Application of high resolution DLP stereolithography for fabrication of tricalcium phosphate scaffolds for bone regeneration

Synthetic or natural polymers have been widely employed as bone TE scaffolds. However, due to their low compressive strength and stiffness, they cannot replace the load bearing function of bone [18–20]. Bioactive glasses, although stronger than most polymeric biomaterials, can be associated with poor sinterability [21] and low fracture toughness [22].

Thus, ceramic scaffolds based on calcium phosphates such as hydroxyapatite (HA) or β-tricalicum phosphate (β-TCP) are considered. They are chemically similar to the inorganic component of bone, and are known to promote osteoblast adhesion, proliferation, and formation of extracellular matrix (ECM), i.e. they are osteoconductive [23]. Bulk compressive strengths of these calcium phosphates have been reported in a wide range of 5–500 MPa [9], which hinders comparison of HA and β-TCP.

In its stoichiometric and crystalline form, HA only shows limited biodegradability, whereas β-TCP has faster dissolution kinetics [24, 25]. This rationale substantiates the use of β-TCP when manufacturing biodegradable bone TE scaffolds.

Scaffold topology can be precisely modified by employing additive manufacturing approaches in order to combine the prerequisite of permeability with desired high compressive strengths.

Commonly used methods for TCP scaffolds include extrusion based systems [26–29], where a ceramic paste is deposited through a nozzle in a predefined pattern. Solidification is followed by application of a next layer, incrementally building up the part. Besides the relatively low resolution of approximately 100–200 μm [30], certain design limitations exist due to the dimensions of the cylindrical material threads, which construct every xy-plane.

Selective laser sintering (SLS) has also been employed to fabricate TCP or biphasic calcium phosphate scaffolds [31, 32]. In this process, a powder bed is exposed to a laser beam, which locally fuses a layer of the material. Although there are less design constraints than in extrusion based techniques and superior resolutions can be attained, disadvantages such as poor surface qualities or insufficient sintering densities remain [33, 34].

TCP structures for TE have also been attained by binder jetting [35, 36]. Here, the material in the powder bed is coalesced in a layer-by-layer fashion by a liquid binder solution. Similarly to SLS, poor mechanical properties due to remaining high porosities are, however, problematic [37].

Stereolithography (SLA) can also be employed to create calcium phosphate scaffolds [38, 39]. These systems utilize a photocurable ceramic slurry, where each layer is cured by exposure to light. Subsequent thermal processing removes the organic binder phase and yields a dense specimen. The digital light processing (DLP) based exposure mechanism of the presently used system ensures simultaneous irradiation of the entire desired cross section, leading to shorter processing times when compared to laser illuminated SLA. Its main advantage over other additive manufacturing systems is higher resolution [40, 41]. DLP is being investigated extensively for manufacturing of calcium phosphate structures in bone TE [17, 39, 42] and this study provides first examples of the technology for fabrication of complex high resolution porous architectures and high strength scaffolds.

The current work focuses on the use of DLP SLA to fabricate 3D scaffolds of various porosities and geometries using β-TCP for applications in bone tissue regeneration. Four different geometries, a rectilinear Grid, a hexagonal Kagome, a Schwarz primitive and a hollow Schwarz are designed with 400 μm pores and 75% or 50% porosity. However after initial screening for design stability and mechanical properties, only a rectilinear Grid structure, a hexagonal Kagome structure are found to be reproducible and showed higher mechanical properties and thus studied in detailed (please see supplementary data available online at stacks.iop.org/BMM/14/045018/mmedia for more details).

The rectilinear grid structure and hexagonal geometries of 50% and 75% porosities were evaluated using compressive testing and µ-CT to ascertain mechanical and structural properties, respectively. Complete chemical analysis pre and post fabrication was carried out using XPS and XRD. Pre-osteoblast MC3T3 cells were used to evaluate biocompatibility of 3D printed scaffolds by DLP process and their ability to form bone in short term culture.

