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Enhancement of mechanical strength of TCP-alginate based bioprinted constructs
journal of the mechanical behavior of biomedical materials 103 (2020) 103533
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Enhancement of mechanical strength of TCP-alginate based bioprinted constructs Jie-Liang Song 1, Xin-Ye Fu 1, Ali Raza, Nai-An Shen, Ya-Qi Xue, Hua-Jie Wang *, Jin-Ye Wang ** School of Biomedical Engineering, Shanghai Jiao Tong University, 800 Dongchuan Road, Shanghai, 200240, China
A R T I C L E I N F O
A B S T R A C T
Keywords: Bioprinting PCL supporter Unit-assembly model Alginate-TCP bioink Bone defect
To overcome the mechanical drawback of bioink, we proposed a supporter model to enhance the mechanical strength of bioprinted 3D constructs, in which a unit-assembly idea was involved. Based on Computed To mography images of critical-sized rabbit bone defect, the 3D re-construction was accomplished by a sequenced process using Mimics 17.0, BioCAM and BioCAD software. 3D constructs were bioprinted using polycaprolactone (PCL) ink for the outer supporter under extrusion mode, and cell-laden tricalcium phosphate (TCP)/alginate bioink for the inner filler under air pressure dispensing mode. The relationship of viscosity of bioinks, 3D bio printing pressure, TCP/alginate ratio and cell survival were investigated by the shear viscosities analysis, live/ dead cell test and cell-counting kit 8 measurement. The viscosity of bioinks at 1.0 s 1-shear rate could be adjusted within the range of 1.75 � 0.29 Pa⋅s to 155.65 � 10.86 Pa⋅s by changing alginate concentration, cor responding to 10 kPa–130 kPa of printing pressure. This design with PCL supporter could significantly enhance the compressive strength and compressive modulus of standardized 3D mechanical testing specimens up to 2.15 � 0.14 MPa to 2.58 � 0.09 MPa, and 42.83 � 4.75 MPa to 53.12 � 1.19 MPa, respectively. Cells could maintain the high viability (over 80%) under the given printing pressure but cell viability declined with the increase of TCP content. Cell survival after experiencing 7 days of cell culture could be achieved when the ratio of TCP/alginate was 1 : 4. All data supported the feasibility of the supporter and unit-assembly model to enhance mechanical properties of bioprinted 3D constructs.
1. Introduction
2016; Blaeser et al., 2017; Duarte Campos et al., 2016; Lee et al., 2014). Bioink is one of the key restrictions on the bioprinting technique. The requirement for the balance between rheological property and shape retention prevents the application of a number of native biomaterials as bioinks, especially for hard tissue implant fabrication, although they satisfy the requirements of biocompatibility and biodegradability (Compaan et al., 2016; Grande, 2017; Jang et al., 2017; Jia et al., 2014; Law et al., 2018; Pati et al., 2014; Rhee et al., 2016; Zehnder et al., 2015; Kim et al., 2018). In another word, bioprinting is a layer-by-layer printing technique and the lower layer needs to support the upper layer (Aljohani et al., 2018). Otherwise the structure will collapse. Especially, in view of the complex mechanical requirement of tissues and organs, the bioprinted 3D constructs with suitable mechanical properties are still a challenge (Carter and Spengler, 1978; Vanderhooft et al., 2009). The 3D structure design is a possible breakthrough point for
The bioprinting technique benefits the precise control over multiple compositions, spatial distributions, and architectural accuracy/ complexity, by layer-by-layer printing of cells-laden biomaterial (bio inks) (Blaeser et al., 2017; Moroni et al., 2018; Murphy and Atala, 2014). From this point of view, it is likely to achieve the highly biomimetic objective at structural features, mechanical properties and even function of native tissues (Cui et al., 2018). Although 3D printing has a predicted $ 23.0 billion market in 2022, the bioprinting technique is now in the ascendant, characterizing with more “comfortable” printing conditions for cells’ survival, such as mild printing temperature and pressure (Chang et al., 2008; Derakhshanfar et al., 2018; Zhao et al., 2015). Various engineered tissues or organs have been obtained by the bio printing technique, such as cartilage, skin tissue, bone tissue, muscle, neural, osteochondral and pancreatic tissue and renal (Abbadessa et al.,
* Corresponding author. ** Corresponding author. E-mail addresses: [email protected] (H.-J. Wang), [email protected] (J.-Y. Wang). 1 These authors contributed equally to this work. https://doi.org/10.1016/j.jmbbm.2019.103533 Received 19 March 2019; Received in revised form 7 November 2019; Accepted 12 November 2019 Available online 13 November 2019 1751-6161/© 2019 Elsevier Ltd. All rights reserved.
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Journal of the Mechanical Behavior of Biomedical Materials 103 (2020) 103533
mechanical improvement of bioprinted 3D constructs (Daly et al., 2016; Hu et al., 2018). The hybrid scaffold consisting of a synthetic polymer as supporter with cell-laden hydrogels is a typical and successful design (Lee et al., 2013). For example, Schuurman et al. established a state-of-the-art printing technology to obtain a polycaprolactone (PCL)/alginate hybrid tissue construct with over 6 MPa of Young’s modulus, which matched the range with cartilage 4.1 MPa and trachea 3.33 MPa (Schuurman et al., 2011). In this design, alginate/cells bioinks were filled into the grooves formed by the PCL strands and PCL played a key important role in supplying a mechanical support to the 3D con structs, while alginate worked as the matrix to localize cells. Instead of alginate, Pati et al. used decellularized adipose tissue matrix/mesen chymal stem cells bioink to print a dome-shaped structure with PCL, which had 122.56 � 20.23 kPa of compressive modulus (Pati et al., 2015). In addition, Campos et al. recently reported a model composed of hollow cylinders with three supporting pins and their negatives (Campos et al., 2019). The mixture of agarose and alginate was chosen as the supporting material, and the mixture of agarose and type I collagen was chosen as bioink. The highest compressive modulus was over 600 kPa. Although efforts in 3D bioprinting field have been made to expand the applicability of cell-laden hydrogels from soft tissue regeneration to the more demanding field of hard tissue regeneration, there are few reports focusing on bone tissue engineering due to their relatively weak me chanical properties. In this work, we proposed a supporter model to enhance the me chanical strength of bioprinted 3D constructs, in which a unit-assembly idea was involved (Scheme 1). Here, the unit was composed of the PCL supporter and cell-laden tricalcium phosphate (TCP)-alginate filler. The former formed the up-down pipelines, columnar linkers, sides and bot tom of bioprinted 3D construct. The latter acted as bioinks to fill in the
PCL frame. After that, the units were assembled together to form the whole 3D constructs by the columnar linkers. The unit-assembly idea had the advantages of less damage to cells due to shorter time for printing of units and easy correction on eventual failures in the fabri cation or in vitro culture. 2. Materials and methods 2.1. Materials PCL (Mw ¼ 80,000 Da) was purchased from Sigma Aldrich (Natick, MA, USA) and used for holt-melt extrusion-based printing material without further purification. Alginate (medium viscosity, 200–500 mPa⋅s) was purchased from Adamas-beta Co. Ltd. (Shanghai, China) and TCP was purchased from Shanghai No.4 Reagent & H.V. Chemical Co. Ltd. (Shanghai, China). Fluorescein diacetate (FDA) and propidium iodide (PI) were purchased from Yeasen Biotech Co. Ltd. (Shanghai, China). Cell-counting kit 8 (CCK8) was purchased from Beyotime Biotechnology (Shanghai, China). 2.2. Cells The human hepatocarcinoma (BEL-7402) cells were obtained from Shanghai Institute of Biochemistry and Cell Biology (Shanghai, China). Cells were routinely cultured in RPMI1640 medium (containing 10% of new borne bovine serum, 100 units/mL penicillin and 100 μg/mL streptomycin). After 80% of confluency, cells were released from the bottom using 0.25% trypsin solution and collected by centrifugation at 1000 rpm for 5 min. BEL-7402 cells were dispersed into fresh medium to get 1 � 106 cells/mL density for bioinks preparation.
Scheme 1. Illustration of the design of 3D constructs and 3D bioprinting process. 2
Journal of the Mechanical Behavior of Biomedical Materials 103 (2020) 103533
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2.3. Animals New Zealand rabbits (3.0–4.0 kg) were purchased from Songlian Experimental Animal Farm (Shanghai, China). All animal procedures were approved by the Institutional Animal Care Committee. All guide lines met ethical standards required by law and also complied with the guidelines for the use of experimental animals in China. The 20-mm length of critical-size radial bone defect model in rabbit was made under anaesthesia via intramuscular administration with 3% pentobarbital sodium (30 mg/kg) according to Tu et al. (2009). The actual digital information of radial bone defect was obtained by micro-CT (SkyScan1176, Bruker, Germany) with 90 kV of voltage, 278 μA of current, 17.93 μm of isotropic voxel size and 1336 � 2000 pixel of field of view. All images were archived as Tiff files.
was magnified three times in order to test the unit-assembly idea and possibility to scale up for use in human-scale bone defect. As shown in Scheme 1, BioCAD software was used to process the unit-assembly model. Here, the re-constructed 3D structure contained 15 units, which had 4 mm height. A bottom, sides, up-down channels and columnar linkers of the units formed the supporter structure, which was printed using PCL ink. The thickness of the supporter layer on bottom and sides was 2 mm and 0.33 mm, respectively. The inner diameter and height of channels was 2.0 mm and 4.0 mm, respectively. The height of columnar linkers was 4.7 mm, including an upper ridge (2.0 mm in diameter and 1.2 mm in height) and a round groove (2.0 mm in diameter and 1.5 mm in height) at the bottom of the linker. In addition, the inner filler structure was bioprinted using cell-laden TCP/alginate bioinks. The units were linked with each other through columnar linkers.
