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Article

3D-Printed Bioceramic Scaffolds with High Strength and High Precision

by
Huifeng Shao
1,2,3,*,
Jinyuan Shi
2,
Zhiqiang Huang
2,
Weibo Yang
3 and
Honghua Wang
3
1
State Key Laboratory of Fluid Power and Mechatronic Systems, School of Mechanical Engineering, Zhejiang University, Hangzhou 310027, China
2
School of Mechanical Engineering, Hangzhou Dianzi University, Hangzhou 310018, China
3
Zhejiang Guanlin Machinery Limited Company, Huzhou 313300, China
*
Author to whom correspondence should be addressed.
Crystals 2023, 13(7), 1061; https://doi.org/10.3390/cryst13071061
Submission received: 12 June 2023 / Revised: 27 June 2023 / Accepted: 3 July 2023 / Published: 5 July 2023
(This article belongs to the Section Inorganic Crystalline Materials)
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Versions Notes

Abstract

:
Due to the increasing cases of bone damage and bone graft demand, bone-repair technology has great social and economic benefits and the manufacturing of artificial bone implants has become a focus in the domain of regenerative therapy. Considering that the traditional manufacturing process cannot effectively control the overall size of the scaffold, the diameter and shape of micropores, and the interoperability of micropores, 3D printing technology has emerged as a focal point of research within the realm of bone tissue engineering. However, the printing accuracy of extrusion-based biological 3D printing techniques is low. In this research, we utilized three-dimensional printing technology to develop high-precision magnesium-containing silicate (CSi-Mg) scaffolds. The precision of this innovative method was scrutinized and the influence of pore size on scaffold strength was systematically analyzed. Furthermore, the influence of the pore architecture on the sidewalls of these 3D-printed scaffolds was evaluated in terms of mechanical properties. The CSi-Mg scaffold, post a 3-week immersion in a simulated body of fluid, demonstrated a high modulus of elasticity (exceeding 404 MPa) and significant compressive strength (beyond 47 MPa). Furthermore, it exhibited commendable bioactivity and biodegradability. These results suggest that the high-precision 3D-printed CSi-Mg scaffolds hold great promise for addressing challenging bone defect cases.

1. Introduction

There are a huge number of bone injury cases each year and patients who suffer from trauma, congenital bone developmental abnormalities, or bone tissue excision due to cancer treatment are prone to bone defects that exceed the body’s self-repair capacity during relevant treatments, which causes them great pain [1,2]. Meanwhile, with the aggravation of the population aging, the challenges brought by bone defects are increasingly severe. Addressing the issue of bone-defect repair urgently requires a solution [3,4]. Currently, both autogenous bone grafting and allogeneic bone grafting have shown their inherent limitations [5,6]. At the same time, the research and improvement of various artificial bone implants have driven the continuous development of the field of bone-defect repair [7,8,9]. They effectively avoid the disadvantages of autograft and allograft bone transplantation and demonstrate excellent bone-repair effects [10,11,12]. Therefore, they have been widely recognized as one of the most ideal bone-repair treatment methods at present [13].
Bioactive ceramics, a kind of bone implant material, are viewed as one of the most hopeful biomaterials within the realm of bone tissue engineering [14,15,16]. Researchers commonly apply them in bone-defect repair studies and clinical treatments through methods such as porous scaffolds, mixing with other materials, or direct cell compounding [17,18]. The superior mechanical properties of bioactive ceramic scaffolds offer substantial mechanical support to bone tissue [19]. They also possess unparalleled biological properties, with excellent biocompatibility and bioactivity that promote close integration with bone tissue [20,21]. Additionally, the bioactive ions produced by the degradation of bioactive ceramic scaffolds can promote bone tissue growth, making them one of the most promising bone-repair materials currently available. Over the past few decades, researchers have explored many bioactive scaffolds for bone repair, with bioactive silicate materials being one of the most important parts that have been extensively studied and have achieved significant research results, especially magnesium-doped calcium silicate [22,23,24].
Biologically active scaffolds with interconnected pore structures are of great significance for bone repair. Not only do pores provide room for tissue expansion but they also encourage cell adhesion, guide the direction of cell growth, and facilitate nutrient transport. Therefore, biologically active ceramic materials are usually made into porous scaffold forms to adapt to human bone tissue [25,26]. However, subtractive manufacturing technologies, such as milling and molding processes, are difficult to utilize to manufacture porous scaffolds with a controllable pore size and shape, controllable pore interconnectivity, and controllable porosity. However, the emergence of 3D printing technology perfectly overcomes the above problems [27,28]. Three-dimensional printing technology can manufacture scaffolds for bone-defect repair with a suitable pore size and pore requirements while ensuring their mechanical and biological performance [29]. In addition, additive manufacturing technology has a higher material utilization rate, can save expensive biological materials, and supports personalized customization and rapid production [30]. This technology has the advantages of a short process cycle and the ability to greatly improve the efficiency of patient treatment. In the bioprinting equipment market, bioprinting equipment based on extrusion-based bioprinting technology is the most widely used and mature technology [31]. However, the low manufacturing accuracy of extrusion-based bioprinting technology limits its development.
One of the inevitable drawbacks of conventional extrusion-based 3D printing technology is its relatively low accuracy. A common strategy to improve the precision of extrusion-based 3D printing is to use finer nozzles for deposition. However, when this manufacturing process is applied to the production of bioactive ceramic scaffolds, it is prone to nozzle clogging, uneven material flow, and low yield. Sintering is one of the steps in the manufacturing process of ceramic precursors and scaffolds. Sintering densifies the ceramic powder, thereby affecting the shrinkage rate and mechanical strength of the ceramic scaffold [32]. In our previous work, we studied the impact of different sintering temperatures on the mechanical properties of the ceramic scaffold [33]. Through in-depth research in this field, a new sintering strategy for manufacturing high-precision bioceramic scaffold active materials has been proposed, based on extrusion-based 3D printing technology.
In this study, we successfully prepared high-precision and high-strength CSi-Mg material scaffolds using a novel sintering strategy based on an extrusion-based 3D printing technique. Our research aimed to explore the effects of different slurry-filling densities and multiple layers on the mechanical properties of the high-precision CSi-Mg scaffold and further evaluate the degradation and biological performance of the scaffold. This new method effectively overcomes the low precision of extrusion-based 3D printing and we expect it to become a better choice for preparing bio-ceramic scaffolds using extrusion-based 3D printing.