Out of the four different scaffold designs employed, Schwarz primitive and hollow Schwarz were not continued due to their low mechanical reproducibility and the fractures found in the Hollow Schwarz architectures after sintering (see supplementary data for more details). The rectilinear grid structure (G) and a hexagonal Kagome structure (K) were created with two different porosities; 50 vol% (G50 and K50) and 75 vol% (G75 and K75). Additionally, one non-porous bulk specimen (F0) was manufactured. Pore sizes were set to 400 μm and cubes of 6 mm edge length were built for general analysis, 5 mm cubes were used for cell culture experiments and 6 × 6 × 12 mm specimens were employed in compressive testing unless stated otherwise. The computer aided design (CAD) of the G50 and G75 structures were created in the Simpleware CAD module (Synopsis^®, Inc.) using the internal structure wizard. K50 and K75 unit cells were designed in SOLIDWORKS 2017 SP03 according to Zok et al [43] and then propagated with the linear pattern function.

A ceramic additive manufacturing system CeraFab 7500 (Lithoz GmbH, Austria), as previously described in literature [40, 41], was employed with a commercial LithaBone TCP 380D slurry. The β-TCP powder was suspended in an organic matrix consisting of acrylates and methacrylates with a radical photoinitiator absorbing in the blue visible region [44]. Thus, after selective exposure to visible light, the slurry is cured. In this DLP SLA approach, as shown in figures 1(a)–(e), a digital micromirror device acts as a dynamic mask by selectively enabling or disabling single mirrors of a large mirror array. While the pixel size is fixed to 40 μm in the lateral xy-plane by the dedicated optical system, the layer height in z-direction can be adjusted from 10 to 100 μm. For this study a layer height of 25 μm was selected due to considerations of build speed and part resolution.

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The manufactured green parts were cleaned from excess uncured slurry by application of pressurized air, post cured in a UV chamber for 2 min and then thermally processed in a Nabertherm P330 furnace. Degradation of the organic matrix and its volatilization as well as sintering was conducted in a 96 h process at temperature rates which were varied between 0.17 and 0.52 °C min⁻¹. The maximum sintering temperature of 1200 °C was held for 2 h. All scaffold designs were treated with experimentally determined scaling factors prior to manufacturing to account for shrinkage during thermal treatment of 1.267 in x, y directions and of 1.281 in z direction. Over-polymerization during curing was also taken into account by scaling factors determined for each scaffold individually, as depicted in table 1.

A Bruker Skyscan 1172 micro-CT apparatus (Department of Cell and Developmental Biology, UCL) with a 0.5 mm aluminum filter was employed. Voltage was set to 50 kV and 200 μA current was used to measure and average 4 frames with 0.5° rotation steps and voxel sizes of 9 μm. Data was reconstructed in the NRecon (Bruker microCT) and internal dimensions were measured from these raw images. Four different locations within the scaffolds (at the center and on the surface, in xy planes and in planes parallel to the z axis) were measured in quadruplicates. After processing in CT-Analyser 1.13 (Bruker-microCT), porosities were determined by considering volume of interests consisting only of complete unit cells.

SEM images were produced on a JEOL JSM-6301F system (Department of Chemistry, UCL) at a working distance of 7.5–8.0 mm and with applied voltage of 5.0 kV. To assure conductivity, samples were gold plasma coated for 1.5 min at 20 mA and connected to the carrier plate with a copper wire. Measurements of the images were taken in the ImageJ software and grain sizes were determined using the lineal intercept method according to Wurst and Nelson [46].

Bulk density measurements were carried out on the non-porous F0 structures with the Archimedean method. A Sartorius density determination kit YDK 01 was used according to the manufacturer's specifications. Briefly, 5 replicate specimens were weighed 5 times in air and subsequently 5 times when immersed in water. The density of the sintered part could then be calculated from its mass and its buoyancy according to DIN EN 1389:2004-03 [47].

XPS measurements were carried out on powdered and homogenized samples with Al Kα radiation using a ThetaProbe MKII spectrometer (Thermo Scientific, Department of Chemistry, UCL) with an electron/ion flood gun. Spectra were acquired after 2 min of argon-ion etching to remove adventitious carbon. High resolution scans were taken at pass energies of 50 eV with steps of 0.100 eV. Calibration according to C 1 s at 284.8 implemented in the CasaXPS software.