2.4. Preparation and rheological analysis of bioinks
2.6. 3D bioprinting
Alginate and TCP powder (size < 100 μm) were sterilized by ultra violet (UV) irradiation for 30 min and mixed with the abovementioned cell suspension. The final concentration of alginate was 2%, 4% and 6% (w/v), respectively. The TCP/alginate ratio was listed in Table 1. The rheological properties of bioinks were analysed using a rota tional rheometer (MCR302, Anton Paar, Graz, Austria) (n ¼ 3). The steady-state shear viscosities of samples were conducted over a range of shear rate 1–1000 s 1 and all measurements were conducted at room temperature.
3D bioprinting process was performed using 3D bioprinter (3D Dis covery, RegenHU, Switzerland) under aseptic conditions. Here, the revolution of 3D bioprinter was 50 μm. Firstly, the melted PCL was extruded at 90 � C via a HM-300H print-head with 0.33 mm diameter and 20 r/s extruder speed. The layer height was 0.25 mm. The pressure and moving speed of nozzle were 0.3 MPa and 5 mm/s, respectively. Sec ondly, CF300H print-head without the valve was used for bioinks printing under continuous air pressure dispensing mode and the pres sure was adjusted in the range of 10–130 kPa according to the rheo logical properties of bioinks. Here, the valve of CF300H print-head was removed in order to match the cell-alginate/TCP bioinks and 19 G nozzle (0.9 mm-size of diameter) was used for CF300H print-head. The moving speed and temperature of nozzle were 5 mm/s and 37 � C, respectively. Thirdly, the printed units were immersed into 0.1 M CaCl2 solution for 20 min, and then rinsed with PBS three times. The units were further cultured in a 12-well plate in a 5% CO2 incubator at 37 � C. Finally, the units were assembled together one by one to form the whole defected bone as shown in the Fig. 1H.
2.5. Computer-aided 3D reconstruction of bone defect Mimics 17.0 software (Materialise, Belgium) was used to process micro-CT images for re-construction of bone defect, including image reading, filtering, threshold segmentation and 3D model reconstruction. The image processing was in detailed described in re sults and discussion section. The files were saved as STL format. BioCAM software (RegenHU, Switzerland) was used to read 3D reconstructed STL format file and the digital information of bone defect
2.7. Compressive properties of 3D constructs
Table 1 Viscosities at 1.0 s 1-shear rate and 3D bioprinting pressures for various bioink formulations (n ¼ 3). Groups
Alginate (% w/v)
TCP: Alginate (w: w)
Viscosity at 1.0 s 1shear rate (Pa⋅s)
Printing pressure (kPa)
TCP: 2% ALG (0 : 4) TCP: 2% ALG (1 : 4) TCP: 2% ALG (8 : 4) TCP: 4% ALG (0 : 4) TCP: 4% ALG (1 : 4) TCP: 4% ALG (8 : 4) TCP: 6% ALG (0 : 4) TCP: 6% ALG (1 : 4) TCP: 6% ALG (8 : 4)
2
0:4
1.75 � 0.29
10
2
1:4
2.44 � 0.45
10
2
8:4
2.19 � 0.17
10
4
0
39.19 � 4.89
45
4
1:4
33.28 � 4.26
45
4
8:4
33.80 � 4.05
50
6
0:4
123.04 � 37.37
120
6
1:4
115.82 � 3.39
120
6
8:4
155.65 � 10.86
130
The cylinder samples with the standard size (d ¼ 10 mm, h ¼ 10 mm) matching the requirement of the ISO 13314 : 2011 had been prepared using the abovementioned method. Briefly, the designed cylinders were fabricated by printing the supporter PCL perimeter and the different alginate/TCP inks within. The mechanical tests were carried out on a universal testing machine (Z020, ZwickRoell, Germany) at a rate of 1.0 mm/min (n ¼ 3). The resulting properties were calculated according to the following formula:
σ ¼ Eε where σ is the stress, ε is the strain, E is the compressive modulus which was calculated using slope of the linear region of the stress strain curve. 2.8. Cell survival and viability After experiencing a 4 h of cell culture, the absolute viability of the bioprinted cells in the standardized 3D units (14 mm � 14 mm � 4 mm) was tested using the Live/Dead assay kit according to the protocol (n ¼ 3) (Yeasen Biotech Co. Ltd., Shanghai, China). Briefly, 1 ml of FDA/ PI staining solution was added into each well of 12-well plate and incubated for 5 min at room temperature. After rinsing three times with PBS (pH 7.2–7.4, 0.1 M), the cells were counted under the fluorescent microscope (IX-70, Olympus, Japan). The live cells were green (FDA) and the dead cells were red (PI). The absolute cells’ viability0days was expressed as the percentage of the live cells number/total cells number. The relative viability of the bioprinted cells in the standardized 3D units after cell culture for 4 h and 7 days was measured using CCK-8 3
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Journal of the Mechanical Behavior of Biomedical Materials 103 (2020) 103533
Fig. 1. 3D re-construction of bone defect of rabbit by a sequenced process using Mimics, BioCAM and BioCAD software and the bioprinted 3D constructs. (A) Bone defect model in rabbit; (B) Micro-CT images of bone defect (1) and typical micro-CT cross section images (2); (C) 3D re-construction of bone defect by Mimics 17.0 software through a four step process, including acquisition, filtering, segmentation and charac terization process. Front view (1), Top view (2), Left view (3), Isometric projection (4); (5–8) threshold segmentation-dynamic regional growth, twenty-six connected body; the magnetic lasso method to re-construct defected section (5), top view (6), front view (7), left view (8); (D) Reading of the 3D digital information by BioCAM software. (E) 3D processing by BioCAD software. (F1) The bioprinted 3D constructs. (F2) Unit of 3D constructs.
detection kit. Briefly, the fresh medium containing 10% of CCK-8 solu tion was added to each well of 12-well plate and the cells were incubated at 37 � C for another 4 h. The OD values at 450 nm were recorded using a microplate reader (Multiskan FC, Thermo, USA). The relative and ab solute cells’ viability was calculated using the following formulas: Relative cells’ viability ¼ OD7days/OD0days � 100%
3. Results and discussion 3.1. Interface establishment of 3D bioprinter and micro-CT image In this work, the sequenced processing steps for re-construction of bone defect model were performed using Mimics 17.0, BioCAM and BioCAD software (Fig. 1) (McCormick et al., 2014; Sun et al., 2004). Firstly, the four-step process for the transformation from the micro-CT image (Tiff format) to 3D model file (STL format) was accomplished using Mimics 17.0, including image acquisition, filtering, segmentation and characterization process (Fig. 1C). Here, the median filter produces the best noise reduction effect without damaging the edge of bone tissue compared to other modes. After that, image segmentation is considered to be a crucial step to identify the region of interest, which can be per formed by thresholding, edge, region or cluster-based methods. Dy namic threshold segmentation and gray value segmentation were used. Both methods could easily and accurately separate bone tissue from other parts of the image. However, it was much difficult to generate the mask of bone defect because there were little differences between the gray value of the bone defect area and the surroundings, and the boundary of bone defect area was not clear. Therefore, the live wire algorithm is used to draw the outline of the target area manually, and then edited the generated masks in each layer as shown in Fig. 1C5 - C8. The digital 3D model of bone defect is re-constructed and identified by
(1)
Absolute cells’ viability7days ¼ Relative cells’ viability � Absolute cells’ via bility0days � 100% (2)
2.9. Statistical analysis All data were reported as the mean � standard deviation. The data were analysed by one-way factorial analysis of variances (one-way ANVOA) and multiple comparisons. Significant effects were defined using Fisher’s method as a post-hoc test, with p < 0.05 considered as being statistically significant.
