials for use in bone engineering applications [8]. As well as material type, the architectural design of scaffolds is important to facilitate cell and tissue growth. The scaffold provides the general shape and structure of the tissue to be replaced and must promote cell adhesion and subsequent tissue growth by allowing the diffusion of nutrients and cells throughout the scaffold. The scaffold also has to provide some mechanical support for the growing tissue and be strong enough to withstand the forces within the implantation site [6]. Scaffold optimization involves altering the microarchitecture, such as the material crystallinity or the microporosity, and/or the macroarchitecture of the scaffold. Altering the macroarchitecture can be achieved by changing the porosity, pore size, and pore interconnectivity, to match the characteristics of the native tissue whilst retaining scaffold integrity [9]. A common problem encountered when using scaffolds in tissue engineering is the rapid formation of tissue on the outer edge of the scaffold at the expense of tissue in the centre, which becomes necrotic, which is thought to be due to limitations of cell penetration and nutrient and waste exchange [10,11]. This can be addressed by altering the culture conditions used to grow the tissue, for example using a flow perfusion culture system [12], but this is only relevant to tissue engineering in vitro. A further method of addressing this is to incorporate a design within the scaffold that will improve nutrient and cell transfer to the scaffold center, both in vitro and in vivo. Recent studies have incorporated aligned channels into the general structure of the scaffold to achieve this goal [13–17]. However, there have been no reports comparing cell infiltration into a range channel diameters, within porous scaffolds for bone engineering, to determine the optimum for this application. The aim of this study was to assess osteoblast-like cell penetration into central aligned channels of varying diameter situated within random porous HA scaffolds. These porous HA scaffolds were manufactured using a modified slip-casting methodology, seeded with human osteoblast-like cells and cultured for 8 days. Cell coverage in channels of each diameter was assessed using scanning electron microscopy (SEM) and subsequent image analysis to determine the optimum channel diameter to promote cell growth within the centre of these tissue engineering scaffolds.
2. Methods 2.1. Scaffold fabrication Porous HA ceramics were fabricated using a modified slip-casting route incorporating foaming of the slip and
setting using methylcellulose. HA powder, batch P214R was supplied by Plasma Biotal Ltd., Tideswell, Derbyshire. HA slips were dispersed using Decon 90 (Sigma; UK), Viscalex and Dispex A40 (Allied Colloids) at a concentration of 70 mg g1 of HA powder. Suspensions were prepared by adding the HA powder gradually and with continuous stirring to the correct volume of water, which contained the dispersants. The mixture was homogenised by manual stirring. When a uniform mass was achieved, mechanical agitation was applied. Further homogenisation was carried out for 5 min after addition of the methylcellulose setting agent (4% w/v Methocels, Fluka). Ceramic slips were foamed by vigorous stirring following the addition of the foaming agent Tergitol TMN10 (Sigma; UK) into the slurry. Foamed slip was poured into moulds (24-well tissue culture polystyrene plates) and immediately placed into a 65 C oven overnight to facilitate setting and drying. Green ceramics were removed from the moulds and stainless steel needles with differing diameters inserted into the ceramics such that a single centrally aligned channel was achieved in each scaffold. Green ceramics were fired in a high temperature furnace in air using a heating rate of 3 C min1 up to 250 C then held at temperature for 30 min for complete burnout of the organic components. The ceramics were then heated at the same ramp rate up to the sintering temperature of 1350 C which was held for 3 h followed by cooling to room temperature. Ceramics were washed by ultrasonic agitation in ultra pure water for 15 min prior to sterilisation by autoclaving at 120 C for 20 min. To determine the average porosity of the sintered ceramics the densities of at least 12 scaffolds were measured and compared to the theoretical density of HA (3.156 g cm3). Pore size and window size distributions were measured by taking the average of 100 measurements from SEM studies for each scaffold type. Channel diameters were also determined by taking measurements during SEM (from the top view of the scaffold). XRD spectra were taken of the powders and of the fabricated scaffolds using a Philips Siemens D500 Diffractometer. 