Samples were crushed by pistil and mortar and the fine, homogeneous powder was packed into a 0.3 mm capillary. Powder XRD spectra were then acquired on a Stoe STADI-P system (Department of Chemistry, UCL) in transmission mode with Cu Kα₁ irradiation. The 2 θ range was 2.000°–69.815° with 0.495° steps and scan rate was set to 10.0 s step⁻¹. Data analysis in the FullProf [48] software was carried out with starting parameters for β-TCP from [49] and for HA from [50].

An in-house ZwickRoell Z005 compression machine was used for compressive testing according to ASTM standard D695-15 [51]. Crosshead speed was 1.3 mm min⁻¹ in order to be within a representative region in alignment with speeds of 0.33–8 mm min⁻¹ used in literature [27, 37]. A pre-load of 10 N was applied prior to compression and samples were rectangular prisms with dimensions of 6 × 6 × 12 mm unless stated otherwise. At least 18 replicates were measured as fired, maximum stress at fracture was taken, and Weibull analysis was carried out according to Quinn and Quinn [52].

Murine pre-osteoblastic MC3T3-E1 cells (Subclone 4, CRL-2593; ATCC, UK) were cultured in Minimum Essential Medium without ascorbic acid (α-MEM; A10490-01; Thermo Fisher Scientific Inc., UK). 10% FCS as well as 1000 U ml⁻¹ penicillin and streptomycin supplements were added and incubation was implemented at 37 °C in 5% CO₂ atmosphere. Culture medium was renewed every 2–3 d and at 70% confluency cells were passaged in a ratio of 1:6 by detaching with a 0.05% trypsin/EDTA solution in PBS.

2.25 × 10⁵ passage 3 cells were seeded per scaffold, which had been sterilized in 70% ethanol. 30 μl of a 7.6 × 10⁶ cells ml⁻¹ suspension were administered onto the scaffolds or to the 2D control wells. After resting at 25 °C for 15 min, 400 μl of α-MEM was added and the cells were incubated for 1 h. An additional 500 μl of medium was then included and a 12 h incubation period followed. Scaffolds were then transferred to clean wells in order to exclude any cells seeded on the well floors and the medium was replaced with osteogenic medium. α-MEM was additionally supplemented with 100 nM dexamethasone, 50 μg ml⁻¹ ascorbic acid, and 10 mM β-glycerophosphate to yield the differentiation medium. Only one of the 2D conditions remained in the basal α-MEM. Subsequently, medium was changed every 2–3 d.

A Quanti-iT PicoGreen assay (P7589; Thermo Fisher Scientific Inc., UK) was utilized for DNA quantification according to the specifications. Cells were lyzed with 1% Triton X-100 (Sigma Aldrich, USA) and after addition of the reagent, diluted 1:200 in the provided puffer solution. Fluorescence was recorded with excitation at 480 nm and emission at 520 nm. For quantification, a standard curve was prepared with MC3T3-E1 cells at passage 4. Measurements were taken at time points of day 1, 7, and 14.

F-actin filaments were stained by Alexa Fluor 594 phalloidin (A12381; Thermo Fisher Scientific Inc., UK) and counterstaining of DNA was carried out by 4', 6-diamidino-2-phenylindole dihydrochloride (DAPI; D1306; Thermo Fisher Scientific Inc., UK) according to the manufacturers' specifications. An in-house ApoTome.2 fluorescence microscope (Carl Zeiss Microscopy GmbH) was utilized for imaging. Images were taken after 1, 7, and 14 d of cell culturing.

The Alkaline Phosphatase assay (ab83369; Abcam, UK) was used according to the manufacturer's specifications at 1, 7, and 14 d of cell culture. Briefly, Para-nitrophenylphosphate (pNPP) was employed as a substrate to asses ALP activity and optical density was recorded at 405 nm.

A Sircol Soluble Collagen assay (S1000; Biocolor, UK) was employed for quantification of total collagen deposition at days 7 and 14. The assay was not feasible as suggested from the manufacturer, as TCP dissolved in the isolation reagents and subsequently interfered with the measurements. An alternative protocol was used [53] where the specimens were fixed in methanol at −20 °C over night and then directly exposed to 1 ml Sircol dye for 4 h at 25 °C. After washing in PBS and in 0.1% acetic acid, images were taken with a reflective light microscope (Zeiss Stemi 508). For quantification, the dye was dissolved in 250 μl of the alkali reagent and 100 μl of the solution was placed in 96 well plates to measure absorbance at 555 nm.