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Table 1 lists the corresponding viscosity at 1.0 s 1-shear rate and 3D bioprinting pressure. Generally speaking, the viscosities of as-prepared samples highly rely on the concentration of alginate. The curves for bioinks containing 2% alginate appear to be relatively flat, suggesting minimal shear-thinning property. When alginate concentration is over 4%, the curves become steep, suggesting that a higher pressure will be needed for printing. In the work reported by Wu et al. the viscosity of pure alginate solution was also verified to be positively relative to the concentration of alginate and the viscosity at 1.0 s 1-shear rate varied in the range of 0–10 Pa⋅s when the concentration of alginate increased from 2% to 6% (Wu et al., 2018). Here, the viscosities at 1.0 s 1-shear rate of bioinks are in the range of 1.75 � 0.29–155.65 � 10.86 Pa⋅s and increase with the increase of alginate concentration. The addition of TCP doesn’t significantly affect shear-thinning property of bioinks. Freeman and Kelly demonstrated that molecular weight and ionic crosslinker were two important factors affecting the viscosity and printability of alginate bioinks (Freeman and Kelly, 2017). Our alginate belonged to medium viscosity product, while the alginate used by Wu et al. was pharmaceutical grade and belonged to low viscosity product (Wu et al., 2018). Therefore, it was the main reason to explain the difference of rheological properties of bioinks between two groups. The original CF300H printhead with a valve (3D Discovery, RegenHU) is very suitable to print inks with lower viscosity, especially cell solution. As for our bioinks, the relatively high viscosity and exis tence of TCP powder very easily led to block of the printhead. Especially for the printing of big-size samples, it was very difficult to continuously print a complete 3D structure. To solve this problem, we removed the valve and used a 10 cc cartridge, in combination with a suitable nozzle (19 G), to print our bioinks. Meanwhile, because the relatively high viscosity, the bioinks couldn’t flow freely without air pressure. It was noted that we could ignore the free-flow problem of bioinks when we controlled the moving speed of nozzle at more than 5 mm/s. The higher viscosity was associated with the higher printing pressure. As shown in Table 1, the printing pressure for 3D bioprinting of bioinks ranges from 10 kPa to 130 kPa. The bioinks containing 2% alginate can be easily pushed out, but can’t keep shape without the supporter due to high fluidity (Fig. 3A1-A3). PCL supporter and the post-crosslinking process using CaCl2 enable these bioinks constructing the 3D units without collapse, but all designed PCL channels become closed (Fig. 3a1-a3). On the contrary, the bioinks containing 6% alginate have the lower fluidity and can keep shape well (Fig. 3C1-C3). However, the printed bioinks shrink after crosslinking by immersion in 0.1 M CaCl2 solution for 20 min, and could not completely attach to the PCL supporter (Fig. 3c1c3). While 4% of alginate concentration shows a perfect 3D structure (Fig. 3B and b). The compressive strength and compressive modulus of bioprinted 3D samples vary from 2.15 � 0.14 MPa to 2.58 � 0.09 MPa, and 42.83 � 4.75 MPa to 53.12 � 1.19 MPa, respectively. All values satisfy the mechanical demand of cancellous bone (strength: 1.5–38 MPa and modulus: 10–1570 MPa) as shown in Table 2 (Yusop, 2012). 3D struc ture design play the key role to improve mechanical properties of printed 3D constructs. For example, Schuurman et al. reported the typical groove (synthesis polymer)/filler (cell-laden bioink) and layer-by-layer structure that had over 6 MPa of Young’s modulus (Schuurman et al., 2011). Xu et al. accomplished the mechanical improvement of 3D printed constructs with 1.76 MPa of Young’s modulus, which was four times higher than that of alginate constructs (0.41 MPa) (Xu et al., 2013). Their structure was characterized with an alternative layers structure of PCL and cell-laden hydrogel. In this study, Young’s modulus of obtained 3D mechanical testing specimens was evidently improved by comparison with pure alginate (less than 0.7 MPa) (Ren et al., 2018; Roushangar et al., 2018), and there were no significant differences no matter if there was TCP, indicating that PCL was responsible for the mechanical strength of 3D constructs. Actually it is still a challenge for mechanical test of irregular samples in the related fields. In general, the precise testing depends on testing method, and the
BioCAM (Fig. 1D). Upon early diagnosis, the exact anatomical model of target can be acquired through CT or magnetic resonance imaging (MRI) imaging modalities without surgical intervention and with minimum radiation exposure (Fu et al., 2017). Based on CT or MRI images, the design of tissue constructs before bioprinting will play an extremely crucial role in determining the properties of bioprinted constructs, which must conform to injury-specific geometrical dimensions (Fig. 1B and C) (Nam et al., 2015; Teodori et al., 2017). Besides Mimics, there are other design software for 3D modelling, such as TSIM, Solidworks, 3D Slicer, MATLAB and OsiriX. 3.2. Design of 3D constructs Scheme 1 shows the design of 3D constructs by BioCAD software (Fig. 1D) and printing process. This study not only supplies a supporter model to enhance the mechanical strength of 3D-bioprinted structure, but also a unit-assembly idea is involved. Most hydrogels fail to establish 3D structures with big size or good mechanical properties because of their inherently incongruous property between rheology and shape retention. The supporter design is an alternative to improve this status, such as the work of Daly et al. Lee et al., Schuurman et al. and pati et al. (Daly et al., 2016; Lee et al., 2013; Pati et al., 2015; Schuurman et al., 2011). The hybrid scaffold consisting of a synthetic polymer as sup porter with cell-laden hydrogels was the basis of their 3D designs. The Young’s modulus even could be enhanced up to megapascales (MPa) level (Schuurman et al., 2011). In this study, we combined the supporter structure with unit-assembly structure. The designed 3D structure is divided into 15 sections that were named as units with the columnar linker and up-down channels. The upper ridge of columnar linker could match with the bottom groove of adjacent columnar liner, enabling the proper fixing of neighbouring units. The up-down channels allowed solution to pass through. There were two advantages for this design. Firstly, the units could be bioprinted in a relative shorter time by comparison with the whole 3D structure. This would support a higher survival rate to cells due to the adverse effects from environmental factors during the printing process. Secondly, one unit could be bioprinted into several identical backups and easily replaced if some of them met with damage or infection before use. 3.3. Preparation of bioinks and printing of 3D constructs We studied the rheological properties of as-prepared bioinks. Fig. 2 shows the viscosity vs. shear rate curves for all bioink formulations, and
Fig. 2. The viscosity vs. shear rate curves for all bioink formulations (n ¼ 3). 5
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3.4. Cells’ viability The standardized units were used to compare cell viabilities after 3D printing and 7 days cell culture (Fig. 4 and Fig. 5). The property of hydrogels themselves, and bioprinting parameters, such as nozzle tem perature, printing time, nozzle diameter, and dispensing pressure would affect the viability of cells in 3D bioprinted constructs (Chang et al., 2008; Zhao et al., 2015). In the view of biocompatibility of alginate, many literatures have demonstrated this point (Morimoto et al., 2017). The presence of alginate will help the precision location of cells during 3D bioprinting process. In addition, our data also supported that algi nate was able to help shield cells from shear stress during the extrusion process (Moroni et al., 2018). BEL-7402 cells in alginate group (TCP: alginate ¼ 0 : 4) had over 85% of cell viability after printing and it got to over 95% in TCP: 2% alginate ¼ 0 : 4 group. Cell death is a key challenge for 3D bioprinting technique. The printing pressure is the key factor that firstly should be considered. In the preliminary work, we tried to print cell suspension without any matrix materials under different printing pressures. The results revealed that the cells viability decreased greatly when the printing pressure was over 120 Pa. Nair et al. used 1.5% w/v alginate solution to study the effect of printing pressure on cells’ survival and gave a mathematical model (Nair et al., 2009). They found that live cells, injury cells and dead cells got to over about 85%, 3% and 7.5%, respectively, using 34.5 kPa pressure and 400 μm diameter of nozzle. However, when the printing pressure was set as 275.8 kPa, live cells, injury cells and dead cells only got to about 60%, 5% and 35%, respectively. As for TCP: alginate bio inks, the suitable rheological properties enabled them to be printed under less than 120 Pa and our data suggested that the effect of printing pressure within the range of 10–120 Pa on cell viability could be ignored. In the preliminary study, we also tried to improve the printing ac curacy and found the bioink was the key factor. It greatly affected the printing pressure (regardless of print modes) and the accuracy (including the accuracy in X, Y and Z). The former decides cells’ viability that has been proved in several groups, including ours. The latter in volves the printability and shape retention. At the same conditions, the increase of accuracy means the decrease of the used nozzle size and the subsequently increase of printing pressure. In the present study, we could get over 90% of cells’ survival rate at the given conditions (with less than 120 Pa of printing pressure). When we tried to change a smaller size of nozzle, the improvement of accuracy was at the expense of the cells’ survival rate under air pressure dispensing mode, even an extru sion mode was applied. Additionally, it was not neglectable that the cell viability decreased with the increase of TCP (Fig. 4). TCP, as another component of bioinks, is one of the most widely used biomaterial in bone tissue engineering. TCP has demonstrated osteogenic properties, phase stability and strong bond formation with the host bone tissue in different studies (Del Rosario et al., 2015; Fahimipour et al., 2017; Mojahedian et al., 2016; Park et al., 2016; Yang et al., 2015). With the increase of TCP concen tration, the decrease in cells’ viability was clearly observed (Fig. 4). In 2% alginate-containing groups, cell viability declined from 97.28% � 3.00% (TCP: 2% ALG (0: 4)) to 81.33% � 13.18% (TCP: 2% ALG (8: 4)), while cell viability declined from 88.49% � 12.47% (TCP: 6% ALG (0 : 4)) to 1.32% � 2.31% (TCP: 6% ALG (8 : 4)) in 6% alginate-containing groups. Jiang et al. reported that the high shear stress was the main reason to induce cell membrane damage and decrease cell viability during 3D bioprinting process (Jiang et al., 2017). The existence of hard TCP powder might enhance the shear stress of bioinks. The factors from inks and 3D structure on cell viability have been demonstrated by several groups (Berg et al., 2018; Habib et al., 2018; Wu et al., 2016; Wu et al., 2018). The work of Schuurman et al. showed that the cells viability in 3D constructs made of PCL/C20A4 cells-laden alginate bioinks was over 80% after bioprinting process (Schuurman
Fig. 3. Optical images of different 3D units with (Lowercase letters) or without (Capital letters) PCL supporter. Table 2 Compressive strength and modulus of bioprinted 3D constructs (n ¼ 3). TCP: 2% ALG (0 : 4) TCP: 2% ALG (1 : 4) TCP: 2% ALG (8 : 4) TCP: 4% ALG (0 : 4) TCP: 4% ALG (1 : 4) TCP: 4% ALG (8 : 4) TCP: 6% ALG (0 : 4) TCP: 6% ALG (1 : 4) TCP: 6% ALG (8 : 4)
Compressive strength (MPa)
Compressive modulus (MPa)
2.42 � 0.08 2.52 � 0.11 2.15 � 0.14 2.37 � 0.09 2.40 � 0.05 2.32 � 0.09 2.58 � 0.09 2.48 � 0.18 2.45 � 0.02
48.50 � 7.55 52.64 � 6.64 42.90 � 6.19 44.27 � 2.18 46.73 � 8.36 42.83 � 4.57 53.12 � 1.19 47.56 � 8.31 49.92 � 2.19
production, size and shape of the test specimen. Therefore, the stan dardized regular shape for mechanical tests of specimen is required. In the present study, the method supplied by ISO 13314:2011 had been utilized for preparation and compressive test of our units. The cylinder samples had the standard size (d ¼ 10 mm, h ¼ 10 mm) and were pre pared using bioprinting method. We can expect that the basic mechan ical properties obtained from the standard smaller units experiments will help to determine the strength and other coefficients of the big irregular structure while application, by simulation such as finite element methods. This is the normal and similar way often used in En gineering Mechanics. Therefore, it could be concluded that the me chanical properties of TCP-alginate based bioprinted constructs should be comparable to those of the cancellous bone.