2. Materials and Methods

2.1. Preparation of Sintered CSi-Mg (SCM) Powders and Unsintered CSi-Mg (UCM) Powders

We created CSi-Mg precipitates utilizing a traditional chemical precipitation method [34], followed by deionized water, ethanol washing, and overnight drying at 60 °C. The ceramic precursor obtained after drying was sintered at 920 °C for 180 min and naturally cooled. Afterward, the SCM powder was obtained by ball milling with 3.5 mm diameter zirconia balls in a planetary ball mill (MP-2L; Chishun Sci&Tech Co., Nanjing, China) for 6 h and then drying. The synthesis method of UCM powder is similar to SCM powder, except for the elimination of the ceramic precursor’s calcination step. All the chemical materials were purchased from Sinopharm Chemical Reagent Co., Ltd., Shanghai China.

2.2. Preparation of CSi-Mg Scaffolds

Initially, 2 slurries were formulated by blending 5.4 g of SCM powder and 5.4 g of UCM powder, respectively, with 4.0 g of 6% polyvinyl alcohol (PVA, Sigma-Aldrich, Shanghai, China) solution. Then, these slurries were used to prepare SCM scaffold samples and UCM scaffold samples, respectively, through extrusion-based three-dimensional printing technology in a layer-by-layer manner. The three-dimensional modeling software was employed to design scaffold models with a porous structure; the prepared slurries were filled into the 3D printer. To investigate the influence of pores on the mechanical characteristics and bone-regrowth capability of bioceramic scaffolds, two bioceramic scaffolds sporting differing lateral pore shapes were conceived using CAD software: a double-layer scaffold (0°/0°/90°/90°) and a triple-layer scaffold (0°/0°/0°/90°/90°/90°). Scaffolds with varied lateral pore structures were created by adjusting the deposition angles from 0° to 90° following one or more printing layers. The double-layer and triple-layer scaffolds were acquired by altering the deposition angles of the second and third layers, respectively. The dispensing apparatus was programmed to move at a speed of 6 mm/s, with the thickness of each layer established at 0.25 mm. Following this, in order to eliminate surplus moisture within the pores, the scaffold samples were subjected to an overnight drying process at a temperature of 70 °C. They were then sintered using the same heating program in an air atmosphere at 1140 °C, controlled by a microcontroller in a temperature furnace (heating rate of 2 °C /min; holding for 45, 60, and 150 min at 320 °C, 500 °C, and 920 °C, respectively; and ultimately maintaining the desired temperature for a period of 4 h), and naturally cooled.

2.3. Characterization of the Bioceramic Powders and Scaffolds

The phase composition of SCM and UCM powders was confirmed using an X-ray diffractometer (XRD; Rigaku Co., Tokyo, Japan) under the conditions of 40 kV/40 mA. Data were gathered within a range of 20–60°, with a step size of 0.02°/2θ and a dwell time of 1.5 s, to ascertain the crystal phases of the powder. The content of inorganic ions in SCM and UCM powders was evaluated using an Inductively Coupled Plasma Optical Emission Spectrometer (ICP-OES; Thermo Icap 6000 Series, Waltham, MA, USA). First, a layer of gold was sprayed onto the surface of the sample using an ion sputter (E-1045; Hitachi Co., Tokyo, Japan). The powders were examined under a Scanning Electron Microscope (SEM, S-4800; Tokyo, Japan) at an accelerating voltage of 10 kV. The dispersion of the particle sizes in purified water was evaluated using dynamic light scattering (DLS, Malvern Instrument 3000E, Shanghai, China) techniques. We manually measured the lines in the SEM images and the thickness of each line was obtained by calculating the average of six measurement points in the image. The shrinkage rate was determined by measuring the dimensions of the scaffold in the SEM images after sintering and comparing them to the dimensions of the scaffold that were designed in the 3D model.