All cell culture experiments were analyzed after subtraction of negative controls. Data were taken from 3 sample replicates, measured in duplicates, and given as mean ± standard deviation (SD). Statistical significance was evaluated by analysis of variance or the Student's t-test in SigmaPlot 12.0. At P < 0.005, statistical significance was acknowledged and at P < 0.001, data were declared as highly significant.

Rectilinear Grid structures and hexagonal Kagome structures were fabricated in 50 and 75 vol% porosities according to the CADs displayed in figures 2(a) and (b). While unit cell size as well as strut thickness varied across the different designs, smallest pore diameter was chosen as a fixed parameter at 400 μm. This value was selected as it is at the lower limit of pore sizes necessary to enable neovascularization [12, 13, 16]. μ-CT data, compare figure 2(c), shows all scaffolds exhibiting complete interconnected porosity and homogeneous pore architecture. Pore dimensions deviated from the desired 400 μm by less than 6% (figure 2(d)) and the unit cell dimensions (figure 2(e)) of all architectures correspond to the CAD with a relative error of less than 2%. This high accuracy was achieved due to the employment of experimentally determined scaling factors and the reproducible nature of the LCM manufacturing process. When regarding the porosity of the scaffolds in figure 2(f), it is noticeable that the Grid architectures correlate better to the desired values with maximum relative deviations 3% than the Kagome scaffolds. The latter showed deviations from specifications of up to 16%. The source of this reduced porosity could be over-curing of the organic matrix caused by scattered light during the manufacturing process. When the energy threshold needed for polymerization is exceeded, curing can occur even outside of the designated area [54]. In the more intricate Kagome architecture, this phenomenon does not scale linearly with porosity as it does for the Grid structure, necessitating a more complex approximation.

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SEM images in figures 2(g) and (h), show non-homogeneous surface with the 20 ± 2 μm ridges stemming from the layer-by-layer technique of the additive manufacturing process and grain sizes of 3.83 ± 0.48 μm are identified (figure 2(i)). Furthermore, the visually compact microstructure, as inferable from the absence of intergranular pores, is confirmed by density values of 3.050 ± 0.006 g cm⁻³ (99.5% of crystallographic density [55]) obtained from Archimedean measurements of bulk specimens. This high strut density is vital for the scaffolds' mechanical integrity. Especially for ceramics, high density can be correlated to a low amount of flaws in the microstructure, which leads to high strength materials [56, 57].

XPS spectra in figure 3(a) show high purity of the TCP material before and after the 3D printing process. All organic components introduced as the binder phase during manufacturing are removed in the thermal processing step, as indicated by disappearance of the C 1s peak. This is beneficial as it eliminates potential cytotoxicity issues often associated with (meth-)acrylate chemistry [58, 59].

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The calcium (Ca) to phosphorous (P) ratio of stoichiometric TCP is 1.50. The present material, however, exhibits a slight calcium deficiency, compare table 2. Calcium vacancies [60] or traces of magnesium substitutions [61] are common explanations of this phenomenon.

The XRD spectrum of the sintered specimens in figure 3(b) confirms high purity and crystallinity of β-TCP without traces of HA or any other phases. Benefits of the demonstrated high purity and absence of amorphous phases are not limited to considerations of cytotoxicity and mechanical integrity but are also a prerequisite for reproducible degradation rates, which are greatly dependent on stoichiometry and crystallinity [62].

Compressive strengths of the scaffolds, as displayed in the Weibull plot in figure 4(a), yielded results ranging from 6.75 to 44.7 MPa (compare table 3) with lower porosity scaffolds exhibiting higher values. The fact that the rectilinear Grid structures withstand more than twice the amount of loading than the Kagome scaffolds, could be due to inhomogeneous stress distributions in the latter architectures. When regarding the stress–strain diagrams in figure 4(b), one can see pronounced discontinuities for both Kagome structures. These so called pop-in events have been extensively characterized for components with small heterogeneous zones such as in layered composite materials where cracks can be arrested when reaching a tougher material [63, 64] and could have a similar origin in the present scaffolds when individual layers of unit cells fail under loading.