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Fig. 4. Live (green)/dead (Red) cell assay on 3D bioprinted constructs by FDA/PI staining under fluorescent microscope (left) and cells’ viability quantified by cell counting method (right). The 3D constructs were cultured for 4 h after bioprinting, and then cell viability was analysed. The scale bar was 500 μm * means p < 0.05 and ** means p < 0.005. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
et al., 2011). Moreover, the cell viability continuous reduced in 3 days-culture. They attributed the decrease to the heat effect during PCL printing. Wu et al. found that cell viabilities of fibroblasts and hepatoma cells in 3D bioprinted alginate/cellulose nanocrystals-constructs decreased from 71.00% to 58.91 and from 67.06% to 49.51% after 3 days, respectively (Wu et al., 2018). They considered that the lack of cell-binding sites in the hydrogel was responsible for the declining viability over time. Pati et al. considered that the cushioning effect of dECM gel helped its laden cells to get a high survival (>95%) (Pati et al.,
2014). However, a light decrease (>90%) of cell viability was observed at day 7 and day 14. In another work of Pati et al. they found that the cell viability in the central layer of 3D bioprinted constructs marginally decreased at day 14 (84.43 � 2.16%) compared to day 1 (93.25 � 3.34%), which was attributed to less exchange of nutrient, oxygen and cell information (Pati et al., 2015). Lee et al. also found the same phenomena and their data supported that the shear stress and crosslinking agent (CaCl2) resulted in cell death (Lee et al., 2013). Meanwhile, it was the pore structure but not the thermal shock during 7
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Fig. 5. Cell viability of bioprinted3D constructs after 7 days culture.(A) Relative cells’ viability that was calculated by formula (1); (B) Absolute cells’ viability that was calculated by formula (2).* means p < 0.05 and ** means p < 0.005.
Declaration of competing interest
deposition of the hot PCL ink to affect the cell’s viability. Habib et al. reported that the 3D bioprinted alginate and alginate/carboxymethyl cellulose-constructs at day 15 had approximately 70% and 80% of cell viability, respectively, which was attributed to their porosity difference (Habib et al., 2018). In the present study, cell’s viability at day 7 was significantly decreased by comparison with that at day 0 (Fig. 5). Moreover, the cell viability in TCP: 4% alginate groups without considering TCP content was evidently higher than other groups. Collectively, we considered that cells couldn’t grow and proliferate without enough space if the cell-laden structure was a dense structure. In the following work, we could improve cell survival by change the con centration of bioinks or other matrix, producing suitable porosity and helping cells survival. Except the mechanical improvement by PCL supporter, this study nevertheless demonstrated that this unit-assembly idea would help cell compatible hydrogel materials to be bioprinted into a 3D structure, with the advantages of less damage to cells due to shorter time for printing of units and easy correction on eventual failures in the fabrication or in vitro culture. Although this idea with different junction points and layer interfaces between blocks might imply increased mechanical failure chances, temporary forms of adjunct surgical fixation techniques, as one of the 4Fs (Form, Fixation, Formation, and function) for biomaterials scaffolds, could be applied in clinic, thus also helping solving this problem (Kolambkar et al., 2011).
The authors have no conflicts of interest. Acknowledgment This work was supported by the Shanghai Municipal Science and Technology Commission, China (13JC1403400, 15540723900, 18490740200), and Shanghai Municipal Education Commission, China (Gaofeng Biomedical Engineering Grant, ZXGF082101), GKW project, China (163-15-ZD-09) and National Key Research and Development project, China (SQ2019YFE010621). References Abbadessa, A., Mouser, V.H.M., Blokzijl, M.M., Gawlitta, D., Dhert, W.J.A., Hennink, W. E., Malda, J., Vermonden, T., 2016. A synthetic thermosensitive hydrogel for cartilage bioprinting and its biofunctionalization with polysaccharides. Biomacromolecules 17, 2137–2147. https://doi.org/10.1021/acs.biomac.6b00366. Aljohani, W., Ullah, M.W., Zhang, X., Yang, G., 2018. Bioprinting and its applications in tissue engineering and regenerative medicine. Int. J. Biol. Macromol. 107, 261–275. https://doi.org/10.1016/j.ijbiomac.2017.08.171. Berg, J., Hiller, T., Kissner, M.S., Qazi, T.H., Duda, G.N., Hocke, A.C., Hippenstiel, S., Elomaa, L., Weinhart, M., Fahrenson, C., Kurreck, J., 2018. Optimization of cellladen bioinks for 3D bioprinting and efficient infection with influenza A virus. Sci. Rep. 8, 13877. https://doi.org/10.1038/s41598-018-31880-x. Blaeser, A., Duarte Campos, D.F., Fischer, H., 2017. 3D bioprinting of cell-laden hydrogels for advanced tissue engineering. Curr. Opin. Biomed. Eng. 2, 58–66. https://doi.org/10.1016/j.cobme.2017.04.003. Campos, D.F.D., Philip, M.A., Gurzing, S., Melcher, C., Lin, Y.Y., Schoneberg, J., Blaeser, A., Theek, B., Fischer, H., Betsch, M., 2019. Synchronized dual bioprinting of bioinks and biomaterial inks as a translational strategy for cartilage tissue engineering. 3D Print. Addit. Manuf. 6 (2), 63–71. https://doi.org/10.1089/ 3dp.2018.0123. Carter, D.R., Spengler, D.M., 1978. Mechanical properties and composition of cortical bone. Clin. Orthop. Relat. Res. 135, 192–217. https://doi.org/10.1097/00003086197809000-00041. Chang, R., Nam, J., Sun, W., 2008. Effects of dispensing pressure and nozzle diameter on cell survival from solid freeform fabrication-based direct cell writing. Tissue Eng. A 14, 41–48. https://doi.org/10.1089/ten.a.2007.0004. Compaan, A.M., Christensen, K., Huang, Y., 2016. Inkjet bioprinting of 3D silk fibroin cellular constructs using sacrificial alginate. ACS Biomater. Sci. Eng. 3, 1519–1526. https://doi.org/10.1021/acsbiomaterials.6b00432. Cui, H., Miao, S., Esworthy, T., Zhou, X., Lee, S.J., Liu, C., Yu, Z.X., Fisher, J.P., Mohiuddin, M., Zhang, L.G., 2018. 3D bioprinting for cardiovascular regeneration and pharmacology. Adv. Drug Deliv. Rev. 132, 252–269. https://doi.org/10.1016/j. addr.2018.07.014. Daly, A.C., Cunniffe, G.M., Sathy, B.N., Jeon, O., Alsberg, E., Kelly, D.J., 2016. 3D bioprinting of developmentally inspired templates for whole bone organ engineering. Adv. Healthc. Mater. 5, 2353–2362. https://doi.org/10.1002/ adhm.201600182. Del Rosario, C., Rodriguez-Evora, M., Reyes, R., Delgado, A., Evora, C., 2015. BMP-2, PDGF-BB, and bone marrow mesenchymal cells in a macroporous beta-TCP scaffold
4. Conclusions In summary, we established the interface between 3D bioprinter and micro-CT images and enhanced the mechanical strength of alginate/TCP bioinks based 3D constructs using PCL supporter. This design allowed bioinks with low viscosity to be bioprinted for fabrication of 3D con structs while supplying sufficient mechanical strength for 3D units. The printing pressure was closely associated with viscosity of bioinks and subsequently affected the cells’ survival after printing. When the bioink viscosity was in the range of 3.86–61.65 Pa⋅s, 3D bioprinting process could achieve both good printability and higher cell survival rate. It could be concluded that the combination use of 3D bioprinting tech nique and the supporter was feasible to obtain personalized substitutes. Author disclosure statement No competing financial interests exist.