2.4. Mechanical Test of Scaffolds

The compressive strength and elastic modulus of the SCM scaffold (7 mm × 7 mm × 7 mm) and UCM scaffold (7 mm × 7 mm × 7 mm) were tested parallel to the pore direction (along the vertical Z direction) on the crosshead of a universal testing machine (Instron 5566, Darmstadt, Germany) at a speed of 0.5 mm/min. The elastic modulus was calculated by analyzing the linear portion of the stress-strain curve.

2.5. Biodegradation Testing In Vitro

Scaffolds of SCM and UCM, each measuring 7 mm × 7 mm × 7 mm, were submerged in a Tris buffer solution maintained at 37 °C with an initial pH of 7.4, following a liquid-to-solid ratio of 100 mL/g. A batch of five samples was then retrieved to conduct measurements on their weight and compressive strength and to assess the levels of Ca2+, Si4+, and Mg2+ present. We also measured the pH values of the remaining soaking solution at intervals of 0, 1, 2, and 3 weeks. The compressive strength of the dehydrated scaffolds was gauged using a universal strength-testing instrument (Instron 5566, Germany). To determine the dry weight, all specimens were initially rinsed with ethanol before being dried at 70 °C for a 2-day period. The weight percentage was then computed as follows: weight (%) = wd/wo ×100%, with wo denoting the dry weight of the original sample and wd signifying the dry weight of the sample after d weeks. On a weekly basis, the concentrations of Ca2+, Si4+, and Mg2+, as well as the pH values, were measured using an inductively coupled plasma atomic emission spectrometer (ICP-ACE, Optima 2100, PerkinElmer, Waltham, MA, USA) and a specialized pH meter (FE20K, Mettler Toledo, Zurich, Switzerland) for electrolytes.

2.6. Cell Viability In Vitro

Samples measuring 10 mm in diameter and 1 mm in thickness were utilized in the cell culture. Sprague–Dawley rats were used to isolate and collect bone marrow mesenchymal stem cells (BMSCs), which were obtained from the femurs and tibiae. The cells were immersed in a proliferating medium, encompassing α-modified Eagle’s medium (αMEM) (Gibco BRL, Bethesda, MD, USA) and 10% fetal bovine serum (FBS) (Gibco BRL, Bethesda, MD, USA), and were cultivated under humid conditions at a temperature of 37 °C and a carbon dioxide concentration of 5%. Subsequently, the cells were subcultured using 0.25% trypsin (Amresco, Solon, OH, USA) and 1 mM ethylene glycol bis (2-aminoethyl ether)-N,N,N′,N′-tetraacetic acid (EGTA) (Amresco, Solon, OH, USA) and passaged at 70–80% confluency. All the BMSC experiments were performed using the third passage cultures.
The BMSCs were inoculated onto the specimens situated in 24-well plates, with a density of 1 × 104 cells for each well. Following incubation periods of 1, 4, and 7 days, a Cell-Counting Kit-8 (CCK-8, Dojindo, Kumamoto, Japan) assay was utilized to enumerate the cells and assess cell viability. The absorbance was measured using a microplate reader (Infinite F50, TECAN, Hombrechtikon, Switzerland) at 450 nm.

2.7. Statistical Analysis

All the aforementioned data were reported as the average value ± standard deviation (SD) and were subjected to one-way analysis of variance (ANOVA) scrutiny. In all instances, outcomes were deemed statistically significant when the p-value was less than 0.05.