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As an indication of the reliability of the scaffolds, Weibull moduli of ≥10 would be desired for ceramic materials [57]. These slightly lower values of 3.27–7.40 (compare table 3) are, however, in accordance with values of 3–10 previously found in literature for similar TCP structures [65] and could stem from reduced reproducibility when not working on bulk specimens.

When comparing to expected compressive strength values for bone at similar densities, as plotted in figure 5, one can see that the scaffolds are not as strong as compact bone, which is often stated to be in the range of 100–180 MPa [9, 10]. However, literature values of comparable scaffolds, i.e. of pure β-TCP structures with pore sizes larger than 50 μm, also do not reach these desired strengths. To our best knowledge, especially the Grid structure has noticeably higher compressive strengths than any similar scaffolds found in literature, which range from 1 MPa [66] to 20 MPa [27], compare figure 5.

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3.4.1. Cell growth and morphology

Seeding efficiency of the murine pre-osteoblastic cells, as seen from DNA amount after 1 d of cell culture, shown in figure 6(a), was at an average of 20% ± 2% for the 3D scaffolds. While this limits the comparability to the 2D control conditions (supplementary figure 1) seeded on tissue culture plastic, the more than two-fold increase of seeding efficiency of the scaffolds when compared to the 8% of seeded cells adhering to the non-porous β-TCP cubes is advantageous.

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After seeding, cell growth could be observed until day 7 of the experiment and the cells maintained that amount at day 14, as evident from the quantification in figure 6(a). When comparing the individual scaffold types, it appears that a larger DNA content was found on higher porosity Kagome scaffolds than on lower porosity Grid scaffolds. These trends are, however, not statistically significant. It is, thus, inferable that all scaffolds are compatible to the cells, which do not only adhere and grow on the TCP support but also change their morphology over the course of the experiment, as seen when regarding stained cells in figure 6(b) (see supplementary figure 2 for all stained scaffolds).

Due to the corresponding size scale of the 20 μm layers originating from the additive manufacturing process, alignment of cell nuclei was observed along these grooves [74], as depicted in figure 6(c). Larger scale ridges originating from pixilation due to the digital micro mirror device cause alignment primarily of the actin filaments. Both of these phenomena could be advantageous to cell adhesion and could have further effects on the cell's phenotype or differentiation behavior [75].

F-actin growth into the pores, as shown in confocal fluorescence images in figure 6(d), appears to be more extensive for the hexagonal Kagome pores than for the cuboid Grid pores. Although this trend was not quantified, it is in accordance with curvature driven tissue growth previously described for MC3T3-E1 cells on HA [76].

3.4.2. ALP activity and collagen deposition

In order to preliminarily asses osteogenic behavior, ALP activity was investigated and quantified compare figure 7(a). While 2D conditions had maximum ALP activity at day 7 (compare supplementary figure 1), the maximum of the 3D scaffolds is delayed by a week and increased by a factor of >2, when normalized to DNA. The porous scaffolds had elevated activity already 1 day after seeding and ALP activity remained high throughout the entire experiment. This could be explained by inhomogeneous seeding resulting in inconsistent areas of confluency and thereby in a broader range of onset-time of differentiation. The higher ALP activity especially for high porosity scaffolds does, however, strongly indicate osteogenic differentiation.

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Total collagen deposition (figures 7(b) and (c)) indicates formation of ECM. Three times as much collagen was deposited by cells in osteogenic medium than in basal medium, as discernible from the 2D control in supplementary figure 1. Furthermore, cells on 3D porous scaffolds deposited 1.5–1.7 times as much collagen as cells in 2D wells at day 14.

Both scaffold types support cell growth until day 14 with pronounced alignment along manufacturing layers and confirmed ALP activity as well as collagen deposition. Although the hexagonal architectures appear to have larger amount of cell growth, no statistically significant differences in cellular behavior can be found to the rectilinear types. The effect of scaffold porosity is limited to the higher ALP activity in more porous structures.