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Journal of the Mechanical Behavior of Biomedical Materials journal homepage: http://www.elsevier.com/locate/jmbbm
Enhancement of mechanical strength of TCP-alginate based bioprinted constructs Jie-Liang Song 1, Xin-Ye Fu 1, Ali Raza, Nai-An Shen, Ya-Qi Xue, Hua-Jie Wang *, Jin-Ye Wang ** School of Biomedical Engineering, Shanghai Jiao Tong University, 800 Dongchuan Road, Shanghai, 200240, China
A R T I C L E I N F O
A B S T R A C T
Keywords: Bioprinting PCL supporter Unit-assembly model Alginate-TCP bioink Bone defect
To overcome the mechanical drawback of bioink, we proposed a supporter model to enhance the mechanical strength of bioprinted 3D constructs, in which a unit-assembly idea was involved. Based on Computed To mography images of critical-sized rabbit bone defect, the 3D re-construction was accomplished by a sequenced process using Mimics 17.0, BioCAM and BioCAD software. 3D constructs were bioprinted using polycaprolactone (PCL) ink for the outer supporter under extrusion mode, and cell-laden tricalcium phosphate (TCP)/alginate bioink for the inner filler under air pressure dispensing mode. The relationship of viscosity of bioinks, 3D bio printing pressure, TCP/alginate ratio and cell survival were investigated by the shear viscosities analysis, live/ dead cell test and cell-counting kit 8 measurement. The viscosity of bioinks at 1.0 s 1-shear rate could be adjusted within the range of 1.75 � 0.29 Pa⋅s to 155.65 � 10.86 Pa⋅s by changing alginate concentration, cor responding to 10 kPa–130 kPa of printing pressure. This design with PCL supporter could significantly enhance the compressive strength and compressive modulus of standardized 3D mechanical testing specimens up to 2.15 � 0.14 MPa to 2.58 � 0.09 MPa, and 42.83 � 4.75 MPa to 53.12 � 1.19 MPa, respectively. Cells could maintain the high viability (over 80%) under the given printing pressure but cell viability declined with the increase of TCP content. Cell survival after experiencing 7 days of cell culture could be achieved when the ratio of TCP/alginate was 1 : 4. All data supported the feasibility of the supporter and unit-assembly model to enhance mechanical properties of bioprinted 3D constructs.
1. Introduction
2016; Blaeser et al., 2017; Duarte Campos et al., 2016; Lee et al., 2014). Bioink is one of the key restrictions on the bioprinting technique. The requirement for the balance between rheological property and shape retention prevents the application of a number of native biomaterials as bioinks, especially for hard tissue implant fabrication, although they satisfy the requirements of biocompatibility and biodegradability (Compaan et al., 2016; Grande, 2017; Jang et al., 2017; Jia et al., 2014; Law et al., 2018; Pati et al., 2014; Rhee et al., 2016; Zehnder et al., 2015; Kim et al., 2018). In another word, bioprinting is a layer-by-layer printing technique and the lower layer needs to support the upper layer (Aljohani et al., 2018). Otherwise the structure will collapse. Especially, in view of the complex mechanical requirement of tissues and organs, the bioprinted 3D constructs with suitable mechanical properties are still a challenge (Carter and Spengler, 1978; Vanderhooft et al., 2009). The 3D structure design is a possible breakthrough point for
The bioprinting technique benefits the precise control over multiple compositions, spatial distributions, and architectural accuracy/ complexity, by layer-by-layer printing of cells-laden biomaterial (bio inks) (Blaeser et al., 2017; Moroni et al., 2018; Murphy and Atala, 2014). From this point of view, it is likely to achieve the highly biomimetic objective at structural features, mechanical properties and even function of native tissues (Cui et al., 2018). Although 3D printing has a predicted $ 23.0 billion market in 2022, the bioprinting technique is now in the ascendant, characterizing with more “comfortable” printing conditions for cells’ survival, such as mild printing temperature and pressure (Chang et al., 2008; Derakhshanfar et al., 2018; Zhao et al., 2015). Various engineered tissues or organs have been obtained by the bio printing technique, such as cartilage, skin tissue, bone tissue, muscle, neural, osteochondral and pancreatic tissue and renal (Abbadessa et al.,
* Corresponding author. ** Corresponding author. E-mail addresses: [email protected] (H.-J. Wang), [email protected] (J.-Y. Wang). 1 These authors contributed equally to this work. https://doi.org/10.1016/j.jmbbm.2019.103533 Received 19 March 2019; Received in revised form 7 November 2019; Accepted 12 November 2019 Available online 13 November 2019 1751-6161/© 2019 Elsevier Ltd. All rights reserved.
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mechanical improvement of bioprinted 3D constructs (Daly et al., 2016; Hu et al., 2018). The hybrid scaffold consisting of a synthetic polymer as supporter with cell-laden hydrogels is a typical and successful design (Lee et al., 2013). For example, Schuurman et al. established a state-of-the-art printing technology to obtain a polycaprolactone (PCL)/alginate hybrid tissue construct with over 6 MPa of Young’s modulus, which matched the range with cartilage 4.1 MPa and trachea 3.33 MPa (Schuurman et al., 2011). In this design, alginate/cells bioinks were filled into the grooves formed by the PCL strands and PCL played a key important role in supplying a mechanical support to the 3D con structs, while alginate worked as the matrix to localize cells. Instead of alginate, Pati et al. used decellularized adipose tissue matrix/mesen chymal stem cells bioink to print a dome-shaped structure with PCL, which had 122.56 � 20.23 kPa of compressive modulus (Pati et al., 2015). In addition, Campos et al. recently reported a model composed of hollow cylinders with three supporting pins and their negatives (Campos et al., 2019). The mixture of agarose and alginate was chosen as the supporting material, and the mixture of agarose and type I collagen was chosen as bioink. The highest compressive modulus was over 600 kPa. Although efforts in 3D bioprinting field have been made to expand the applicability of cell-laden hydrogels from soft tissue regeneration to the more demanding field of hard tissue regeneration, there are few reports focusing on bone tissue engineering due to their relatively weak me chanical properties. In this work, we proposed a supporter model to enhance the me chanical strength of bioprinted 3D constructs, in which a unit-assembly idea was involved (Scheme 1). Here, the unit was composed of the PCL supporter and cell-laden tricalcium phosphate (TCP)-alginate filler. The former formed the up-down pipelines, columnar linkers, sides and bot tom of bioprinted 3D construct. The latter acted as bioinks to fill in the
PCL frame. After that, the units were assembled together to form the whole 3D constructs by the columnar linkers. The unit-assembly idea had the advantages of less damage to cells due to shorter time for printing of units and easy correction on eventual failures in the fabri cation or in vitro culture. 2. Materials and methods 2.1. Materials PCL (Mw ¼ 80,000 Da) was purchased from Sigma Aldrich (Natick, MA, USA) and used for holt-melt extrusion-based printing material without further purification. Alginate (medium viscosity, 200–500 mPa⋅s) was purchased from Adamas-beta Co. Ltd. (Shanghai, China) and TCP was purchased from Shanghai No.4 Reagent & H.V. Chemical Co. Ltd. (Shanghai, China). Fluorescein diacetate (FDA) and propidium iodide (PI) were purchased from Yeasen Biotech Co. Ltd. (Shanghai, China). Cell-counting kit 8 (CCK8) was purchased from Beyotime Biotechnology (Shanghai, China). 2.2. Cells The human hepatocarcinoma (BEL-7402) cells were obtained from Shanghai Institute of Biochemistry and Cell Biology (Shanghai, China). Cells were routinely cultured in RPMI1640 medium (containing 10% of new borne bovine serum, 100 units/mL penicillin and 100 μg/mL streptomycin). After 80% of confluency, cells were released from the bottom using 0.25% trypsin solution and collected by centrifugation at 1000 rpm for 5 min. BEL-7402 cells were dispersed into fresh medium to get 1 � 106 cells/mL density for bioinks preparation.
Scheme 1. Illustration of the design of 3D constructs and 3D bioprinting process. 2
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2.3. Animals New Zealand rabbits (3.0–4.0 kg) were purchased from Songlian Experimental Animal Farm (Shanghai, China). All animal procedures were approved by the Institutional Animal Care Committee. All guide lines met ethical standards required by law and also complied with the guidelines for the use of experimental animals in China. The 20-mm length of critical-size radial bone defect model in rabbit was made under anaesthesia via intramuscular administration with 3% pentobarbital sodium (30 mg/kg) according to Tu et al. (2009). The actual digital information of radial bone defect was obtained by micro-CT (SkyScan1176, Bruker, Germany) with 90 kV of voltage, 278 μA of current, 17.93 μm of isotropic voxel size and 1336 � 2000 pixel of field of view. All images were archived as Tiff files.
was magnified three times in order to test the unit-assembly idea and possibility to scale up for use in human-scale bone defect. As shown in Scheme 1, BioCAD software was used to process the unit-assembly model. Here, the re-constructed 3D structure contained 15 units, which had 4 mm height. A bottom, sides, up-down channels and columnar linkers of the units formed the supporter structure, which was printed using PCL ink. The thickness of the supporter layer on bottom and sides was 2 mm and 0.33 mm, respectively. The inner diameter and height of channels was 2.0 mm and 4.0 mm, respectively. The height of columnar linkers was 4.7 mm, including an upper ridge (2.0 mm in diameter and 1.2 mm in height) and a round groove (2.0 mm in diameter and 1.5 mm in height) at the bottom of the linker. In addition, the inner filler structure was bioprinted using cell-laden TCP/alginate bioinks. The units were linked with each other through columnar linkers.