3. Results and Discussion

3.1. Primary Characterization of the Bioceramic Powders and Scaffolds

The manufacturing process of a SCM scaffold is illustrated in Figure 1A while the process for producing a high-precision UCM scaffold is depicted in Figure 1B. The typical fabrication process of a SCM scaffold involves the sintering of ceramic precursors. We propose a novel approach for producing high-precision CSi-Mg scaffolds that involves eliminating the step of sintering ceramic precursors. This method is not only applicable to CSi-Mg scaffolds but can also be applied to other inorganic ceramic materials.
Figure 2A shows the XRD patterns of crushed powders of the SCM and UCM scaffolds after sintering at 1140 °C. The XRD data reveals that both the SCM and UCM scaffold powder samples exhibit high crystallinity. Their XRD spectra coincide with wollastonite 2M (β-CSi; PDF# 27-0088), which indicates the existence of a pure wollastonite phase. Scanning Electron Microscopy (SEM) observations show that the particle sizes of the SCM and UCM powders, both pre- and post-calcination, were under 5 μm (Figure 2B,C). It is noteworthy that the Mg content measured in the CSi-Mg powder was 2.11 wt%, aligning closely with the theoretical value (2.12 wt%) computed based on stoichiometry, where 10% of Ca in wollastonite is replaced by Mg. Particle size distribution was measured using Dynamic Light Scattering. Figure 2D,E show that the particle sizes of SCM powder and UCM powder are narrowly distributed within the range of 900–1700 nm. In Figure 2A, we can see that the compositions of both the SCM and UCM powders are essentially CSi-Mg materials. In our previous work, using different sintering strategies could lead to phase transitions in wollastonite scaffolds [19]. Therefore, we maintained a dwell time of 150 min at 920 °C during the sintering of the scaffolds to prevent possible changes in the material components that could be caused by rapid heating. This ensured that the composition of the UCM scaffolds would not change due to rapid heating.
Figure 3 shows the surface and cross-sectional views of the SCM scaffold and UCM scaffold. The scaffold was manufactured using nozzle diameters of 0.4 mm, 0.5 mm, and 0.6 mm. As the nozzle diameter increased, the lines of the scaffold became thicker. Among the nozzle sizes, when using a nozzle diameter of 0.4 mm, the lines of the SCM scaffold were the thinnest, with a line diameter of approximately 279.1 ± 18.9 μm and a shrinkage rate of about 69.8 ± 4.7%. The line diameter of the UCM scaffold was approximately 190.3 ± 11.7 μm and its shrinkage rate was about 47.6 ± 2.9%. When the nozzle diameter was 0.5mm, the scaffold line width was approximately 300.8 ± 17.5 μm with a shrinkage rate of 60.2 ± 3.5% for the SCM scaffold; there was a line width of approximately 218.4 ± 15.3 μm with a shrinkage rate of 43.7 ± 3.1% for the UCM scaffold. When using a nozzle diameter of 0.6 mm, the scaffold lines were the thickest, with a line diameter of approximately 315.0 ± 17.9 μm for the SCM scaffold and a shrinkage rate of about 52.5 ± 3.0%. The line diameter of the UCM scaffold was approximately 233.5 ± 16.7 μm and its shrinkage rate was about 38.9 ± 2.8%.
We can see in Figure 3 that the lines of the UCM scaffold are finer than those of the SCM scaffold, which is due to the higher shrinkage of the UCM scaffold. This is advantageous for achieving high-precision manufacturing using extrusion-based 3D printing technology. This confirms our hypothesis. Given that ceramic scaffolds shrink after sintering [20], we hypothesized that changing the sintering strategy of the ceramic precursor may influence the shrinkage of the scaffold. Our idea is to achieve greater shrinkage and, thus, enhance precision by canceling the sintering step of the ceramic. That is, we would convert the two shrinkage occurrences (the first occurring during the sintering of the ceramic precursor and the second during the sintering of the scaffold) into one shrinkage occurance (occurring during the sintering of the scaffold). This could result in a larger instance shrinkage after sintering, thereby producing scaffolds of higher precision.
Based on the data shown in Figure 3, we have plotted a summary table, Table 1, which shows the shrinkage rate of the lines of the SCM scaffolds and UCM scaffolds at different nozzle diameters. In Table 1, we can observe that regardless of the nozzle diameter, the shrinkage of the UCM scaffold is greater than that of the SCM scaffold. This finding provides evidence that our idea is feasible. Additionally, we have also compared the shrinkage rates of the two scaffold pore sizes.
As shown in Figure 4, surface and cross-sectional views of scaffolds manufactured using filling densities of 0.4, 0.5, and 0.6 are presented. As the filling density increases, the pore size of the scaffolds becomes smaller. When the filling density was 0.4, the scaffold exhibited the largest pore size, with a pore size of about 500.5 ± 12.5 μm and a shrinkage rate of 66.7 ± 1.7% for the SCM scaffold and a pore size of about 397.3 ± 27.3 μm and a shrinkage rate of 53.0 ± 3.6% for the UCM scaffold. When the filling density is 0.5, the pore size was approximately 385.0 ± 40.7 μm with a shrinkage rate of 77.0 ± 8.1% for the SCM scaffold and there was a pore size of approximately 261.5 ± 21.3 μm with a shrinkage rate of 52.3 ± 4.3% for the UCM scaffold. When the filling density was 0.6, the pore size of the scaffold was the smallest, with a pore size of about 268.8 ± 20.9 μm and a shrinkage rate of 80.6 ± 6.3% for the SCM scaffold and a pore size of about 211.3 ± 26.3 μm and a shrinkage rate of 63.4 ± 7.9% for the UCM scaffold.
As we can see in the graph, the UCM scaffold has smaller pore sizes than the SCM scaffold, which is due to the higher shrinkage of the UCM scaffold. This is beneficial for the manufacturing process of porous scaffolds because it allows for more precise control of the pore sizes and promotes the osteogenic effect of the scaffold.
Drawing from the data in Figure 4, we have constructed a summary table, Table 2. This table displays the shrinkage rate of the pore sizes in both the SCM and UCM scaffolds at varying filling densities. In Table 2, we can observe that regardless of the filling density, the shrinkage of the UCM scaffold is greater than that of the SCM scaffold. This finding further validates our idea that improving the sintering strategy of scaffolds can enhance their precision.
However, due to the tendency of powder to agglomerate, using a 0.4mm nozzle to manufacture the scaffold can easily result in nozzle clogging. Meanwhile, the scaffold produced using a 0.6mm nozzle had thicker lines and less-distinct pores, which can also lead to pore blockage. Therefore, we ultimately chose to use a 0.5mm nozzle to manufacture the scaffold and proceeded to the next experiment.
As shown in Figure 3 and Figure 4, compared to SCM powder, extrusion-based 3D printing technology can produce finer lines and more precise pore sizes using UCM powder. This is beneficial to our manufacturing process for porous scaffolds as it allows for more precise control over pore size and facilitates the production of higher porosity. To achieve this, we propose a multi-layer extrusion printing approach.
Figure 5 illustrates different layers of the scaffolds. It can be seen that the scaffold with a single layer demonstrates the smallest cross-sectional pore size; whereas, the scaffold comprising three layers presents the largest pore size. It is noticeable that the cross-sectional pore size in the scaffold escalates in correlation with the increasing number of layers.
In all of the surface and cross-sectional views, under the same conditions, the UCM scaffold exhibited finer lines, smaller pore sizes, and greater shrinkage compared to the SCM scaffold, further demonstrating the great development potential of the UCM scaffold.
In the process of scaffold manufacturing, process parameters are important factors in determining the final scaffold performance, especially the mechanical strength of the sample. Therefore, we conducted a series of studies to explore the effects of process parameters, such as slurry-filling density and printing layers, on scaffold sample performance. For example, we studied the effect of different filling densities on scaffold performance, as shown in Figure 6A, where samples with filling densities of 0.4, 0.5, and 0.6 are arranged from left to right. We can see that as the filling density increases, the pore size of the scaffold gradually decreases. In addition, we also studied the effect of different layers on scaffold performance, as shown in Figure 6B, where samples with three, two, and one layers are arranged from left to right. We can see that as the number of layers increases, the pores on the sides of the scaffolds become more and more pronounced. In Figure 6, all UCM scaffolds under the same conditions have greater shrinkage compared to the SCM scaffolds. During scaffold manufacturing, UCM scaffold samples undergo significant shrinkage, which occurs during scaffold sintering and includes shrinkage in length, width, and height. In addition, modifying the powder material enables us to produce scaffolds using various materials, including hydroxyapatite and tricalcium phosphate.