2.2. Cell culture Human osteosarcoma cells (HOSTE85; supplied by ECACC) were maintained in monolayer culture, in a humidified atmosphere at 37 C, 5% CO2, for less than five passages in DMEM supplemented with 10% (v/v) foetal calf serum, 2 mm l-glutamine, 100 U/ml penicillin, 0.1 mg/ml streptomycin, 0.25 mg/ml amphotericin B, 1% (v/v) non-essential amino acids (NEAA, 100 stock solution), and 0.15 mg/ml ascorbic acid (HOS complete medium). Passaging and preparation of single cell suspensions for scaffold seeding was achieved by enzymatic digestion using a 0.25% (v/v) trypsin, 0.02% (w/v) EDTA solution in phosphate buffered
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saline (PBS) pH 7.4. Cell counts were assessed using Trypan blue exclusion and a haemocytometer; live cell numbers were used to determine the final cell concentration for scaffold seeding. 2.3. Scaffold seeding optimization HA scaffolds (with no aligned channel) were soaked overnight in HOS complete medium before being transferred to 24-well tissue culture plastic plates containing 1 ml HOS cell suspension (passage number 90; 4 106 cells/ml). The cell suspension was pipetted through the scaffold matrix and the plates agitated on an orbital shaker (IKAs Schuttler . MTS4, Germany) at 0, 100, or 200 rpm overnight at 5% CO2 and 37 C. Control scaffolds were agitated in the same way, but with media alone. Scaffolds were assayed for cell viability using the Alamar Bluet Assay, washed three times in PBS, and stored frozen for DNA analysis to quantify cell number. 2.4. Relative cell viability determination Relative cell viability following seeding was determined using the Alamar Bluet assay, using a method adapted from that reported by Fields and Lancaster [18]. Scaffolds were removed from culture, washed three times in PBS and incubated with 10% Alamar Bluet in Hank’s balanced salt solution without phenol red (HBSS; pH 7) for 90 min (Alamar Bluet working solution). Aliquots (100 ml) of Alamar Bluet working solution were placed in a 96-well plate and the fluorescence measured at excitation wavelength 530 nm and emission wavelength 595 nm using a fluorescence plate reader (F2 Microplate Fluorescence, Absorbance and Luminescence System, Labtech, UK). Using cell number data generated with the Hoescht DNA assay, the viability per cell was determined.
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related to cell number using the standard curve of cell number versus average fluorescence (nm). 2.6. Scaffold seeding and subsequent culture for image analysis HA scaffolds were soaked overnight in HOS complete medium before being transferred to a 24-well tissue culture plastic plate containing 1 ml HOS cell suspension (passage number 87; 4 106 cells/ml). The cell suspension was pipetted through the scaffold prior to agitation at 100 rpm on an orbital shaker (IKAs Schuttler . MTS4, Germany), overnight at 37 C, 5% CO2. The seeded scaffolds were cultured statically in 6-well tissue culture plastic plates for a period of 8 days; scaffolds were turned over every day to encourage cell growth on both sides of the scaffold. Constructs (quadruplicate) were sacrificed for image analysis at day 8; scaffolds were washed three times in PBS and fixed in 3% (v/v) glutaraldehyde at 4 C. 2.7. SEM analysis Preparation of cultured scaffolds for SEM was carried out according to the method described by Robinson and Gray [20]. In brief, fixed scaffolds were washed in PBS, cut longitudinally through the central aligned channel, and fixed further in 1 ml 1% (v/v) osmium tetroxide in PBS for 2 h. Scaffolds were washed with distilled water, dehydrated through a series of ethanol concentrations (25%, 50%, 75%, 90%, 95%, 100% (v/v) in distilled water), and chemically dried using hexamethyldisilazane (HMDS; Sigma, Poole UK). Dried scaffolds were sputter-coated with gold prior to analysis using a scanning electron microscope (Philips XL30; 10 kV). Images were taken along the length of the scaffold channel at 300 and 400 magnification for all scaffold diameters. The length of each channel was determined using the measurement facility on the SEM software.