2.4. Preparation and rheological analysis of bioinks
2.6. 3D bioprinting
Alginate and TCP powder (size < 100 μm) were sterilized by ultra violet (UV) irradiation for 30 min and mixed with the abovementioned cell suspension. The final concentration of alginate was 2%, 4% and 6% (w/v), respectively. The TCP/alginate ratio was listed in Table 1. The rheological properties of bioinks were analysed using a rota tional rheometer (MCR302, Anton Paar, Graz, Austria) (n ¼ 3). The steady-state shear viscosities of samples were conducted over a range of shear rate 1–1000 s 1 and all measurements were conducted at room temperature.
3D bioprinting process was performed using 3D bioprinter (3D Dis covery, RegenHU, Switzerland) under aseptic conditions. Here, the revolution of 3D bioprinter was 50 μm. Firstly, the melted PCL was extruded at 90 � C via a HM-300H print-head with 0.33 mm diameter and 20 r/s extruder speed. The layer height was 0.25 mm. The pressure and moving speed of nozzle were 0.3 MPa and 5 mm/s, respectively. Sec ondly, CF300H print-head without the valve was used for bioinks printing under continuous air pressure dispensing mode and the pres sure was adjusted in the range of 10–130 kPa according to the rheo logical properties of bioinks. Here, the valve of CF300H print-head was removed in order to match the cell-alginate/TCP bioinks and 19 G nozzle (0.9 mm-size of diameter) was used for CF300H print-head. The moving speed and temperature of nozzle were 5 mm/s and 37 � C, respectively. Thirdly, the printed units were immersed into 0.1 M CaCl2 solution for 20 min, and then rinsed with PBS three times. The units were further cultured in a 12-well plate in a 5% CO2 incubator at 37 � C. Finally, the units were assembled together one by one to form the whole defected bone as shown in the Fig. 1H.
2.5. Computer-aided 3D reconstruction of bone defect Mimics 17.0 software (Materialise, Belgium) was used to process micro-CT images for re-construction of bone defect, including image reading, filtering, threshold segmentation and 3D model reconstruction. The image processing was in detailed described in re sults and discussion section. The files were saved as STL format. BioCAM software (RegenHU, Switzerland) was used to read 3D reconstructed STL format file and the digital information of bone defect
2.7. Compressive properties of 3D constructs
Table 1 Viscosities at 1.0 s 1-shear rate and 3D bioprinting pressures for various bioink formulations (n ¼ 3). Groups
Alginate (% w/v)
TCP: Alginate (w: w)
Viscosity at 1.0 s 1shear rate (Pa⋅s)
Printing pressure (kPa)
TCP: 2% ALG (0 : 4) TCP: 2% ALG (1 : 4) TCP: 2% ALG (8 : 4) TCP: 4% ALG (0 : 4) TCP: 4% ALG (1 : 4) TCP: 4% ALG (8 : 4) TCP: 6% ALG (0 : 4) TCP: 6% ALG (1 : 4) TCP: 6% ALG (8 : 4)
2
0:4
1.75 � 0.29
10
2
1:4
2.44 � 0.45
10
2
8:4
2.19 � 0.17
10
4
0
39.19 � 4.89
45
4
1:4
33.28 � 4.26
45
4
8:4
33.80 � 4.05
50
6
0:4
123.04 � 37.37
120
6
1:4
115.82 � 3.39
120
6
8:4
155.65 � 10.86
130
The cylinder samples with the standard size (d ¼ 10 mm, h ¼ 10 mm) matching the requirement of the ISO 13314 : 2011 had been prepared using the abovementioned method. Briefly, the designed cylinders were fabricated by printing the supporter PCL perimeter and the different alginate/TCP inks within. The mechanical tests were carried out on a universal testing machine (Z020, ZwickRoell, Germany) at a rate of 1.0 mm/min (n ¼ 3). The resulting properties were calculated according to the following formula:
σ ¼ Eε where σ is the stress, ε is the strain, E is the compressive modulus which was calculated using slope of the linear region of the stress strain curve. 2.8. Cell survival and viability After experiencing a 4 h of cell culture, the absolute viability of the bioprinted cells in the standardized 3D units (14 mm � 14 mm � 4 mm) was tested using the Live/Dead assay kit according to the protocol (n ¼ 3) (Yeasen Biotech Co. Ltd., Shanghai, China). Briefly, 1 ml of FDA/ PI staining solution was added into each well of 12-well plate and incubated for 5 min at room temperature. After rinsing three times with PBS (pH 7.2–7.4, 0.1 M), the cells were counted under the fluorescent microscope (IX-70, Olympus, Japan). The live cells were green (FDA) and the dead cells were red (PI). The absolute cells’ viability0days was expressed as the percentage of the live cells number/total cells number. The relative viability of the bioprinted cells in the standardized 3D units after cell culture for 4 h and 7 days was measured using CCK-8 3
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Fig. 1. 3D re-construction of bone defect of rabbit by a sequenced process using Mimics, BioCAM and BioCAD software and the bioprinted 3D constructs. (A) Bone defect model in rabbit; (B) Micro-CT images of bone defect (1) and typical micro-CT cross section images (2); (C) 3D re-construction of bone defect by Mimics 17.0 software through a four step process, including acquisition, filtering, segmentation and charac terization process. Front view (1), Top view (2), Left view (3), Isometric projection (4); (5–8) threshold segmentation-dynamic regional growth, twenty-six connected body; the magnetic lasso method to re-construct defected section (5), top view (6), front view (7), left view (8); (D) Reading of the 3D digital information by BioCAM software. (E) 3D processing by BioCAD software. (F1) The bioprinted 3D constructs. (F2) Unit of 3D constructs.
detection kit. Briefly, the fresh medium containing 10% of CCK-8 solu tion was added to each well of 12-well plate and the cells were incubated at 37 � C for another 4 h. The OD values at 450 nm were recorded using a microplate reader (Multiskan FC, Thermo, USA). The relative and ab solute cells’ viability was calculated using the following formulas: Relative cells’ viability ¼ OD7days/OD0days � 100%
3. Results and discussion 3.1. Interface establishment of 3D bioprinter and micro-CT image In this work, the sequenced processing steps for re-construction of bone defect model were performed using Mimics 17.0, BioCAM and BioCAD software (Fig. 1) (McCormick et al., 2014; Sun et al., 2004). Firstly, the four-step process for the transformation from the micro-CT image (Tiff format) to 3D model file (STL format) was accomplished using Mimics 17.0, including image acquisition, filtering, segmentation and characterization process (Fig. 1C). Here, the median filter produces the best noise reduction effect without damaging the edge of bone tissue compared to other modes. After that, image segmentation is considered to be a crucial step to identify the region of interest, which can be per formed by thresholding, edge, region or cluster-based methods. Dy namic threshold segmentation and gray value segmentation were used. Both methods could easily and accurately separate bone tissue from other parts of the image. However, it was much difficult to generate the mask of bone defect because there were little differences between the gray value of the bone defect area and the surroundings, and the boundary of bone defect area was not clear. Therefore, the live wire algorithm is used to draw the outline of the target area manually, and then edited the generated masks in each layer as shown in Fig. 1C5 - C8. The digital 3D model of bone defect is re-constructed and identified by
(1)
Absolute cells’ viability7days ¼ Relative cells’ viability � Absolute cells’ via bility0days � 100% (2)
2.9. Statistical analysis All data were reported as the mean � standard deviation. The data were analysed by one-way factorial analysis of variances (one-way ANVOA) and multiple comparisons. Significant effects were defined using Fisher’s method as a post-hoc test, with p < 0.05 considered as being statistically significant.
