Mechanical and Biocompatibility Properties of Sintered Titanium Powder for Mimetic 3D-Printed Bone Scaffolds.

A composite comprising Ti and NaCl powders was sintered similar to a three-dimensional (3D)-printed patient-customized artificial bone scaffold. Additionally, a proper microstructure of the mimetic scaffold and the optimum processing parameters for its development were analyzed. The mechanical properties of the metal-based porous-structured framework used as an artificial bone scaffold were an optimum replacement for the human bone. Thus, it was confirmed that patient-customized scaffolds could be manufactured via 3D printing. The 3D-printed mimetic specimens were fabricated by a powder-sintering method using Ti for the metal parts, NaCl as the pore former, and polylactic acid as the biodegradable binder. Scanning electron microscopy (SEM) images showed that pores were formed homogeneously, while X-ray computed tomography confirmed that open pores were generated. The porosity and pore size distribution were measured using a mercury porosimeter, while the flexural strength and flexural elastic modulus were calculated using the three-point bending test. Based on these measurements, a pore-former content of 15 vol % optimized the density and flexural strength to 2.52 g cm-2 and 283 MPa, respectively, similar to those of the actual iliac bone. According to the 3D-printing production method, a selective laser-sintering process was applied for the fabrication of the mimetic specimen, and it was determined that the microstructure and properties similar to those of previous metal specimens could be achieved in the as-prepared specimen. Additionally, a decellularized extracellular matrix (dECM) was used to coat the surfaces and interiors of the specimens for evaluating their biocompatibilities. SEM image analysis indicated that the adipose-derived stem cells grew evenly inside the pores of the coated specimens, as compared with the bulky Ti specimens without the dECM coating. The doubling time at 65% was measured at 72, 75, and 83 h for specimens with pore-former contents of 5, 10, and 15 vol %, respectively. The doubling time without the pore former was 116 h. As compared with the specimens without the pore former (73 h), 15% of the dECM-coated specimens showed a doubling time of 64%, measured at 47 h.

Introduction

There has been an increasing demand for three-dimensional (3D)-printed patient-customized artificial bone scaffolds for the replacement of bone structures due to bone loss;[1,2] however, a proper scaffold microstructure and the optimum processing parameters have not yet been completely realized. Currently, artificial bone scaffolds used in practical applications are primarily fabricated from ceramic materials because their chemical composition is similar to that of real bones.[3,4] However, ceramic materials have mechanical and formability-related disadvantages and are biologically unstable and incompatible.[5,6] Hence, the development of a metal-based bone scaffold without these problems requires further investigation.

Nickel, cobalt, zirconium, aluminum, titanium, and their alloys are currently being used in the fabrication of metallic artificial bone scaffolds.[7] Among them, Ti is considered the most prominent material because of its biocompatibility and excellent weight-to-strength ratio.[8−10]

In terms of manufacturing patient-customized artificial bone scaffolds, it is important to identify a suitable fabrication method based on the shape, strength, and density of the bones, defects, and patient requirements. The most widely used method for molding metals is the casting method.[11,12]

The casting method is inexpensive and allows easy fabrication; however, molded metal exhibits a microstructure that is too compact to build bone scaffolds. Compactly constructed bone scaffolds are too dense and exhibit mechanical strengths greater than those of actual bones.[13] As a result, when inserted into the body, such metal scaffolds are incompatible with the surrounding bones because of the differences in their physical properties, thus resulting in stress shielding and implant loosening at the interface.[14−16]

To compensate for these shortcomings, various methods to secure pores in the scaffold microstructure have been employed, as demonstrated in previous studies, for modulating the properties of the scaffold.[17−19] Fujibayashi et al.[20] conducted a study on controlling the density through the formation of a Ti foam structure. They also conducted a study on the production of a bioactive Ti structure by controlling the density and strength of porous Ti blocks and Ti fiber-mesh cylinders using NaOH as the pore former. Using the methods reported in previous papers, we found that it was easy to control the physical properties of the scaffold by securing a large pore. However, an artificial bone scaffold should ideally exhibit fine pores with a controllable pore distribution.[21,22] In addition, smaller artificial bone scaffolds pose limitations to the possible detailing of the fabrication process. Another disadvantage is that bone scaffolds must be produced individually, and it is inconvenient and time-consuming to mold individual shapes for customized fabrications.