2.5. Cell number determination 2.8. Image analysis and calculations For the determination of the number of cells within scaffolds, scaffolds were washed three times in PBS, frozen and lyophilized. HOSTE85 cell pellets (4 106 cells) were also frozen and lyophilized for use as a standard curve. Scaffolds and cell pellets were digested overnight in 1 ml papain solution (1.06 mg/ml) at 60 C as described by Kim and colleagues [19]. Serial dilutions of the digested cell pellet were prepared as standards (0, 1.25 105, 2.5 105, 5 105, 1 106, 2 106, 4 106 cells/ml). The amount of DNA within the scaffolds and standards was determined using Hoechst 33258 dye [19]. Fluorescence was measured at excitation wavelength 355 nm and emission wavelength 460 nm on a fluorescence plate reader. The amount of DNA was
Image analysis was carried out using Leica QWin image analysis software. The pixel to micron ratio was established by measuring the micron bar on a representative SEM image and this ratio used to calibrate the image analysis software. Using the channel diameter and length measurements taken from the SEM images, a measurement frame was imposed on the image to represent the area of the channel for measurement (Fig. 1B). The SEM images of the channel are a 2-D projection of the curved hemi-cylinder of the channel in cross-section. When calculating the relative cell coverage, the bottom of the channel surface should be used since this surface is perpendicular to the SEM electron
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beam (a line parallel to the bottom of the channel). However, to maximise the data available the line was broadened to two parallel lines, at a distance equal to the radius of the channel, from the line parallel to the bottom of the channel. The area between these two lines gave a suitable compromise between a more accurate representation of cell coverage while still maintaining less than 5% error between the actual surface area and the projected area. Whilst every attempt was made to section the channels into perfect hemi-cylinders this was difficult to achieve. By using the method above, which analyses an area dependent on the actual radius of the channel rather than an arbitary projection of a sectioned channel, consistent comparisons between different channel diameters were made as this method ensures that the same ratio of total channel surface area was always analysed. The channel length was divided into three sections to allow cell coverage in the upper (uppermost during seeding), middle, and lower sections of the channel to be determined (the sum of which giving the total cell coverage). Using the image analysis software, cell coverage was measured by drawing around the cells within the measurement frame; the software then determines the cell coverage area marked by the operator. The cell area coverage is also represented as a percentage of the total channel area within the measurement frame.
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porosity of 85% (Fig. 1A). The pore size variability was approximately gaussian with a standard deviation of approximately 100 mm; the average pore size was 282711 mm (7SEM). The size of the interconnecting windows was measured to be 7274 mm (7SEM). Inserting needles of different diameter (800, 500, 330 and 250 mm) into the HA scaffolds in the green state generated a series of scaffolds after sintering with a single central aligned channel of diameters 421721.5, 314749.1, 19877.5, and 170730.5 mm (7SEM), respectively. XRD spectra of the HA powder taken before and after sintering confirmed the presence of crystalline HA with no evidence of decomposition of the powder to bTCP (Fig. 2). Both spectra diffraction patterns correspond exactly to the JCPDS reference standard 9432 for hydroxylapatite. The broad peaks at the low 2y range correspond to the presence of some amorphous/ nanocrystalline hydroxylapatite. Very slight changes in peak intensity can be observed. 3.2. Number of cells seeded and relative cell viability
3. Results
Optimum seeding conditions for these HA scaffolds (without an aligned channel) were determined by comparing the average number of cells seeded with the viability per cell data at 0, 100, and 200 rpm (Fig. 3). The number of cells seeded increased with increasing agitation speed but in contrast, cell viability decreased with increasing agitation, over the speeds analysed. The optimum seeding speed for this study was chosen to be 100 rpm, where the average number of cells seeded was 1.5 106 cells and the viability per cell was 0.00770.002 nm/cell (7SEM).
3.1. Scaffold specifications
3.3. Total cell coverage in central aligned channels
Porous HA scaffolds (approximately 9 mm diameter 5 mm height) were fabricated with an average
The total percentage cell coverage generally increased with increasing channel diameter ranging from 3876%
2.9. Statistical analysis Statistical significance of results was assessed using GraphPad InStat v.3.00 (GraphPad Software Inc, San Diego, CA, USA). The Tukey–Kramer multiple comparison post-test was performed.
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in the largest channel diameter (421 mm) to 2273% in the smallest (170 mm; 7SEM). There was no statistical significant difference in total percentage cell coverage over the range of channel diameters (Fig. 4A). Total cell area coverage decreased from 376 103740 103 to 88 10379 103 mm2 (7SEM) from the smallest to the largest channel diameter, respectively. The total cell area coverage in the largest channel was statistically significantly greater when compared with the other channel diameters (Fig. 4B). The relationship between total cell area coverage and channel diameter follows a linear relationship (Fig. 4C) which may be extrapolated to show that the minimum channel diameter required for cell penetration is 82720 mm. 3.4. Cell coverage in channel sections Cell coverage in the upper, middle and lower sections of each scaffold was assessed. In general, both the percentage cell coverage and the average cell area coverage in each section of the channel increased with increasing channel diameter (Fig. 5A and B, respectively). Cell coverage in the middle section of the channel was consistently lower than that seen in the upper and lower sections of the channel. It is important to note that the cell coverage in the middle section of the channel was greatest in the 421 mm channel (3679%; 121 103732 103 mm2; 7SEM) and was at it’s lowest in the 170 mm diameter channel (1974%; 26 1037 6 103 mm2; 7SEM). Data from each section was analysed for statistical significance, comparing like with like. There was no statistical significant difference in the average percentage cell coverage between channel sections. Using cell area data, the average cell area coverage in the upper section was statistically greater in the largest channel (163 10377 103 mm2; 7SEM) compared with that in the upper section of the other three channel diameters (98 10372 103 mm2, po0:05; 50 10377 103 mm2, po0:001; and 3 33 10 78 103 mm2, po0:001 respectively; 7SEM).