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Table 1 lists the corresponding viscosity at 1.0 s 1-shear rate and 3D bioprinting pressure. Generally speaking, the viscosities of as-prepared samples highly rely on the concentration of alginate. The curves for bioinks containing 2% alginate appear to be relatively flat, suggesting minimal shear-thinning property. When alginate concentration is over 4%, the curves become steep, suggesting that a higher pressure will be needed for printing. In the work reported by Wu et al. the viscosity of pure alginate solution was also verified to be positively relative to the concentration of alginate and the viscosity at 1.0 s 1-shear rate varied in the range of 0–10 Pa⋅s when the concentration of alginate increased from 2% to 6% (Wu et al., 2018). Here, the viscosities at 1.0 s 1-shear rate of bioinks are in the range of 1.75 � 0.29–155.65 � 10.86 Pa⋅s and increase with the increase of alginate concentration. The addition of TCP doesn’t significantly affect shear-thinning property of bioinks. Freeman and Kelly demonstrated that molecular weight and ionic crosslinker were two important factors affecting the viscosity and printability of alginate bioinks (Freeman and Kelly, 2017). Our alginate belonged to medium viscosity product, while the alginate used by Wu et al. was pharmaceutical grade and belonged to low viscosity product (Wu et al., 2018). Therefore, it was the main reason to explain the difference of rheological properties of bioinks between two groups. The original CF300H printhead with a valve (3D Discovery, RegenHU) is very suitable to print inks with lower viscosity, especially cell solution. As for our bioinks, the relatively high viscosity and exis tence of TCP powder very easily led to block of the printhead. Especially for the printing of big-size samples, it was very difficult to continuously print a complete 3D structure. To solve this problem, we removed the valve and used a 10 cc cartridge, in combination with a suitable nozzle (19 G), to print our bioinks. Meanwhile, because the relatively high viscosity, the bioinks couldn’t flow freely without air pressure. It was noted that we could ignore the free-flow problem of bioinks when we controlled the moving speed of nozzle at more than 5 mm/s. The higher viscosity was associated with the higher printing pressure. As shown in Table 1, the printing pressure for 3D bioprinting of bioinks ranges from 10 kPa to 130 kPa. The bioinks containing 2% alginate can be easily pushed out, but can’t keep shape without the supporter due to high fluidity (Fig. 3A1-A3). PCL supporter and the post-crosslinking process using CaCl2 enable these bioinks constructing the 3D units without collapse, but all designed PCL channels become closed (Fig. 3a1-a3). On the contrary, the bioinks containing 6% alginate have the lower fluidity and can keep shape well (Fig. 3C1-C3). However, the printed bioinks shrink after crosslinking by immersion in 0.1 M CaCl2 solution for 20 min, and could not completely attach to the PCL supporter (Fig. 3c1c3). While 4% of alginate concentration shows a perfect 3D structure (Fig. 3B and b). The compressive strength and compressive modulus of bioprinted 3D samples vary from 2.15 � 0.14 MPa to 2.58 � 0.09 MPa, and 42.83 � 4.75 MPa to 53.12 � 1.19 MPa, respectively. All values satisfy the mechanical demand of cancellous bone (strength: 1.5–38 MPa and modulus: 10–1570 MPa) as shown in Table 2 (Yusop, 2012). 3D struc ture design play the key role to improve mechanical properties of printed 3D constructs. For example, Schuurman et al. reported the typical groove (synthesis polymer)/filler (cell-laden bioink) and layer-by-layer structure that had over 6 MPa of Young’s modulus (Schuurman et al., 2011). Xu et al. accomplished the mechanical improvement of 3D printed constructs with 1.76 MPa of Young’s modulus, which was four times higher than that of alginate constructs (0.41 MPa) (Xu et al., 2013). Their structure was characterized with an alternative layers structure of PCL and cell-laden hydrogel. In this study, Young’s modulus of obtained 3D mechanical testing specimens was evidently improved by comparison with pure alginate (less than 0.7 MPa) (Ren et al., 2018; Roushangar et al., 2018), and there were no significant differences no matter if there was TCP, indicating that PCL was responsible for the mechanical strength of 3D constructs. Actually it is still a challenge for mechanical test of irregular samples in the related fields. In general, the precise testing depends on testing method, and the
BioCAM (Fig. 1D). Upon early diagnosis, the exact anatomical model of target can be acquired through CT or magnetic resonance imaging (MRI) imaging modalities without surgical intervention and with minimum radiation exposure (Fu et al., 2017). Based on CT or MRI images, the design of tissue constructs before bioprinting will play an extremely crucial role in determining the properties of bioprinted constructs, which must conform to injury-specific geometrical dimensions (Fig. 1B and C) (Nam et al., 2015; Teodori et al., 2017). Besides Mimics, there are other design software for 3D modelling, such as TSIM, Solidworks, 3D Slicer, MATLAB and OsiriX. 3.2. Design of 3D constructs Scheme 1 shows the design of 3D constructs by BioCAD software (Fig. 1D) and printing process. This study not only supplies a supporter model to enhance the mechanical strength of 3D-bioprinted structure, but also a unit-assembly idea is involved. Most hydrogels fail to establish 3D structures with big size or good mechanical properties because of their inherently incongruous property between rheology and shape retention. The supporter design is an alternative to improve this status, such as the work of Daly et al. Lee et al., Schuurman et al. and pati et al. (Daly et al., 2016; Lee et al., 2013; Pati et al., 2015; Schuurman et al., 2011). The hybrid scaffold consisting of a synthetic polymer as sup porter with cell-laden hydrogels was the basis of their 3D designs. The Young’s modulus even could be enhanced up to megapascales (MPa) level (Schuurman et al., 2011). In this study, we combined the supporter structure with unit-assembly structure. The designed 3D structure is divided into 15 sections that were named as units with the columnar linker and up-down channels. The upper ridge of columnar linker could match with the bottom groove of adjacent columnar liner, enabling the proper fixing of neighbouring units. The up-down channels allowed solution to pass through. There were two advantages for this design. Firstly, the units could be bioprinted in a relative shorter time by comparison with the whole 3D structure. This would support a higher survival rate to cells due to the adverse effects from environmental factors during the printing process. Secondly, one unit could be bioprinted into several identical backups and easily replaced if some of them met with damage or infection before use. 3.3. Preparation of bioinks and printing of 3D constructs We studied the rheological properties of as-prepared bioinks. Fig. 2 shows the viscosity vs. shear rate curves for all bioink formulations, and
Fig. 2. The viscosity vs. shear rate curves for all bioink formulations (n ¼ 3). 5
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Journal of the Mechanical Behavior of Biomedical Materials 103 (2020) 103533
3.4. Cells’ viability The standardized units were used to compare cell viabilities after 3D printing and 7 days cell culture (Fig. 4 and Fig. 5). The property of hydrogels themselves, and bioprinting parameters, such as nozzle tem perature, printing time, nozzle diameter, and dispensing pressure would affect the viability of cells in 3D bioprinted constructs (Chang et al., 2008; Zhao et al., 2015). In the view of biocompatibility of alginate, many literatures have demonstrated this point (Morimoto et al., 2017). The presence of alginate will help the precision location of cells during 3D bioprinting process. In addition, our data also supported that algi nate was able to help shield cells from shear stress during the extrusion process (Moroni et al., 2018). BEL-7402 cells in alginate group (TCP: alginate ¼ 0 : 4) had over 85% of cell viability after printing and it got to over 95% in TCP: 2% alginate ¼ 0 : 4 group. Cell death is a key challenge for 3D bioprinting technique. The printing pressure is the key factor that firstly should be considered. In the preliminary work, we tried to print cell suspension without any matrix materials under different printing pressures. The results revealed that the cells viability decreased greatly when the printing pressure was over 120 Pa. Nair et al. used 1.5% w/v alginate solution to study the effect of printing pressure on cells’ survival and gave a mathematical model (Nair et al., 2009). They found that live cells, injury cells and dead cells got to over about 85%, 3% and 7.5%, respectively, using 34.5 kPa pressure and 400 μm diameter of nozzle. However, when the printing pressure was set as 275.8 kPa, live cells, injury cells and dead cells only got to about 60%, 5% and 35%, respectively. As for TCP: alginate bio inks, the suitable rheological properties enabled them to be printed under less than 120 Pa and our data suggested that the effect of printing pressure within the range of 10–120 Pa on cell viability could be ignored. In the preliminary study, we also tried to improve the printing ac curacy and found the bioink was the key factor. It greatly affected the printing pressure (regardless of print modes) and the accuracy (including the accuracy in X, Y and Z). The former decides cells’ viability that has been proved in several groups, including ours. The latter in volves the printability and shape retention. At the same conditions, the increase of accuracy means the decrease of the used nozzle size and the subsequently increase of printing pressure. In the present study, we could get over 90% of cells’ survival rate at the given conditions (with less than 120 Pa of printing pressure). When we tried to change a smaller size of nozzle, the improvement of accuracy was at the expense of the cells’ survival rate under air pressure dispensing mode, even an extru sion mode was applied. Additionally, it was not neglectable that the cell viability decreased with the increase of TCP (Fig. 4). TCP, as another component of bioinks, is one of the most widely used biomaterial in bone tissue engineering. TCP has demonstrated osteogenic properties, phase stability and strong bond formation with the host bone tissue in different studies (Del Rosario et al., 2015; Fahimipour et al., 2017; Mojahedian et al., 2016; Park et al., 2016; Yang et al., 2015). With the increase of TCP concen tration, the decrease in cells’ viability was clearly observed (Fig. 4). In 2% alginate-containing groups, cell viability declined from 97.28% � 3.00% (TCP: 2% ALG (0: 4)) to 81.33% � 13.18% (TCP: 2% ALG (8: 4)), while cell viability declined from 88.49% � 12.47% (TCP: 6% ALG (0 : 4)) to 1.32% � 2.31% (TCP: 6% ALG (8 : 4)) in 6% alginate-containing groups. Jiang et al. reported that the high shear stress was the main reason to induce cell membrane damage and decrease cell viability during 3D bioprinting process (Jiang et al., 2017). The existence of hard TCP powder might enhance the shear stress of bioinks. The factors from inks and 3D structure on cell viability have been demonstrated by several groups (Berg et al., 2018; Habib et al., 2018; Wu et al., 2016; Wu et al., 2018). The work of Schuurman et al. showed that the cells viability in 3D constructs made of PCL/C20A4 cells-laden alginate bioinks was over 80% after bioprinting process (Schuurman
Fig. 3. Optical images of different 3D units with (Lowercase letters) or without (Capital letters) PCL supporter. Table 2 Compressive strength and modulus of bioprinted 3D constructs (n ¼ 3). TCP: 2% ALG (0 : 4) TCP: 2% ALG (1 : 4) TCP: 2% ALG (8 : 4) TCP: 4% ALG (0 : 4) TCP: 4% ALG (1 : 4) TCP: 4% ALG (8 : 4) TCP: 6% ALG (0 : 4) TCP: 6% ALG (1 : 4) TCP: 6% ALG (8 : 4)
Compressive strength (MPa)
Compressive modulus (MPa)
2.42 � 0.08 2.52 � 0.11 2.15 � 0.14 2.37 � 0.09 2.40 � 0.05 2.32 � 0.09 2.58 � 0.09 2.48 � 0.18 2.45 � 0.02
48.50 � 7.55 52.64 � 6.64 42.90 � 6.19 44.27 � 2.18 46.73 � 8.36 42.83 � 4.57 53.12 � 1.19 47.56 � 8.31 49.92 � 2.19
production, size and shape of the test specimen. Therefore, the stan dardized regular shape for mechanical tests of specimen is required. In the present study, the method supplied by ISO 13314:2011 had been utilized for preparation and compressive test of our units. The cylinder samples had the standard size (d ¼ 10 mm, h ¼ 10 mm) and were pre pared using bioprinting method. We can expect that the basic mechan ical properties obtained from the standard smaller units experiments will help to determine the strength and other coefficients of the big irregular structure while application, by simulation such as finite element methods. This is the normal and similar way often used in En gineering Mechanics. Therefore, it could be concluded that the me chanical properties of TCP-alginate based bioprinted constructs should be comparable to those of the cancellous bone.