Metal 3D printing can optimally control the pore size and distribution for the production of an appropriate bone scaffold.[23] In particular, 3D-printing methods using selective laser sintering can be applied to correct the defect by adjusting the output and speed of the laser source.[24] Although 3D printing is well suited for the production of fine patient-specific artificial bone scaffolds in a single process, various types of specimens must be manufactured under various conditions to understand the microstructure and physical characteristics of the artificial bone scaffold. Therefore, in this study, specimens similar to 3D-printed specimens were fabricated. Experiments using 3D-printed mimetic specimens obtained through the powder-sintering method were conducted, which allowed easy modification of the process conditions and were inexpensive, and shortened the processing time.[25,26]

Using mimetic 3D-printed specimens prepared using the powder-sintering method, we investigated the density and strength of the bone scaffold. These specimens were fabricated using sodium chloride as the pore former, which could be easily eliminated to produce specimens with various porosities.

Spotted as a promising bio ink decellularized extracellular matrix (dECM) are used as cell proliferation agent. The most important function of dECM is structural support for tissue and organ morphogenesis. In addition, it can be also improving cell proliferation and tissue regeneration in artificial bone scaffolds.[27−29]

Moreover, multiple protein adhesion domains of the dECM can localize and present soluble growth factors. In addition, tissue-specific ECMs extend distinct chemical signals into surrounding cells, triggering distinct cellular responses. A more detailed explanation and contents on dECM can be found in the review paper researched by Cho et al.[30]

The internal and external surfaces of an open-pored Ti scaffold were coated with dECM to conduct cell-proliferation experiments. We found that the biocompatibility of the scaffold was secured by improving cell proliferation and tissue regeneration around the scaffold.

In this study, mimetic 3D-printed artificial bone scaffold specimens with various densities and strengths were fabricated to test the suitability of 3D printing for the production of patient-specific bone scaffolds. In addition, we produced a biocompatible artificial bone scaffold by coating the scaffold with dECM bioink.

Material and Experiments

To verify the suitability of 3D printing for the production of biomimetic specimens, the physical properties and microstructures of the mimetic 3D-printed specimens were compared with those of actual 3D-printed metal specimens. Figure shows a schematic of the experimental procedure used in the present research.

Figure 1

Schematic of the steps in the experimental procedure.

Ti powder with particle sizes ranging from 50 to 100 μm and six types of three-base NaCl (0%, 10%, and 15% by volume and 100 μm, 100–200 μm, and 200 μm of size) as the pore former were mixed and stirred well for 24 h at 1000 rpm in a Y-shaped tube. The mixture was subsequently placed in a cemented carbide mold (9 × 35 mm). Thereafter, 3 g of polylactic acid (PLA, MakerBot flexible filament) was dissolved in acetone (30 mL), and 0.6 mL of the mixture was placed in the mold and allowed to disperse for 5 min. After 10 min of compression at 300 MPa using a compressor, the NaCl was removed and placed in deionized water for 1 d. Postdrying, the prepared specimens were processed via sintering in a horizontal furnace for 2 h at 1100 °C and a heating rate of 5 °C min–1.

Prior to 3D printing, the 3D modeling tool “Meta Sequoia” was used to model a pellet shape (9 × 35 × 3 mm) suitable for the three-point bending test. Thereafter, two types of Ti-pellet specimens, with porosities of 0% and 27%, were manufactured using a metal 3D printing machine (ORLAS Creator RA).

The porosity and density of the fabricated specimens were measured via mercury porosimetry (Quantachrome, PM33GT). Thereafter, the surface was analyzed using scanning electron microscopy (SEM) (Su-70) and X-ray computed tomography (CT) for determining the pore distribution and internal structure, respectively.

The breaking strengths of the fabricated specimens were measured for a span length of 25 mm and at a moving speed of 1 mm min–1 in the three-point flexural mode of a universal testing machine (UTM). The compressive yield strengths of the compressive specimens were measured at a moving speed of 1 mm min–1 in compressive mode. Thereafter, the breaking and compressive strengths were measured using an MTDI analysis program (Data analysis ver. 2.43).