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Fig. 4. Average total percentage cell coverage (A) and average total cell area coverage (B and C) within each aligned channel of different diameter. Statistical significance using the Tukey–Kramer multiple comparison post-test po0:05; po0:01; po0:001 (error bars indicate SEM; n ¼ 4).
In the middle section, average cell coverage area in the 421 mm diameter channel (121 103732 103 mm2; 7SEM) was significantly greater than that in the middle section of the 170 mm channel only (26 1037 6 103 mm2, po0:05; 7SEM). No statistically significant difference in the average cell coverage between the lower sections was observed. This data is represented schematically in Fig. 6A and B.
4. Discussion A common problem encountered with tissue engineering strategies using scaffolds is the rapid formation of tissue on the outer edge of the scaffold whilst the tissue in the centre becomes necrotic. A way of addressing this problem is to incorporate a specific macroarchitecture design into the scaffold to improve nutrient and cell transfer to the scaffold centre. This may be achieved by inserting aligned channels into the body of porous scaffolds creating a scaffold with both a random and controlled architecture. One concern when fabricating
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Fig. 5. Average percentage cell coverage (A) and average cell area coverage (B) in the upper ( ), middle ( ), and lower ( ) sections within each aligned channel of different diameter. Statistical significance using the Tukey–Kramer multiple comparison post-test comparing average cell area coverage comparing like with like. No statistically significant difference was observed when comparing average percentage cell coverage data. Comparing upper sections, average cell area coverage (mm2) was significantly greater in the 421 mm with that in all other channel diameters po0:05 and po0:001: Comparing middle sections, average cell area coverage was significantly greater in the 421 mm than that in the smallest channel diameter po0:05: No significant differences were observed when comparing cell coverage in the lower sections of the channels (error bars indicate SEM; n ¼ 4).
such scaffolds where part of the main body of the scaffold is removed is the potential loss of mechanical strength. Subsequent studies have demonstrated that the insertion of 13 channels within these scaffolds (432 mm in diameter) actually enhanced the mechanical strength of the scaffolds (when tested in the direction of the channels) almost two-fold when compared with scaffolds with no channels (unpublished observations). It is postulated that this increase in mechanical strength is due to strut formation around the channel on insertion of the needles during fabrication. The purpose of this study was to assess osteoblast-like cell penetration into central aligned channels of different diameter, situated within random porous HA scaffolds, to determine the optimum channel diameter for this application and subsequent bone engineering projects. Initial experiments determined the optimum seeding speed for these
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Fig. 6. Schematic representation of the average percentage cell coverage (A) and average cell area coverage (B) in the upper, middle, and lower sections within each aligned channel of different diameter (Fig. 5). Shaded boxes represent percentage cell coverage ( ) 50– 41%, ( ) 40–31%, ( ) 30–21%, ( ) 20–11% and cell coverage area ( ) 200 103–151 103 mm2, ( ) 150 103–101 103, ( ) 100 103–51 103 mm2, ( ) 50,000–0 mm2. Representative SEM images from the middle section of both the 421 mm (a) and the 170 mm (b) diameter channel. Arrow heads indicate location of cells.