3.2. Effect of Fill Densities on the Mechanical Property

Figure 7 illustrates the effect of different fill densities on the compressive strength of SCM and UCM scaffolds. Changes in fill density lead to alterations in the porosity of the bioactive scaffold, which can impact the scaffold’s integration with the surrounding bone tissue, bone tissue regeneration, and the stability and mechanical characteristics of the bioactive scaffold. If the fill density is too low, it can cause a reduction in scaffold stability and mechanical properties, thus affecting the repair outcome. Conversely, if the fill density is too high, it may result in a decrease in the scaffold pore size and hinder the scaffold’s osteogenic effect. Therefore, we selected groups with fill densities of 0.4, 0.5, and 0.6 for testing. In the graph, it can be observed that both SCM and UCM scaffolds possess considerable compressive strengths at different fill densities and that compressive strength increases with increasing fill density. When the fill density was set to 0.4, the compressive strength of the SCM scaffold was (~51.5 MPa) while that of the UCM scaffold was (~47.6 MPa); when the fill density was set to 0.5, the compressive strength of the SCM scaffold was (~93.1 MPa) while that of the UCM scaffold was (~70 MPa); and when the fill density was set to 0.6, the compressive strength of the SCM scaffold was (~120.6 MPa) while that of the UCM scaffold was (~108.4 MPa). In summary, the compressive strength of the scaffold was the lowest at a fill density of 0.4 and the highest at a fill density of 0.6, which is attributed to changes in porosity. At the same time, we noticed that under the same conditions (the same number of layers, the same filling density), the mechanical strength of the UCM scaffold was slightly lower than that of the SCM scaffold. As in our previous work, changing the sintering curve of the scaffold in this case will have an impact on the mechanical properties of the scaffold [33]. We hypothesize that this might be due to the elimination of the sintering step of the ceramic precursor in the production of the UCM scaffold.