All scaffolds were coated with dECM bioink (T&R Biofab Co. Ltd., Seoul, Korea). The dECM bioink solution was prepared using bone bioink, and the component (acetic acid) was mixed for pH adjustment. The scaffolds were coated via centrifugation at 2500 rpm and 4 °C for 3 min. The coated scaffolds were incubated at 37 °C for 45 min for gelation and subsequently freeze-dried. Adipose-derived stem cells (ADSCs), at a density of 7000 cells cm–2, were seeded into the scaffold. The doubling time was measured for each dECM-coated and uncoated specimen using the Cell Counting Kit-8 (CCK-8, Dojin Laboratories, Kumamoto, Japan). The use of ADSCs for this study was approved by Seoul St. Mary’s Hospital of Korea (KCMC06BR067).

Results and Discussion

A Ti sample was fabricated (9 × 35 × 3 mm) for the three-point bending test, as shown in Figure a. The SEM images of the specimens prepared with different volume percentages of the pore former (0, 5, 10, and 15 vol %) are shown in Figure b–e, respectively. As the amount of pore former increased, the number of pores inside the specimen increased, confirming that NaCl acted as the pore former. In addition, after observing the network of the Ti powder structure, we confirmed that the specimen produced by the compression sintering process retained its shape.

Figure 2

(a) Photograph of the porous Ti scaffold used for the three-point bending test; SEM images of Ti scaffold with (b) 0 vol % pore former, (c) 5 vol % pore former, (d) 10 vol % pore former, and (e) 15 vol % pore former. (f) Porosity and density versus the quantity of the pore former.

Figure f shows the porosity and density graph for the various pore former contents, indicating that the porosity of the specimen increased with increasing pore-former content. In the case of Ti-0% (without pore formers), the specimen exhibited 19% porosity, which is characteristic of a powder specimen, caused by the formation of fine pores with sizes <10 μm between the powder particles. As the porosity increased, the density of the specimen decreased. For Ti-15%, the actual cortical density (2 g cm–3)[31] of the specimen was approximately 2.52 g cm–3, thus indicating that the Ti-15% specimen was the most biocompatible.

The density, porosity, and physical strength were measured by a mercury porosimeter, and the results of the three-point bending test for the specimen were compared numerically with those for the natural iliac bone (Table ). It was confirmed that the microstructure and physical strength of the specimen were similar to those of the iliac bone.

Table 1

Mechanical Properties of the Ti Scaffolds Depending on the Pore Former Fraction

NaCl (vol %)	density (g cm–3)	porosity (%)	flexural strength (MPa)	flexural elastic modulus (GPa)
0	3.25	19.0	445	50
5	3.23	24.0	362	40
10	2.81	37.3	291	38
15	2.52	40.6	283	32
bone	∼2	–	103–238	17–23
bulk Ti	4.51	–	200–1000	100–110

Figure shows the flexural strength and elastic modulus of the Ti specimens in the three-point bending test. The flexural strength of Ti-0% (pure Ti specimen) (445 MPa) was higher than that of the real bone (100 to 200 MPa). The specimen appeared superior in strength; however, when there is a large difference in the physical strength at the interface between the artificial bone scaffold and natural human long bone, stress shielding occurred, and the strength of the contacted long bone gradually weakened. Conversely, the Ti-15% specimen demonstrated a flexural strength of 283 MPa and elastic modulus comparable to those of the actual long bone, thus indicating that the Ti-15% specimen exhibited compatible mechanical properties.[32] Furthermore, the results of the compressive test were similar to those of flexural strength test, as shown in Figure . These findings indicate that the artificial bone scaffold was able to withstand impacts regardless of the direction of the applied stress. Accordingly, a study of the parameters associated with 3D-printed specimens containing 15% NaCl will be conducted in the future.

Figure 3

Mechanical strength relative to the porosity of the specimens.

To ensure the validity of the mimetic 3D-printed specimens, the physical properties of the specimens fabricated by selective laser sintering (SLS) using the 3D printer ORLAS Creator RA and the mimetic 3D-printed specimens obtained by compression sintering were compared. After the process conditions were set to secure a porosity similar to that of the compression-sintered specimen, the SLS metal 3D-printed specimen was produced and compared with the compression-sintered specimen (Figures S1 and S2). The 3D-printing process carried out under various conditions will be reported in a future study.