HA scaffolds using HOSTE85 cells. The number of cells seeded into the scaffolds increased with increasing agitation speed in contrast with cell viability, which was compromised at the higher speeds. Although cell viability was highest with a static seeding, there were concerns that not enough cells would be present for subsequent image analysis. For this reason, 100 rpm was selected as the optimum seeding speed for this study (over the speeds analyzed) and although initial cell viability was compromised, it was hypothesized that this would increase with time in culture. Overall, total cell coverage of the channel walls increased with increasing channel diameter. If channel diameter were the sole influencing factor on cell coverage, then the cell coverage area would be linearly proportional to channel diameter. This is shown in Fig. 4C. However, the graph
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line does not pass through the origin, intercepting the Xaxis at 82 mm. This suggests that below a critical channel diameter, or even pore size, of 82 mm, no cell coverage within the scaffold would be observed. A further factor that affects cell survival within tissue engineering scaffolds is nutrient flow. However, the linear relationship between cell coverage and channel diameter suggests that this was not a limiting factor in this study, as a deviation from this linear relationship would have been observed. Other studies have demonstrated that osteoblasts do not require high media flow rates [21,22] as compared with other cell types, for example chondrocytes [23]. This would explain why no such deviation was observed as the channel diameter decreased and gives confidence that the extrapolation to the X-axis is valid. Additional confirmation is provided by our observation that no cell penetration was detected within identical scaffolds containing no aligned channels. The average window size of these scaffolds was 72 mm in diameter, which is below the 82 mm cut off suggested in this study. This would explain the lack of cell infiltration into the scaffold from the central aligned channels themselves (unpublished observations). Importantly, cell coverage in the middle section of the channel increased with increasing channel diameter; the cell coverage area in this section of the largest channel being significantly greater than that in the smallest channel diameter assessed. Since Fig. 4C suggests that access of nutrients is not a rate limiting step with the conditions investigated, then the increased cell coverage in the middle section of the largest channel diameter is probably due to the initial increased cell seeding into the channel. This study may be compared to that published by Frosch and colleagues [24] where they assessed the infiltration of human osteoblast cells into a range of channel diameters (from 300 to 1000 mm) within titanium implants containing no other form of porosity. They found that cell ingrowth and subsequent matrix formation increased with increasing channel diameter up to 600 mm; however matrix production was suboptimal in the 1000 mm channel. There have been a number of reports published where aligned channels have been used for bone tissue engineering scaffolds. Lin and colleagues [13] used steel rods coated with poly (l-lactide-co-dl-lactide) that, upon removal of the rods, generated an interconnected porous polymer scaffold with axially orientated channels measuring 100 mm in diameter. Using osteoblastlike MC3T3-E1 cells to study in vitro attachment, viable cells were largely seen on the periphery of the scaffold although some viable cells were observed in the interior. However, these cell attachment studies were carried out to assess the biocompatibility of the material and did not quantify cell penetration into the scaffold. In another study, porous HA scaffolds with channels
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arranged in an orthogonal design (444 mm diameter channels; 44% porosity) and in a radial design (a central channel of 3 mm from which 366 mm diameter channels radiated to the scaffold periphery; 38% porosity) were evaluated in terms of bone regeneration using an in vivo model. Bone regeneration was observed in both designs after 9 weeks and new tissue formation followed the design of the scaffold. In the orthogonal design, new tissue formed an interconnected matrix with the scaffold material whereas tissue formation was largely seen in the central channel of the radial design. However, on assessing bone penetration into different zones of the scaffold they found that tissue production was largely seen throughout either end of both scaffold designs but was restricted to the periphery in central sections. Although bone penetration was generally higher in the orthogonal design, this was not statistically greater than that seen in the radial design [14]. Ma and Zhang [16] generated porous poly(l-lactic acid) scaffolds made using a solid–liquid phase separation technique at either a fixed temperature or uniaxial temperature gradient to produce scaffolds with isotropic and orientated macroarchitecture (with channel diameters in the range of 50–100 mm). These polymer scaffolds were assessed for in vitro tissue formation and cell growth over a 4-week period using MC3T3-E1 osteoblast-like cells. They reported that cell distribution followed the architectural features of the scaffold and that initial tissue formation was enhanced in scaffolds with orientated architecture compared to that with a random pore network. They do not comment on the extent of cell/tissue penetration into the scaffolds but as these were only 1.5 mm thick the effectiveness of such channels in a larger scaffold cannot be assessed. Aligned channels have been used in a range of other areas of tissue engineering. A common application is in nerve regeneration where such channels are used to provide spatial guidance and increased surface area for neuron growth [25–31]. In such applications, scaffolds with channel diameters ranging from 60 mm to 1.35 mm have been fabricated. Whilst the optimal pore diameter has yet to be determined for peripheral and spinal cord nerve regeneration, an optimal channel diameter for sciatic nerve regeneration has been reported to be 100 mm [31]. Survival of hepatocytes on three-dimensional synthetic polymer scaffolds manufactured with an internal network of aligned channels (800 mm diameter) has also been published [32]. Previous studies have reported the positive influence of aligned architecture within tissue engineering scaffolds on cell and tissue growth. It is likely that this occurs due to the initial increased access of cells to such areas and subsequent enhanced exchange of nutrients. As such, there will be a threshold diameter below which no such influence is observed. To date, many studies
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lack the comparison of different channel diameters within scaffolds and the effect on cell penetration into the middle of the construct. This is an aspect that will vary depending on the dimensions of the scaffold and the cell type being used. It is therefore essential to address this when planning to use such macroarchitecture to influence tissue formation.
Acknowledgements The authors would like to acknowledge the Foresight LINK consortium for funding.
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