3.3. Effect of the Numbers of Layers on the Mechanical Property

Figure 8 illustrates the effect of different numbers of layers on the compressive strength of SCM and UCM scaffolds at a fill density of 0.5. The size of the side holes of the scaffold changes with the increasing number of layers; the side holes of the bioactive scaffold can increase the bonding force between the scaffold and the surrounding bone tissue, accelerate bone tissue regeneration, and accelerate the scaffold degradation rate. Side holes that are too large or too small can have adverse effects on the bioactive scaffold. Therefore, we selected one, two, and three layers as the experimental groups. As shown in the figure, both SCM and UCM scaffolds exhibit significant compressive strengths at different numbers of layers; the compressive strength gradually decreases with increasing layers. When the number of layers is set to one, the compressive strength of the SCM scaffold is (~93.1 MPa) and that of the UCM scaffold is (~70 MPa). When the number of layers is set to two, the compressive strength of the SCM scaffold is (~72.3 MPa) and that of the UCM scaffold is (~60.6 MPa). When the number of layers is set to three, the compressive strength of the SCM scaffold is (~51.6 MPa) and that of the UCM scaffold is (~40.6 MPa). In summary, the compression strength of the scaffold is the highest when the number of layers is one and lowest when it is three; this is due to the indirect effect of changes in the side holes on the porosity of the scaffold.

3.4. Biodegradation Testing In Vitro

Figure 9A displays the in vitro biodegradation experiment of the scaffolds. Figure 9B–F show the changes in the pH, inorganic ion release concentration, weight, compressive strength, and elastic modulus of the scaffold during the three-week in vitro biodegradation test.
In Figure 9B, we can observe that with an increase in immersion time, the pH of the degradation solution shifted from neutral 7.4 to alkaline. The pH of the SCM scaffold was around 7.7, slightly lower than the pH of the UCM scaffold. This change is postulated to have occured due to the interaction of SiO34− ions released from the scaffold with H+ ions in the solution [23]. Previous studies suggest that a mildly alkaline microenvironment can enhance the activity and proliferation of osteoblasts [23]. However, if the inorganic ions released from bioactive glass cause the pH of the cell culture medium to become too high, cells will die. Consequently, a slight elevation in pH levels has a positive impact on the proliferation of osteoblasts.
Figure 9C illustrates the variation in inorganic ion release concentration with soaking time. Throughout the degradation process, Ca2+, Si4+, and Mg2+ were gradually liberated as the material broke down. By the conclusion of the soaking period, the amount of Ca release was at the highest, attaining 181.3 ± 4.3 ppm for the SCM scaffold, and 189.8 ± 4.7 ppm for the UCM scaffold. The Si released from the SCM scaffold was 55.2 ± 0.2 ppm while that from the UCM scaffold was 62.3 ± 0.3 ppm. The Mg released from the SCM scaffold was 20.1 ± 0.1 ppm and that from the UCM scaffold was 25.4 ± 0.1 ppm. These data demonstrate that the UCM scaffold can steadily release Ca2+, Si4+, and Mg2+, which can promote bone tissue growth. Meanwhile, we observed that the ion release of UCM scaffolds is higher than that of SCM scaffolds; although, it is generally comparable. This may be attributed to the higher porosity of UCM scaffolds compared to SCM scaffolds under the same conditions.
Figure 9D displays the weight variation during the soaking process. It was found that the mass of the scaffold degraded significantly in the first week. The mass of SCM scaffolds decreased by 3.63–4.59% during the initial week; that of the UCM scaffold decreased by 5.39–6.46%. Subsequently, the rate of weight loss slowed down. After three weeks, the mass of all SCM scaffolds decreased by 9.46–10.36%; that of the UCM scaffold decreased by 11.46–12.78%. We observed that compared to the SCM scaffolds, the UCM scaffolds exhibited faster mass decay. This further supports our previous hypothesis that this may be due to the UCM scaffolds having a higher degree of shrinkage under the same conditions compared to the SCM scaffolds. Greater shrinkage results in thinner lines and, therefore, UCM scaffolds of the same size have more lines. The above results lead to an increased contact area between the UCM scaffolds and the tris solution, thereby accelerating the degradation rate of the UCM scaffolds.
Figure 9E shows the variation of compressive strength in relation to soaking time. During the degradation process of the material, the compressive strength gradually decreased. In this process, the compressive strength of the SCM scaffold reduced by about 39.93%, from approximately 93.09 MPa to approximately 55.92 MPa, while that of the UCM scaffold decreased by about 32.70%, from approximately 70.04 MPa to approximately 47.14 MPa. After the entire 3-week soaking period, the UCM scaffold still had a considerable amount of compressive strength, indicating good mechanical properties. This aligns with our previous conclusion, that the mechanical decay rate of the scaffold is greater in the initial stages and that over time, the rate of mechanical decay gradually decreases [22]. On the other hand, this also corresponds with the trend of mass decay shown in Figure 9D.
Figure 9F shows the variation of elastic modulus in relation to soaking time. The trend of elastic modulus variation is consistent with the trend of compressive strength, showing an overall downward trend with time. After three weeks, the elastic modulus of the SCM scaffold decreased to approximately 424.61 Mpa; meanwhile, the UCM scaffold still had a significant elastic modulus, which decreased to approximately 404.60 MPa. We observed that under the same conditions, the mechanical performance of UCM scaffolds was lower than that of SCM scaffolds, both before and after degradation. We observed in the experiment that the UCM scaffolds and SCM scaffolds had similar elastic moduli, which was because our sintering strategy did not change the material composition of the UCM scaffolds. Meanwhile, we also noticed that the elastic modulus of the UCM scaffolds was slightly lower than that of the SCM scaffolds at all three time points, which was a result of the faster degradation rate of the UCM scaffolds.