Notwithstanding that sufficient similarity to the iliac bone was achieved by adjusting the density and strength of the specimen, additional processing steps were necessary to achieve biocompatibility in the artificial bone scaffold in this study. As such, the specimens were coated with dECM, which was also used for 3D printing, both on the internal and external surfaces of the Ti bone scaffold.

We initially secured the pore size for cell proliferation.[33,34]Figure a–c shows the adjusted pore sizes of the pore former (100 μm, 100–200 μm, and 200 μm, respectively), and pore sizes ranging from 100 to 200 μm were found to be the most advantageous for cell proliferation according to the microstructures observed in the SEM results. However, for specimens with a pore size of 200 μm, a pore size of 300 μm was likewise produced. Thus, we determined that the specimens produced with pore sizes ranging from 100 to 200 μm were the most suitable for building the bone scaffold.

Figure 4

SEM images of the Ti scaffold, depending on the size of the pores: (a) approximately 100 μm, (b) 100–200 μm, and (c) approximately 200 μm. (d) X-ray CT image of the 15 vol % 100–200 μm Ti scaffold.

For the dECM bioink, to evenly coat the interior surface of the Ti scaffold, a network of pores must be formed.[35] To confirm this, X-ray CT of the Ti-15% specimen with pore sizes ranging from 100 to 200 μm was carried out. The results showed that the pores were homogeneous and that the open pores were well-formed (Figure d).

The multilineage differentiation of dECM was excellent; hence, it was possible to secure a certain level of biocompatibility regardless of the part of the body to which the scaffold was applied. According to Cho et al.,[36] an animal-bone-derived stem cell, which induced MSCs on the treated surface, facilitated cell growth and thus exhibited desirable results. This was found to improve osteogenesis of the scaffold.[37]

In addition, the dECM bioink was evenly coated on the scaffold surface and pores.[38−40] As shown in Figure , the dECM was coated on the surface, in the pores, and inside the scaffold. As a result of ADSC proliferation in the dECM-coated scaffold, the cells grew inside the scaffold, thus obtaining pores with a size suitable for cell growth.

Figure 5

SEM images of (a, b) Ti scaffold coated with dECM and (c, d) cell cultures on the dECM-coated Ti scaffold.

The cell-doubling times based on the pore-former contents were measured for the dECM-coated and uncoated specimens (Figure ). The uncoated specimens with no pore former (0 vol %) demonstrated a doubling time of 116 h, as compared with 72, 75, and 83 h for the specimens with pore formers (5, 10, and 15 vol %, respectively). Therefore, macropores were found to affect cell proliferation. However, it was difficult to compare the significant differences obtained owing to the amount of pore former that was used. For the dECM-coated specimens, differences due to the pore size were easily observed. The doubling times were 88, 50, 47 h for 5, 10, 15 vol % pore-former contents, respectively. It was confirmed that 15% of the specimens, as compared with the specimens without voids, exhibited a doubling time of 65% regardless of the coating. Thus, it can be observed that cell proliferation can be increased by increasing the surface area of the scaffold, where cells grow due to the macropores generated upon the addition of the pore former.

Figure 6

Doubling time of ADSCs, according to the pore former content.

Further research is necessary with respect to the development of patient-specific artificial bone scaffolds via 3D printing along with the database obtained in this study. Furthermore, research on the use of stainless steel, which is inexpensive, is likewise important.

Conclusions

In this study, to apply the 3D-printing method, which is suitable for the fabrication of patient-specific scaffolds, a mimetic 3D-printed Ti scaffold was fabricated, and its mechanical properties and microstructure were investigated. The porosity of the scaffold was controlled by adjusting the amount of NaCl used as the pore former, which could readily be eliminated. Thereafter, experiments were performed, which determined that a specimen obtained using 15 vol % pore former achieved a density of 2.52 g cm–3 similar to that of natural human long bone. In addition, the optimum conditions for fabrication, indicated by a similar specimen strength of 283 MPa, were achieved. Additionally, cell-proliferation experiments were conducted by coating the surface of the fabricated scaffold with dECM, a type of bioink. In both the dECM-coated and uncoated specimens, the pore former exhibited an approximately 65% doubling time, thus indicating that macropores affected the cell-proliferation capacity. Through this study, it was possible to understand the properties of the materials required for the fabrication of a Ti scaffold via 3D printing. Furthermore, other biocompatible metals, such as stainless steel 316L, can be used with this method in future studies.