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Fig. 4. Live (green)/dead (Red) cell assay on 3D bioprinted constructs by FDA/PI staining under fluorescent microscope (left) and cells’ viability quantified by cell counting method (right). The 3D constructs were cultured for 4 h after bioprinting, and then cell viability was analysed. The scale bar was 500 μm * means p < 0.05 and ** means p < 0.005. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)
et al., 2011). Moreover, the cell viability continuous reduced in 3 days-culture. They attributed the decrease to the heat effect during PCL printing. Wu et al. found that cell viabilities of fibroblasts and hepatoma cells in 3D bioprinted alginate/cellulose nanocrystals-constructs decreased from 71.00% to 58.91 and from 67.06% to 49.51% after 3 days, respectively (Wu et al., 2018). They considered that the lack of cell-binding sites in the hydrogel was responsible for the declining viability over time. Pati et al. considered that the cushioning effect of dECM gel helped its laden cells to get a high survival (>95%) (Pati et al.,
2014). However, a light decrease (>90%) of cell viability was observed at day 7 and day 14. In another work of Pati et al. they found that the cell viability in the central layer of 3D bioprinted constructs marginally decreased at day 14 (84.43 � 2.16%) compared to day 1 (93.25 � 3.34%), which was attributed to less exchange of nutrient, oxygen and cell information (Pati et al., 2015). Lee et al. also found the same phenomena and their data supported that the shear stress and crosslinking agent (CaCl2) resulted in cell death (Lee et al., 2013). Meanwhile, it was the pore structure but not the thermal shock during 7
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Fig. 5. Cell viability of bioprinted3D constructs after 7 days culture.(A) Relative cells’ viability that was calculated by formula (1); (B) Absolute cells’ viability that was calculated by formula (2).* means p < 0.05 and ** means p < 0.005.
Declaration of competing interest
deposition of the hot PCL ink to affect the cell’s viability. Habib et al. reported that the 3D bioprinted alginate and alginate/carboxymethyl cellulose-constructs at day 15 had approximately 70% and 80% of cell viability, respectively, which was attributed to their porosity difference (Habib et al., 2018). In the present study, cell’s viability at day 7 was significantly decreased by comparison with that at day 0 (Fig. 5). Moreover, the cell viability in TCP: 4% alginate groups without considering TCP content was evidently higher than other groups. Collectively, we considered that cells couldn’t grow and proliferate without enough space if the cell-laden structure was a dense structure. In the following work, we could improve cell survival by change the con centration of bioinks or other matrix, producing suitable porosity and helping cells survival. Except the mechanical improvement by PCL supporter, this study nevertheless demonstrated that this unit-assembly idea would help cell compatible hydrogel materials to be bioprinted into a 3D structure, with the advantages of less damage to cells due to shorter time for printing of units and easy correction on eventual failures in the fabrication or in vitro culture. Although this idea with different junction points and layer interfaces between blocks might imply increased mechanical failure chances, temporary forms of adjunct surgical fixation techniques, as one of the 4Fs (Form, Fixation, Formation, and function) for biomaterials scaffolds, could be applied in clinic, thus also helping solving this problem (Kolambkar et al., 2011).
The authors have no conflicts of interest. Acknowledgment This work was supported by the Shanghai Municipal Science and Technology Commission, China (13JC1403400, 15540723900, 18490740200), and Shanghai Municipal Education Commission, China (Gaofeng Biomedical Engineering Grant, ZXGF082101), GKW project, China (163-15-ZD-09) and National Key Research and Development project, China (SQ2019YFE010621). References Abbadessa, A., Mouser, V.H.M., Blokzijl, M.M., Gawlitta, D., Dhert, W.J.A., Hennink, W. E., Malda, J., Vermonden, T., 2016. A synthetic thermosensitive hydrogel for cartilage bioprinting and its biofunctionalization with polysaccharides. Biomacromolecules 17, 2137–2147. https://doi.org/10.1021/acs.biomac.6b00366. Aljohani, W., Ullah, M.W., Zhang, X., Yang, G., 2018. Bioprinting and its applications in tissue engineering and regenerative medicine. Int. J. Biol. Macromol. 107, 261–275. https://doi.org/10.1016/j.ijbiomac.2017.08.171. Berg, J., Hiller, T., Kissner, M.S., Qazi, T.H., Duda, G.N., Hocke, A.C., Hippenstiel, S., Elomaa, L., Weinhart, M., Fahrenson, C., Kurreck, J., 2018. Optimization of cellladen bioinks for 3D bioprinting and efficient infection with influenza A virus. Sci. Rep. 8, 13877. https://doi.org/10.1038/s41598-018-31880-x. Blaeser, A., Duarte Campos, D.F., Fischer, H., 2017. 3D bioprinting of cell-laden hydrogels for advanced tissue engineering. Curr. Opin. Biomed. Eng. 2, 58–66. https://doi.org/10.1016/j.cobme.2017.04.003. Campos, D.F.D., Philip, M.A., Gurzing, S., Melcher, C., Lin, Y.Y., Schoneberg, J., Blaeser, A., Theek, B., Fischer, H., Betsch, M., 2019. Synchronized dual bioprinting of bioinks and biomaterial inks as a translational strategy for cartilage tissue engineering. 3D Print. Addit. Manuf. 6 (2), 63–71. https://doi.org/10.1089/ 3dp.2018.0123. Carter, D.R., Spengler, D.M., 1978. Mechanical properties and composition of cortical bone. Clin. Orthop. Relat. Res. 135, 192–217. https://doi.org/10.1097/00003086197809000-00041. Chang, R., Nam, J., Sun, W., 2008. Effects of dispensing pressure and nozzle diameter on cell survival from solid freeform fabrication-based direct cell writing. Tissue Eng. A 14, 41–48. https://doi.org/10.1089/ten.a.2007.0004. Compaan, A.M., Christensen, K., Huang, Y., 2016. Inkjet bioprinting of 3D silk fibroin cellular constructs using sacrificial alginate. ACS Biomater. Sci. Eng. 3, 1519–1526. https://doi.org/10.1021/acsbiomaterials.6b00432. Cui, H., Miao, S., Esworthy, T., Zhou, X., Lee, S.J., Liu, C., Yu, Z.X., Fisher, J.P., Mohiuddin, M., Zhang, L.G., 2018. 3D bioprinting for cardiovascular regeneration and pharmacology. Adv. Drug Deliv. Rev. 132, 252–269. https://doi.org/10.1016/j. addr.2018.07.014. Daly, A.C., Cunniffe, G.M., Sathy, B.N., Jeon, O., Alsberg, E., Kelly, D.J., 2016. 3D bioprinting of developmentally inspired templates for whole bone organ engineering. Adv. Healthc. Mater. 5, 2353–2362. https://doi.org/10.1002/ adhm.201600182. Del Rosario, C., Rodriguez-Evora, M., Reyes, R., Delgado, A., Evora, C., 2015. BMP-2, PDGF-BB, and bone marrow mesenchymal cells in a macroporous beta-TCP scaffold
4. Conclusions In summary, we established the interface between 3D bioprinter and micro-CT images and enhanced the mechanical strength of alginate/TCP bioinks based 3D constructs using PCL supporter. This design allowed bioinks with low viscosity to be bioprinted for fabrication of 3D con structs while supplying sufficient mechanical strength for 3D units. The printing pressure was closely associated with viscosity of bioinks and subsequently affected the cells’ survival after printing. When the bioink viscosity was in the range of 3.86–61.65 Pa⋅s, 3D bioprinting process could achieve both good printability and higher cell survival rate. It could be concluded that the combination use of 3D bioprinting tech nique and the supporter was feasible to obtain personalized substitutes. Author disclosure statement No competing financial interests exist.
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