3.5. Cell Viability In Vitro

Figure 10 shows the results of the in vitro cell viability assay. We used the CCK8 to determine the viability of BMSCs in samples incubated for 1, 4, and 7 days. In Figure 10, we can see that the optical density is on an upward trend, indicating that the cells can proliferate on both UCM and SCM scaffolds. This result is similar to that of our previous work [22]. The results shown in the figure indicate that the cell count of the UCM group was close to that of the SCM group on days one, four, and seven, with no significant difference, demonstrating that the UCM scaffold had a similar capability to promote BMSC viability as the SCM scaffold. We believe this is due to our sintering strategy, which did not alter the composition of the powder components. On the other hand, we observed that the BMSC viability in the UCM scaffold experimental group was slightly higher than that in the SCM scaffold group on days one, four, and seven. We attribute this to the UCM having finer lines and smaller pore sizes. This further validates our high-precision ceramic material 3D printing manufacturing strategy, indicating that it holds promising application prospects.
In bone-repair therapy applications, biologically active scaffolds need to have high precision. However, scaffolds manufactured using traditional extrusion-based 3D printing technology have lower precision, which limits their clinical applications. The high-precision extrusion-based 3D printing manufacturing technology developed in this study can manufacture CSi-Mg scaffolds with better line and pore accuracies, thus improving their potential for use in bone-repair therapy. In addition, CSi-Mg ceramics have excellent mechanical properties and can obtain an anti-tumor function by adsorbing hydroxyapatite, displaying controllable biodegradable characteristics and good in vivo mechanical properties [22].
We further studied UCM scaffolds with different filling densities and layers to control their line thickness and pore size, making them more suitable for clinical applications. These research findings suggest that high-precision CSi-Mg scaffolds show promising potential for practical applications and can play an important role in future bone-repair therapy.

4. Conclusions

In this study, we developed an extrusion-based 3D printing strategy to improve the precision of bioceramic scaffolds. Compared to the SCM scaffolds manufactured through traditional sintering processes, the high-precision UCM scaffolds produced by this method exhibit superior line and pore precision. Additionally, in vitro degradation tests demonstrated that the high-precision UCM scaffolds have an excellent compressive strength and modulus of elasticity. After soaking in simulated body fluid for three weeks, the UCM scaffold had a high elastic modulus (over 404 MPa) and significant compressive strength (over 47 MPa). The cell culture experiments also confirmed that UCM scaffolds are more conducive to supporting and guiding cell growth to form new or replacement tissues, compared to SCM scaffolds. We believe that this method provides a promising strategy for manufacturing high-precision bioceramic scaffolds using extrusion-based 3D printing technology. This will broaden the prospects for manufacturing methods in bone tissue engineering; this high-precision CSi-Mg scaffold can be used to treat bone defects, fractures, and other bone injuries, providing a better treatment option for bone-injury repair.

Author Contributions

Conceptualization, H.S.; methodology, J.S.; validation, H.W.; formal analysis, Z.H.; investigation, H.S. and J.S.; data curation, J.S.; writing—original draft preparation, J.S.; writing—review and editing, H.S. and J.S.; supervision, H.S.; project administration, W.Y.; funding acquisition, H.S. All authors have read and agreed to the published version of the manuscript.

Funding

This work was supported by the National Natural Science Foundation of China (51805475), Zhejiang Provincial Natural Science Foundation of China (LY22E050011), Zhejiang Provincial Postdoctoral Research Funds (ZJ2022009), Postdoctoral Science Foundation of China (2022M722736), Fundamental Research Funds for the Provincial Universities of Zhejiang (GK229909299001-304), Open Foundation of the State Key Laboratory of Fluid Power and Mechatronic Systems (GZKF-202102).

Data Availability Statement

Not applicable.

Conflicts of Interest

The authors declare no conflict of interest.

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Figure 1. A concise overview of the research project. (A) A diagrammatic representation of the fabrication procedure for the SCM scaffold. (B) A diagrammatic representation of the fabrication procedure for the UCM scaffold.
Figure 1. A concise overview of the research project. (A) A diagrammatic representation of the fabrication procedure for the SCM scaffold. (B) A diagrammatic representation of the fabrication procedure for the UCM scaffold.
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Figure 2. Characterization of the SCM and UCM powders after sintering at 1140 °C. (A) XRD patterns of the SCM and UCM powders. (B) Scanning electron microscopy image of the SCM powders. (C) Scanning electron microscopy image of the UCM powders. (D) The particle size distribution of the SCM powder. (E) The particle size distribution of the UCM powder.
Figure 2. Characterization of the SCM and UCM powders after sintering at 1140 °C. (A) XRD patterns of the SCM and UCM powders. (B) Scanning electron microscopy image of the SCM powders. (C) Scanning electron microscopy image of the UCM powders. (D) The particle size distribution of the SCM powder. (E) The particle size distribution of the UCM powder.
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Figure 3. SEM micrographs of the cross-section and surface of the SCM scaffold and UCM scaffold fabricated using different nozzle diameters (0.4 mm, 0.5 mm, 0.6 mm).
Figure 3. SEM micrographs of the cross-section and surface of the SCM scaffold and UCM scaffold fabricated using different nozzle diameters (0.4 mm, 0.5 mm, 0.6 mm).
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Figure 4. SEM micrographs of the cross-sections and surfaces of the SCM scaffold and UCM scaffold fabricated using different filling densities (0.4, 0.5, and 0.6).
Figure 4. SEM micrographs of the cross-sections and surfaces of the SCM scaffold and UCM scaffold fabricated using different filling densities (0.4, 0.5, and 0.6).
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Figure 5. SEM micrographs of the cross-section and surfaces of the SCM scaffold and UCM scaffold fabricated using different numbers of layers (three, two, and one).
Figure 5. SEM micrographs of the cross-section and surfaces of the SCM scaffold and UCM scaffold fabricated using different numbers of layers (three, two, and one).
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Figure 6. Images of SCM scaffold and UCM scaffold (A) with different filling densities and (B) with different numbers of layers.
Figure 6. Images of SCM scaffold and UCM scaffold (A) with different filling densities and (B) with different numbers of layers.
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Figure 7. The compressive strengths of scaffolds with different filling densities (* p < 0.05).
Figure 7. The compressive strengths of scaffolds with different filling densities (* p < 0.05).
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Figure 8. The compressive strengths of scaffolds with different numbers of layers (* p < 0.05).
Figure 8. The compressive strengths of scaffolds with different numbers of layers (* p < 0.05).
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Figure 9. (A) Degradation in vitro. (B) Alterations in the pH levels of Tris. (C) The release of ions throughout the degradation period. (D) Changes in weight (%). (E) The compressive strength and elastic modulus (F) of the scaffolds plotted over time.
Figure 9. (A) Degradation in vitro. (B) Alterations in the pH levels of Tris. (C) The release of ions throughout the degradation period. (D) Changes in weight (%). (E) The compressive strength and elastic modulus (F) of the scaffolds plotted over time.
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Figure 10. Cell viability at 1, 4, and 7 days.
Figure 10. Cell viability at 1, 4, and 7 days.
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Table
Table 1. Table of the shrinkage rate of scaffold lines under different nozzle diameters.
Table 1. Table of the shrinkage rate of scaffold lines under different nozzle diameters.
SampleNozzle Diameter
0.4 mm0.5 mm0.6 mm
Line Thickness (μm)Shrinkage Rate (%)Line Thickness (μm)Shrinkage Rate (%)Line Thickness (μm)Shrinkage Rate (%)
SCM scaffold279.1 ± 18.969.8 ± 4.7300.8 ± 17.560.2 ± 3.5315.0 ± 17.952.5 ± 3.0
UCM scaffold190.3 ± 11.747.6 ± 2.9218.4 ± 15.343.7 ± 3.1233.5 ± 16.738.9 ± 2.8
Table
Table 2. Table of the shrinkage rates of scaffold pore sizes under different filling densities.
Table 2. Table of the shrinkage rates of scaffold pore sizes under different filling densities.
SampleFilling Density
0.40.50.6
Pore Size (μm) Shrinkage Rate (%)Pore Size (μm) Shrinkage Rate (%)Pore Size (μm) Shrinkage Rate (%)
SCM scaffold500.5 ± 12.566.7 ± 1.7385.0 ± 40.777.0 ± 8.1268.8 ± 20.980.6 ± 6.3
UCM scaffold397.3 ± 27.353.0 ± 3.6261.5 ± 21.352.3 ± 4.3211.3 ± 26.363.4 ± 7.9
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Shao, H.; Shi, J.; Huang, Z.; Yang, W.; Wang, H. 3D-Printed Bioceramic Scaffolds with High Strength and High Precision. Crystals 2023, 13, 1061. https://doi.org/10.3390/cryst13071061

AMA Style

Shao H, Shi J, Huang Z, Yang W, Wang H. 3D-Printed Bioceramic Scaffolds with High Strength and High Precision. Crystals. 2023; 13(7):1061. https://doi.org/10.3390/cryst13071061

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Shao, Huifeng, Jinyuan Shi, Zhiqiang Huang, Weibo Yang, and Honghua Wang. 2023. "3D-Printed Bioceramic Scaffolds with High Strength and High Precision" Crystals 13, no. 7: 1061. https://doi.org/10.3390/cryst13071061

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